Interatrial shunt with enlarged pars constrictor area

Abstract

The device comprises a bridge made of a biocompatible material that surrounds the necked-down region of the encapsulated shunt. In one embodiment, the bridge is configured to maintain an outer diameter while the diameter of the necked-down region is reduced or increased in vivo to regulate fluid flow rate through the device. The bridge may have an outer diameter configured to be placed within an enlarged septal hole. Methods for regulating fluid flow and for the manufacture of such devices are also provided.

Description

[Technical field]

[0001] (CROSS REFERENCE TO RELATED APPLICATIONS) This application claims priority to U.S. Provisional Patent Application No. 63 / 386,147, filed December 5, 2022, and U.S. Provisional Patent Application No. 63 / 363,015, filed April 14, 2022, the entire contents of each of which are incorporated herein by reference.

[0002] The present technology relates to devices for use in the human body, such as percutaneously implanted devices, including devices for regulating pressure within the circulatory system, such as interatrial shunts for regulating blood pressure in the heart, and methods for the manufacture of such devices. [Background technology]

[0003] For some medical conditions, there is benefit in regulating the flow of fluids in the human body, for example, through a passage between two body cavities. Such passages are typically used in catheter placement procedures, where a catheter is delivered through the patient's vascular system. In some catheter placement procedures, there is benefit in moving from one cavity to another by creating a passage. For example, such a passage can be created between the right side of the heart and the left side of the heart, for example, from the right atrium to the left atrium, and a clinical procedure is performed on the left side of the heart using an approach from the right side of the heart. Such clinical procedures include, for example, arrhythmia ablation procedures in the left atrium and mitral valve repair activities.

[0004] In addition, passages can be created and maintained in the heart wall between the two ventricles to accommodate shunts for redistributing blood from one ventricle to another to address pathologies such as heart failure (HF), myocardial infarction (MI), and pulmonary arterial hypertension (PAH). HF is a physiological state in which cardiac output is insufficient to meet the body's needs or does so only at higher filling pressures. There are many underlying causes of HF, including MI, coronary artery disease, valvular disease, hypertension (such as PAH), and myocarditis. Chronic heart failure is associated with alterations in neurohormonal activation and autonomic control. Although these compensatory neurohormonal mechanisms provide beneficial support for the heart under normal physiological circumstances, they also play a fundamental role in the development and subsequent progression of HF.

[0005] HF is generally classified as either systolic heart failure ("SHF") or diastolic heart failure ("DHF"). In SHF, the pumping action of the heart is reduced or weakened. A common clinical measurement is the ejection fraction, which is a function of blood ejected out of the left ventricle (stroke volume) divided by the maximum volume in the left ventricle at the end of diastole or relaxation. A normal ejection fraction is greater than 50%. Systolic heart failure generally causes a reduced ejection fraction of less than 40%. Such patients have heart failure with reduced ejection fraction ("HFrEF"). Patients with HFrEF may usually have a larger left ventricle due to a phenomenon called "cardiac remodeling" that occurs secondary to higher ventricular pressures.

[0006] In DHF, the heart generally contracts well, with a normal ejection fraction, but is stiffer or less compliant than a healthy heart would be when relaxing and filling with blood. Such patients are considered to have heart failure with preserved ejection fraction ("HFpEF"). This stiffness can prevent blood from filling the heart and produce stasis in the lungs, which can lead to pulmonary venous hypertension and pulmonary edema. HFpEF is more common in patients older than 75 years of age, especially women who suffer from hypertension.

[0007] Both variants of HF have been treated using pharmacological approaches, which typically involve the use of vasodilators to reduce cardiac workload by reducing systemic vascular resistance, and diuretics to prevent fluid accumulation and edema formation and reduce cardiac filling pressure. While no pharmacological therapy has been shown to improve morbidity or mortality in HFpEF, several classes of drugs, including renin-angiotensin antagonists, neprilysin inhibitors, beta-blockers, mineralocorticoid antagonists, and sodium-glucose cotransporter-2 (SGLT2) inhibitors, have had a significant impact on the management of patients with HFrEF. Nevertheless, in general, HF remains a progressive disease, with most patients experiencing deterioration in cardiac function and symptoms over time. In the United States, there are over a million hospitalizations per year from acutely worsening HF, and the mortality rate is higher than most forms of cancer.

[0008] In more severe cases of HFrEF, mechanical circulatory support (MCS) devices, such as mechanical pumps, are used to reduce the load on the heart by performing all or part of the pumping function normally performed by the heart. Chronic left ventricular assist devices ("LVADs"), total artificial hearts, and heart transplants are used as a last resort. However, such assist devices are typically intended to improve the heart's pumping capacity, increase cardiac output to a level compatible with normal lifespan, and sustain the patient until a donor heart for transplant is available. This use of MCS is also known as "bridge to transplant" therapy. More often, MCS is the only therapeutic option, also known as "heart transplant alternative therapy," because the supply of donor hearts for transplant is insufficient relative to the demand. Such mechanical devices allow the propulsion of significant volumes of blood (liters / minute), but are limited by the need for a power source, a relatively large pump, and pose the risk of hemolysis, thrombus formation, and infection. Temporary assist devices, intra-aortic balloons, and pacing devices are also in use.

[0009] Various devices have been developed that use stents to modify blood pressure and flow within a given vessel or between the ventricles of the heart. For example, Ruiz's U.S. Pat. No. 6,120,534 is directed to an intraluminal stent for regulating fluid flow through a body vessel or organ, for example, regulating blood flow through the pulmonary artery and treating congenital heart defects. The stent may include an expandable mesh with balloon expandable lobes or cone-shaped sections joined by shape-memory stenosis regions, which limit flow through the stent. The stenosis regions may be adjusted in vivo, and may also be heated to restore maximum stenosis. Ruiz does not mention the treatment of HF or the reduction of left atrial pressure.

[0010] McNamara's U.S. Patent Publication No. 2013 / 0178784 describes an adjustable pressure relief shunt that can be expanded, for example, via an inflation balloon. The tubular body of the shunt can be plastically deformed in vivo so that the size of the shunt can be repeatedly adjusted by various mechanisms, for example, an elastically wound spring or a series of tabs and a unidirectional mechanical ramp, in response to measurements of the patient's physiological parameters. A significant drawback of the approach described in that patent is the hysteresis effect, i.e., the irreversible change in the underlying crystalline structure that occurs when the shunt is permanently deformed. Importantly, such plastic deformation can lead to stress and fatigue-related fracture of the device. Another drawback of that system is that expanding or reducing the encapsulated shunt after tissue has attached to the shunt over time can result in trauma to the atrial septum.

[0011] U.S. Patent No. 6,468,303 to Amplatz et al. describes a collapsible medical device and associated methods for shunting selected organs and blood vessels. Amplatz explains that the device may be suitable for shunting a patient's cardiac septal defect, for example, by creating a shunt in the atrial septum of a newborn with hypoplastic left heart syndrome ("HLHS"). The patent also explains that increasing the mixing of pulmonary and systemic venous blood improves oxygen saturation, and that the shunt may be subsequently closed with an occlusion device. Amplatz does not mention the treatment of HF or the reduction of left atrial pressure as well as a means for regulating the rate of blood flow through the device.

[0012] Implantable interatrial shunt devices have been used successfully in patients with severe symptomatic heart failure. By diverting or shunting blood from the left atrium ("LA") to the right atrium ("RA"), pressure within the left atrium is reduced or prevented from rising as high as it would otherwise (left atrial decompression). Such implementation would be expected to prevent, alleviate, or limit the symptoms, signs, and associated syndromes of pulmonary congestion. These include severe shortness of breath, pulmonary edema, hypoxia, the need for emergency hospitalization, mechanical ventilation, and death.

[0013] Shunt flow is generally governed by the pressure gradient between the atria and the hydromechanical properties of the shunt device. The latter is typically influenced by the geometry and material composition of the shunt. For example, it has been shown that the general flow quality of similar shunt designs is related to the mean interatrial pressure gradient and the effective orifice diameter.

[0014] Percutaneous implantation of an interatrial shunt generally requires transseptal catheterization immediately prior to shunt device insertion. A transseptal catheterization system is generally placed across the interatrial septum in the region of the fossa ovalis ("FO"), the central and thinnest region of the interatrial septum, from an entry site in the femoral vein. The FO in adults is typically 15-20 mm in its major axial dimension and <3 mm thick, but in some circumstances may be up to 10 mm thick. LA ventricular access may be achieved using a number of different techniques familiar to those skilled in the art, including, but not limited to, needle puncture, stylet puncture, screw needle puncture, and radiofrequency ablation. The passage between the two atria is dilated to facilitate passage of a shunt device having the desired orifice size. Dilation is generally accomplished by advancing a tapered sheath / dilator catheter system or by inflation of an angioplasty type balloon across the FO. This is the same general location where a congenital secundum atrial septal defect ("ASD") would be located.

[0015] US Patent Publication No. 2005 / 0165344 to Dobak, III describes an apparatus for treating heart failure that includes a tubular conduit with an embolic filter or valve, the device being configured to be positioned in an opening in the atrial septum of the heart and allow flow from the left atrium into the right atrium. Dobak discloses that shunting blood can reduce left atrial pressure, thereby preventing pulmonary edema and progressive left ventricular dysfunction and reducing LVEDP. Dobak explains that the device can include deployable retention struts, such as metal arms, that apply slight force on both sides to the atrial septum, clamping or pinching the device to the septum.

[0016] In addition, following implantation of a shunt device into the heart wall, tissue ingrowth, including an endothelial or neointimal layer, typically forms on the device, thereby preventing thrombosis of the shunt device. Shunt adjustment following tissue ingrowth in the atrial septum may present a risk if the shunt is pulled away from the ingrowth tissue in the atrial septum during adjustment. Additionally or alternatively, anchoring in the atrial septum with a relatively large hole, such as that caused by a transseptal intervention such as MitraClip® insertion, may present challenges with shunt fixation. Thus, there is a need for an interatrial shunt that can be safely adjusted and / or can be reliably anchored in a relatively large septal hole. [Prior art documents] [Patent documents]

[0017] [Patent Document 1] U.S. Patent No. 6,120,534 [Patent Document 2] US Patent Application Publication No. 2013 / 0178784 [Patent Document 3] U.S. Patent No. 6,468,303 [Patent Document 4] US Patent Application Publication No. 2005 / 0165344 Summary of the Invention [Means for solving the problem]

[0018] Described herein are devices and methods for making and using improved interatrial shunts to improve treatment and outcomes for patients suffering from cardiovascular and cardiopulmonary disorders, including heart failure (HF) and pulmonary arterial hypertension (PAH). Interatrial shunts may be particularly well suited for in vivo adjustment of the size of the flow path through the shunt and / or for use in relatively large holes in the septum (e.g., following transseptal interventions).

[0019] The device may be designed to shunt blood between a first and a second atrium of a patient to treat a medical condition. The device may include an encapsulated shunt comprising a first diverging end region, a second diverging end region, and a region of reduced diameter disposed therebetween. The encapsulated shunt may be formed from a frame encapsulated in a biocompatible material such as ePTFE such that it defines a passageway and allows blood to flow from the first atrium to the second atrium through the passageway. The device may further include a bridge extending from a first outer surface of the first diverging end region to a second outer surface of the second diverging end region. The bridge may be formed from a biocompatible material and configured to engage the patient's atrial septum.

[0020] The bridge may be configured to engage the patient's atrial septum, rather than the encapsulated shunt itself, so that when the device is adjusted in vivo, the bridge may be configured to remain the same outer diameter, while only the inner diameter of the shunt is modified. The bridge is preferably designed to prevent dehiscence and reduce tissue trauma that may result from such adjustment. The bridge may further be configured to maintain contact with the septal tissue, such that leakage or bypass flow around the outer surface of the shunt is minimized. In addition, the bridge may be used in conjunction with encapsulated shunts that are not in vivo adjustable. For example, incorporating a biocompatible material bridge into the encapsulated shunt increases the outer diameter of the device, thus allowing for implantation of the device in an enlarged septal hole without affecting the inner diameter of the device and the fluid flow rate throughout the device.

[0021] According to another aspect of the present disclosure, a shunt for implantation in the atrial septum to treat a heart disease is provided. The shunt may include a single metal frame including a proximal end region, a distal end region, and an intermediate region disposed therebetween. The proximal and distal end regions may each be configured to self-expand upon deployment such that the intermediate region is positioned at an opening in the atrial septum. The single metal frame may also be heat treated such that the intermediate region is in vivo adjustable between a first diameter and a second diameter greater than the first diameter, and the intermediate region may be selectively thermally shrinkable from the second diameter to a third diameter greater than the first diameter. For example, the intermediate region may be heat treated to exhibit a martensite finish temperature and an austenite finish temperature above the martensite finish temperature such that the intermediate region may be selectively thermally shrinkable by heating to a temperature between the martensite finish temperature and the austenite finish temperature.

[0022] The intermediate region may be malleable at body temperature and may comprise nitinol having an austenite finish temperature of 45-60°C. The proximal and distal end regions may be superelastic and may comprise nitinol having an austenite finish temperature of 5-20°C. The intermediate region may also be heat treated to exhibit different shape memory properties than the proximal and distal end regions. Additionally, the intermediate region may comprise a plastically deformable material such that the intermediate region may be configured to be expanded in vivo via mechanical expansion. The proximal and distal end regions and the intermediate region may define a diabolo-shaped shunt.

[0023] The system may further include a biocompatible material coating the proximal end region, the distal end region, and the intermediate region to define a passageway for allowing blood to flow through the atrial septum via the shunt. A cross-sectional flow area within the passageway may be smallest in the intermediate region. Additionally, the passageway may be sized and shaped to allow a sufficient amount of blood to flow through the atrial septum via the shunt to treat pulmonary arterial hypertension. Additionally or alternatively, the passageway may be sized and shaped to allow a sufficient amount of blood to flow through the atrial septum via the shunt to treat heart failure. In some embodiments, the system may further include a bridge extending from a first outer surface of the proximal end region to a second outer surface of the distal end region. The bridge may be formed from a biocompatible material and configured to engage the atrial septum. Additionally, the bridge may define an outer diameter greater than the second diameter. For example, the bridge may extend to form a gap between an inner surface of the bridge and an outer surface at the intermediate region such that when the intermediate region is expanded in vivo, the gap may be configured to decrease in size and the outer diameter may be configured to remain the same diameter Additionally, or alternatively, the bridge may extend to form a gap between an inner surface of the bridge and an outer surface at the intermediate region such that when the intermediate region is contracted in vivo, the gap is configured to increase in size and the outer diameter is configured to remain in contact with the septum.

[0024] According to another aspect of the disclosure, a system for treating a heart disease is provided that includes a shunt. For example, the system may further include a catheter configured to inject a heated fluid into the shunt in vivo, heat the shunt in vivo, and selectively thermally shrink the intermediate region from a second diameter to a third diameter. For example, the heated fluid may have an initial fluid temperature selected based on at least one of an injection rate of the heated fluid or a size of the catheter to reach a target fluid temperature in the shunt in vivo, thereby heating at least the intermediate region of the shunt to a target shunt temperature. The catheter may also include a temperature sensor disposed on a distal tip of the catheter, the temperature sensor configured to generate a signal indicative of a temperature of the heated fluid delivered into the shunt. The catheter may be sized and shaped to inject the heated fluid at an injection rate selected to minimize dilution and loss of the heated fluid in vivo. In some embodiments, the system may include an inflatable balloon configured to be positioned adjacent the distal end region in an expanded state to block blood flow through the shunt during injection of the heated fluid to minimize dilution and loss of the heated fluid in vivo.

[0025] In addition, the system may further include an automatic injector fluidly coupled to the catheter for injecting heated fluid through the catheter. The automatic injector may be operatively coupled to the temperature sensor and programmed to adjust at least one of a rate or duration of heated fluid injection through the catheter in response to a signal. The automatic injector may be programmed to automatically terminate injection of heated fluid through the catheter when a temperature of the heated fluid delivered within the shunt reaches a predetermined temperature sufficient to selectively thermally shrink an intermediate region of the shunt to a third diameter. The automatic injector may also be programmed to modulate a rate of heated fluid injection through the catheter to maintain a temperature of the heated fluid delivered within the shunt for a predetermined period of time.

[0026] According to yet another aspect of the invention, a method for treating heart disease is provided that may include delivering a shunt in a contracted state to an opening in the atrial septum, the shunt comprising a frame having proximal and distal end regions and a mid-region disposed therebetween, the frame being heat treated such that the mid-region is in vivo adjustable between a first diameter and a second diameter greater than the first diameter, deploying the shunt in the opening to an expanded state, the proximal end region self-expanding in the first atrium and the distal end region self-expanding in the second atrium, the mid-region disposed in the atrial septum, expanding the mid-region of the shunt in vivo from the first diameter to the second diameter, and selectively thermally contracting the mid-region of the shunt in vivo from the second diameter to a third diameter greater than the first diameter.

[0027] Expanding the intermediate region of the shunt may include mechanically expanding the intermediate region of the shunt. Also, selectively thermally contracting the intermediate region of the shunt may include heating the intermediate region to a temperature between a martensite finish temperature and an austenite finish temperature, for example, by injecting a heated fluid into the shunt in vivo via a catheter to heat at least the intermediate region of the shunt. The method may further include measuring a temperature of the heated fluid delivered into the shunt via a temperature sensor disposed at a distal tip of the catheter, and adjusting at least one of a rate or duration of heated fluid injection through the catheter based on the measured temperature of the heated fluid delivered into the shunt. Additionally, the method may include automatically terminating the injection of heated fluid through the catheter when a temperature of the heated fluid delivered into the shunt reaches a predetermined temperature sufficient to selectively thermally contract the intermediate region of the shunt to a third diameter, and / or modulating a rate of heated fluid injection through the catheter to maintain a temperature of the heated fluid delivered into the shunt for a predetermined period of time. The method may further include expanding an inflatable balloon adjacent a distal end region of the shunt to an expanded state to block blood flow through the shunt during injection of the heated fluid to minimize dilution and loss of the heated fluid in vivo. The method may also include allowing a sufficient amount of blood to flow through the shunt and through the atrial septum to treat pulmonary arterial hypertension and / or heart failure. [Brief description of the drawings]

[0028] These and other features, aspects, and advantages of the present invention will become apparent from the following description, appended claims, and the accompanying exemplary embodiments illustrated in the drawings, which are briefly described below.

[0029] [Figure 1A]1A and 1B are side views of hourglass shaped shunts with and without bridges constructed in accordance with the methods of the present invention. [Figure 1B] 1A and 1B are side views of hourglass shaped shunts with and without bridges constructed in accordance with the methods of the present invention.

[0030] [Figure 1C] 1C-1F are side, cross-sectional, and perspective views of the device of FIG. 1B. [Figure 1D] 1C-1F are side, cross-sectional, and perspective views of the device of FIG. 1B. [Figure 1E] 1C-1F are side, cross-sectional, and perspective views of the device of FIG. 1B. [Figure 1F] 1C-1F are side, cross-sectional, and perspective views of the device of FIG. 1B.

[0031] [Figure 2A] 2A and 2B are cross-sectional side views of the device of FIG. 1B. [Figure 2B] 2A and 2B are cross-sectional side views of the device of FIG. 1B.

[0032] [Figure 3A] 3A-3E illustrate exemplary steps for using the device of FIG. 1B within the human body. [Figure 3B] 3A-3E illustrate exemplary steps for using the device of FIG. 1B within the human body. [Figure 3C] 3A-3E illustrate exemplary steps for using the device of FIG. 1B within the human body. [Figure 3D] 3A-3E illustrate exemplary steps for using the device of FIG. 1B within the human body. [Figure 3E] 3A-3E illustrate exemplary steps for using the device of FIG. 1B within the human body.

[0033] [Figure 4A]4A-4D illustrate the construction of shunt-graft assemblies having two and three layer regions. [Figure 4B] 4A-4D illustrate the construction of shunt-graft assemblies having two and three layer regions. [Figure 4C] 4A-4D illustrate the construction of shunt-graft assemblies having two and three layer regions. [Figure 4D] 4A-4D illustrate the construction of shunt-graft assemblies having two and three layer regions.

[0034] [Figure 5A] 5A-5E are side views sequentially illustrating a technique for depositing a first graft portion onto a shunt. [Figure 5B] 5A-5E are side views sequentially illustrating a technique for depositing a first graft portion onto a shunt. [Figure 5C] 5A-5E are side views sequentially illustrating a technique for depositing a first graft portion onto a shunt. [Figure 5D] 5A-5E are side views sequentially illustrating a technique for depositing a first graft portion onto a shunt. [Figure 5E] 5A-5E are side views sequentially illustrating a technique for depositing a first graft portion onto a shunt.

[0035] [Figure 6A] 6A and 6B are side views sequentially illustrating a technique for depositing a second graft portion onto a shunt. [Figure 6B] 6A and 6B are side views sequentially illustrating a technique for depositing a second graft portion onto a shunt.

[0036] [Figure 7A]7A and 7B illustrate a mandrel assembly including a first mandrel portion illustrated in FIG. 7A and a second mandrel portion illustrated in FIG. 7B. [Figure 7B] 7A and 7B illustrate a mandrel assembly including a first mandrel portion illustrated in FIG. 7A and a second mandrel portion illustrated in FIG. 7B.

[0037] [Figure 8A] 8A-8C are side and close-up views illustrating a technique for depositing a third graft portion onto a shunt. [Figure 8B] 8A-8C are side and close-up views illustrating a technique for depositing a third graft portion onto a shunt. [Figure 8C] 8A-8C are side and close-up views illustrating a technique for depositing a third graft portion onto a shunt.

[0038] [Figure 9A] 9A and 9B illustrate a perspective view of a flexible sleeve and a side view of the flexible sleeve mounted onto a shunt-graft assembly. [Figure 9B] 9A and 9B illustrate a perspective view of a flexible sleeve and a side view of the flexible sleeve mounted onto a shunt-graft assembly.

[0039] [Figure 10A] 10A and 10B illustrate a perspective view of the compression shell and a side view of the compression shell mounted onto the flexible sleeve and shunt-graft assembly. [Figure 10B] 10A and 10B illustrate a perspective view of the compression shell and a side view of the compression shell mounted onto the flexible sleeve and shunt-graft assembly.

[0040] [Figure 11A] 11A-11F illustrate a first embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge. [Figure 11B] 11A-11F illustrate a first embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge. [Figure 11C] 11A-11F illustrate a first embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge. [Figure 11D] 11A-11F illustrate a first embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge. [Figure 11E] 11A-11F illustrate a first embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge. [Figure 11F] 11A-11F illustrate a first embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge.

[0041] [Figure 12A] 12A-12D illustrate a second embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge. [Figure 12B] 12A-12D illustrate a second embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge. [Figure 12C] 12A-12D illustrate a second embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge. [Figure 12D] 12A-12D illustrate a second embodiment of the shunt-graft assembly of FIGS. 4A-4D having a bridge.

[0042] [Figure 13] FIG. 13 illustrates another embodiment of a shunt-graft assembly having a bridge.

[0043] [Figure 14A] 14A-14B illustrate images of the encapsulated stent of FIGS. 12A-12D. [Figure 14B]14A-14B illustrate images of the encapsulated stent of FIGS. 12A-12D.

[0044] [Figure 15] FIG. 15 is a graph illustrating incremental thermal contraction of a shunt orifice in accordance with the principles of the present disclosure.

[0045] [Figure 16] FIG. 16 is a flow chart illustrating exemplary method steps for treating cardiac disease in accordance with the principles of the present disclosure. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0046] Detailed Description Described herein are apparatus and methods for making and using an improved inter-atrial shunt to improve treatment and outcomes for patients suffering from cardiovascular and cardiopulmonary disorders such as pulmonary arterial hypertension (PAH) or heart failure (HF). In some aspects, the device has dimensions that can be reduced and increased in vivo.

[0047] The device may include one or more components that can be permanently or temporarily implanted in the human body and adjusted in size, larger or smaller, after implantation. The need for such an adjustable device may arise, for example, in the treatment of pulmonary arterial hypertension (PAH) or heart failure (HF). In PAH, placing a shunt in the interatrial septum allows excess blood pressure in the right atrium to be relieved by allowing some blood to flow from the right atrium to the left atrium through an orifice. In HF, placing a shunt in the interatrial septum allows excess blood pressure in the left atrium to be relieved by allowing some blood to flow from the left atrium into the right atrium through an orifice. In both PAH and HF, interatrial shunts have been shown to effectively reduce symptoms and increase exercise tolerance. Interatrial shunts may also reduce the need for hospitalization and even improve life expectancy.

[0048] However, if the orifice of the interatrial shunt is too small, only very little blood can be transferred and the shunt is relatively ineffective and may provide little or no clinical benefit. In contrast, shunting too much blood ("over-shunt") through an orifice that is too large may lead to serious or even fatal complications over time. For example, in PAH patients, over-shunting may result in systemic oxygen desaturation and its sequelae, including cyanosis, polycythemia with increased blood viscosity, end-organ ischemia, and potentially, death. In HF patients, over-shunting may result in pulmonary hypertension, right ventricular failure, and potentially, death.

[0049] Currently, there is no known method for predicting a given patient's response to a particular shunt orifice size. As previously known, the shunt orifice can be increased in vivo by expanding a suitably designed shunt, for example, by expanding an inflatable balloon catheter or other similar mechanical expansion means within the shunt, provided, however, that the shunt is made of a malleable material and remains expanded due to plastic deformation or some other physical property, so that when the balloon or other expansion means is removed, the amount of elastic rebound or recoil will be low enough so that the desired increment in orifice size is achieved. One disadvantage of this approach is that the orifice size can only be increased. If the shunt starts out too large or is made too large by balloon expansion, but the patient requires a smaller shunt, there is no way to return to a smaller size orifice other than by providing another smaller shunt or placing a smaller shunt within the lumen of the original shunt. This technique is known as a "shunt within a shunt." Thus, finding a suitable shunt orifice size for a given patient is a trial and error process, with the shunt orifice size being selected according to the patient's response, which may be as short as a few minutes or as long as several months, and the shunt orifice size may be increased (e.g., by balloon expansion) or decreased (by providing a new, smaller shunt) depending on the patient's response. Thus, the opportunity to increase or decrease the size of the shunt may be very limited and not reproducible. Furthermore, the extent to which an inflatable balloon catheter can expand the shunt orifice may be limited by the maximum size of the balloon. Thus, what is needed is a means to repeatedly and non-traumatically adjust the orifice size of shunts and other implantable devices in vivo, and in both larger and smaller directions.

[0050] In some examples, the devices provided herein may incorporate technology with adjustable cross-sectional flow areas that can be easily reduced in vivo and / or expanded in vivo in any order as clinically required. Examples of interatrial shunts with adjustable cross-sectional flow areas are described in U.S. Pat. No. 9,724,499 to Rottenberg et al., U.S. Pat. No. 10,898,698 to Eigler et al., WO 2021 / 224736, and U.S. Patent Application Publication No. 2021 / 0121179 to Ben-David et al., each of which is assigned to the assignee of the present application, the entire contents of each of which are incorporated herein by reference. Some examples of the devices include self-expanding superelastic (austenite phase) materials as well as malleable shape memory (martensite phase) materials. When the device is implanted in the human body, for example by transporting the device in a compressed state in a sheath to a desired location and then removing the sheath, the self-expanding superelastic material may automatically expand to its desired size, while the malleable shape memory material may initially remain in a reduced size state. The cross-sectional area of ​​the malleable shape memory material may then be expanded and reduced in vivo as desired, for example by providing a suitable fluid flow rate therethrough, to obtain a cross-sectional area suitable for treating the patient, or to properly secure the device within the patient while allowing repositioning to improve the effectiveness of treatment. A wide variety of devices may be prepared using components that include the self-expanding superelastic material and the malleable shape memory material, each as illustrated herein.

[0051] One complication that may occur when adjusting the size of the shunt over time is tissue trauma to the atrial septum. The encapsulated shunt may be designed to encourage tissue ingrowth and endothelialization, and thus expanding or reducing the encapsulated shunt after tissue has attached to it over time may result in trauma to the atrial septum. Provided herein is a device for adjusting the size of the shunt without disturbing the septal tissue surrounding the device. In particular, the device may include a bridge formed of a biocompatible material that extends between the outer surfaces of the first and second flared end regions and creates a gap between the bridge and the reduced diameter region of the encapsulated shunt. The bridge may be configured to engage the patient's atrial septum, rather than the encapsulated shunt itself, such that when the device is adjusted in vivo, the bridge may be configured to remain the same outer diameter, while only the inner diameter of the shunt is modified. Thus, the bridge prevents dehiscence that may result when the device is adjusted in vivo. Additionally, the bridge may mitigate any bypass flow that may flow around the outside of the device after the diameter of the necked down region is reduced in vivo.

[0052] The bridge described above may also be used with encapsulated shunts that are not in vivo adjustable. Patients who may benefit from an interatrial shunt may also have required or will require a prior transseptal procedure resulting in a hole in the septal wall. Alternatively, patients may have a pre-dilated septal defect that is larger than the delivery system required to implant the device described herein. Incorporating a biocompatible material bridge into an encapsulated shunt increases the outer diameter of the device, thus allowing for implantation of the device in an enlarged septal hole without affecting the inner diameter of the device and the fluid flow rate throughout the device.

[0053] In some examples, the device may be or include an hourglass or "diabol" shaped shunt, which may be optionally encapsulated with a biocompatible material and used to treat a subject suffering from a disorder for which it may be useful to regulate fluid flow, such as CHF or PAH. In some examples, the hourglass shaped shunt may be specifically configured to be securely inserted into the atrial septum, e.g., into an opening through the fossa ovalis, to allow blood flow from the left atrium to the right when the blood pressure in the left atrium exceeds that of the right atrium, or blood flow from the right atrium to the left when the blood pressure in the right atrium exceeds that of the left atrium. As provided herein and described in more detail in PCT Application No. WO 2021 / 224736, incorporated above, the internal dimensions of the hourglass shaped shunt may be suitably adjusted in vivo, e.g., to regulate the flow of fluid therethrough, e.g., to regulate the flow of fluid between the left and right atria through the atrial septum.

[0054] 1A and 1B, a shunt 100 is illustrated with or without a bridge. The shunt 100 may be hourglass or "diabolo" shaped and include a first component 110, a second component 120, and a third component 130 that are fluidly coupled to one another. In some embodiments, the shunt 100 has internal dimensions that can be reduced and increased in vivo. The first component 110 may include a first self-expanding superelastic material, the second component 120 may include a malleable shape memory material, and the third component 130 may include a second self-expanding superelastic material. The malleable shape memory material of the second component 120 may have a first cross-sectional area (or diameter) that allows a first rate of fluid flow through the second component, may be expandable to a second cross-sectional area (or diameter) that allows a second rate of fluid flow through the second component, and may be contractible to a third cross-sectional area (or diameter) that allows a third rate of fluid flow through the second component. The overall rate of fluid flow through the device 100 may also depend on the cross-sectional areas (or diameters) of the first component 110 and the third component 130.

[0055] The shunt 100 may additionally or alternatively be constructed as described in U.S. Pat. Nos. 9,707,382, 9,980,815, and 10,639,459 to Nitzan et al., U.S. Pat. Nos. 10,076,403, 10,251,740, and 11,291,807 to Eigler et al., and U.S. Pat. No. 10,835,394 to Nae et al., each of which is assigned to the assignee of the present application (the entire contents of each of which are incorporated herein by reference).

[0056] The first component 110 may include any suitable number of rings, e.g., rings 112, 113, which are formed from or include a first self-expanding material and may optionally be sinusoidal. The second component 120 may include any suitable number of rings, e.g., ring 114, which are formed from or include a malleable shape memory material and may optionally be sinusoidal. The third component 130 may include any suitable number of rings, e.g., rings 115, 116, which are formed from or include a third self-expanding material and may optionally be sinusoidal. The struts 111, 108 may interconnect the rings of the first component 110, the second component 120, and the third component 130.

[0057] The first component 110 may provide a first diverging end region 102, the third component 130 may provide a second end diverging region 106, and the second component 120 may provide a neck region 104 disposed between the first and second diverging end regions. The inlet and outlet of the device 100 may include flanges 102, 106, and the neck region 104 may include flexible longitudinal struts 111, 108 and a sinusoidal ring 114. The flexible longitudinal struts 111, 108 may allow the flanges to fully expand upon deployment, and the sinusoidal ring 114 may have sufficient strength to maintain its diameter when balloon expanded or heat shrunk.

[0058] In the non-limiting embodiment shown in FIG. 1A, the first diverging end region 102 has a first end region dimension D1, the second diverging end region 106 has a second end region dimension D2, and the narrowing region 104 has a narrowing dimension D3, which may be increased or decreased in a manner as described with reference to the second component 120 illustrated in FIGS. 3A-3D. As shown in FIG. 1A, the narrowing region 104 of the shunt 100 may be significantly narrower than the diverging end regions 102 and 106, e.g., having a smaller cross-sectional area and smaller dimensions than those of the diverging end regions 102 and 106. Also, as shown in FIG. 1A, the shunt 100 may be asymmetric. For example, the shunt 100 may be asymmetric and take advantage of the natural characteristics of the atrial septum and the left and right atrial cavities of the heart. Alternatively, the hourglass shaped shunt 100 may be symmetrical, with the first end region dimension D1 equal to the second end region dimension D2. The first and second diverging end regions 102, 106 may also have either or both straight or curved profiles. For example, the struts 111 have a straight profile and the struts 108 have a curved profile. Additionally, the first and second diverging end regions 102, 106 may assume any angular position consistent with an hourglass configuration.

[0059] The device 100 may optionally be fabricated from a single tube of material that is laser cut to define a plurality of struts and connecting members, e.g., a plurality of sinusoidal rings connected by longitudinally extending struts. The sinusoidal rings and longitudinal struts may be laser cut to form a unitary piece of unitary construction, and different regions of the piece may be heat treated differently from one another in a manner as described elsewhere herein to produce components having different austenite finish temperatures (Af) from one another. Alternatively, the sinusoidal rings of the first component 110, the second component 120, and the third component 130 may be defined separately to form different pieces of material with suitable Af that are subsequently bonded together to form the device 100. The device 100 may also be electropolished to reduce thrombus formation.

[0060] In some embodiments, the first self-expanding superelastic material of the first component 110, the malleable shape memory material of the second component 120, and the second self-expanding superelastic material of the third component 130 may comprise materials different from each other, or may comprise materials that are the same as each other but have different phases from each other. For example, the first component 110, the second component 120, and the third component 130 may independently comprise one or more materials selected from the group consisting of nickel titanium (NiTi), also known as Nitinol, other shape memory alloys, self-expanding materials, superelastic materials, polymers, and the like. In one non-limiting embodiment, the first component 110 and the third component 130 may each comprise a Nitinol alloy having an austenite finish temperature (Af) well below body temperature, which material is in an austenitic superelastic phase while in the human body. In one non-limiting example, the self-expanding superelastic material of the first component 110 and the third component 130 includes Nitinol having an Af of less than 37° C. For example, the Af of the self-expanding superelastic material Nitinol may be between 5 and 20° C. The first component 110, the second component 120, and the third component 130 may be optionally integrally formed from a common frame with each other. For example, the first component 110, the second component 120, and the third component 130 may be initially cut and processed as a single unit from a frame of tubing, sheet, or other suitable configuration that is identical to each other. Portions of that common frame may be heat treated differently from each other to define the first component 110, the second component 120, and the third component 130, for example, in a manner similar to that described in PCT Application No. WO 2021 / 224736, incorporated above.

[0061] The second component 120 may comprise a nitinol alloy having an austenite phase transition temperature Af slightly above body temperature such that the material remains in its martensite malleable shape memory phase while in the body unless and until it is heated to its Af, for example, by injection of lukewarm or hot saline (or other fluid) into the fluid flowing in or through the second component 120, or by application of heat through electrical energy, such as with an RF energy source. In one non-limiting example, the malleable shape memory material of the second component 120 comprises nitinol, having an austenite finish temperature (Af) greater than 37°C. For example, the Af of the nitinol of the malleable shape memory material of the second component 120 may be 45-65°C, e.g., 50-55°C. In some embodiments, lukewarm or hot saline (or other fluid) may be injected using a side-hole catheter positioned through the device 100, close enough to the second component 120 to heat that component to or above its Af. Optionally, an expandable balloon may be placed on the distal end of the side-hole catheter and inflated at the distal opening of the shunt so that blood flow in the shunt is blocked during delivery of the saline. In a similar embodiment, an expandable balloon may be placed proximal to the distal end of the side-hole catheter and inflated at the proximal opening of the shunt so that blood flow in the shunt is again blocked during delivery of the saline. In yet another embodiment, lukewarm or hot saline may be injected through the distal end of the central lumen of the catheter, positioned adjacent to or proximal to the second component 120. Preferably, the distal end of the catheter includes a larger hole than the side hole described above so that the heated saline can be delivered more quickly. Optionally, a separate balloon catheter may be inserted through device 100 such that the balloon is distal to the catheter for delivering the saline. Prior to injecting saline through the distal end of the catheter, a balloon may be expanded at the distal opening of device 100 to block blood flow during delivery of the saline, which may increase the effectiveness of the heating.The use of separate catheters to deliver saline and block blood flow may allow the saline to heat the stent more quickly while the expanded balloon is kept in place.

[0062] In other embodiments, a pair of electrodes may be contacted with the device 100, for example, via a catheter, and operated at an appropriate voltage and frequency to heat the component 120 to or above its Af. In still other embodiments, any other suitable means of locally applying heat to the device 100 may be used, such as a laser, magnetic inductance, electrical resistance, or the like. Heating the device 100 using electrical resistance may include contacting the device with a pair of electrodes, for example, via a catheter, and passing an electrical current through the device, which causes the device to heat. Heating the device 100 using a laser may include illuminating the device with light from a laser, which may be introduced by a catheter. Heating the device 100 using magnetic inductance may include passing an alternating magnetic field through the device, which heats the device, induces eddy currents inside the device. It should be noted that in blood vessels with particularly high blood flow rates (e.g., 2-5 L / min), such as the aorta or internal iliac arteries, it may be useful to heat the device 100 using direct heating methods instead of saline, such as using electrical energy (e.g., direct current (DC), radio frequency (RF)), laser, magnetic inductance, electrical resistance, non-contact radio frequency (RF), dielectric heating, or conductive heating such as local probe heating, before fully heating the device.

[0063] Alternatively, device 100 may include a single Nitinol alloy that has been heat treated to produce a lower Af in the regions corresponding to first component 110 and third component 130, respectively, and a higher Af in the region corresponding to second component 120. The malleable shape memory material of second component 120 may be expandable and contractable using any suitable technique. For example, the malleable shape memory material of second component 120 may be mechanically expanded, for example, using balloon expansion as known in the art. Additionally or alternatively, the malleable shape memory material of second component 120 may be thermally contracted, for example, using saline at or above the Af of the material, or otherwise heated, such as with the use of RF energy or lasers, magnetic inductance, electrical resistance, or the like, in the manner described above.

[0064] The first, second, and third components may be joined, for example, fluidly joined, to one another using any suitable joining manner. For example, any malleable shape memory material may be optionally and independently joined to any self-expanding superelastic material by welding. Additionally or alternatively, any malleable shape memory material may be optionally and independently joined to any self-expanding superelastic material using an encapsulant that may cover at least a portion of at least one of the components and join such components to one another. Additionally or alternatively, any shape memory material and any self-expanding superelastic material may be integrally formed from a common frame with one another.

[0065] The encapsulant may comprise any suitable biocompatible material, such as a polymer or natural material. Examples of polymers suitable for use as the encapsulant include expanded polytetrafluoroethylene (ePTFE), silicone, polycarbonate urethane, Dacron (polyethylene terephthalate), ultra-high molecular weight polyethylene (UHMWPE), and polyurethane. Examples of natural materials suitable for use as the encapsulant include human tissue, such as pericardial tissue from equine, bovine, or porcine sources, or human placenta or other human tissue. The biocompatible material is preferably smooth to discourage thrombus formation, and may optionally be impregnated with carbon to encourage tissue ingrowth. Alternatively, the biocompatible material may form a mesh-like structure to encourage tissue ingrowth and endothelialization. The device may be encapsulated with a biocompatible material in a manner similar to that described in U.S. Patent No. 11,304,831 to Nae et al., entitled "Systems and Methods for Making Encapsulated Hourglass Shaped Stents," the entire contents of which are incorporated herein by reference. Additional methods for encapsulating shunts are described in U.S. Patent No. 10,835,394 to Nae et al., U.S. Patent No. 11,109,988 to Rosen et al., U.S. Patent No. 9,034,034 to Nitzan et al., U.S. Patent No. 9,980,815 to Nitzan et al., and U.S. Patent No. 10,076,403 to Eigler, the entire contents of each of which are incorporated herein by reference.

[0066] In one embodiment, the device is encapsulated with ePTFE. It will be understood by those skilled in the art that ePTFE material has a characteristic microstructure consisting of nodes and fibrils, with the fibril orientation being approximately parallel to the axis of longitudinal expansion. Expanded polytetrafluoroethylene material may be made by ram extruding a compressed mass of particulate polytetrafluoroethylene and extrusion lubricant through an extrusion die to form a sheet or tubular extrudate. The extrudate is then longitudinally expanded and heated to or above the crystalline melting point of polytetrafluoroethylene, i.e., 327°C, for a period of time sufficient to form a node-fibril microstructure and sinter the ePTFE material. Heating may be performed in a vacuum chamber to prevent or inhibit oxidation of the device. Alternatively, heating may be performed in a nitrogen-rich environment. A furnace may be used to heat the encapsulated device. Alternatively, or in addition, a mandrel on which the encapsulated device rests may be used to heat the encapsulated device.

[0067] Referring now to FIG. 1B, the device shown in FIG. 1A is modified to add a bridge at the constriction region 104 configured to engage the patient's atrial septum. The device 100 is preferably encapsulated with a graft material to create the shunt-graft assembly and a passageway through which blood can flow, the graft material being omitted from FIG. 1B to better illustrate the location of the bridge 240. The bridge 240 may be made of a biocompatible material, such as a polymer or natural material as described above, and may be the same material as the material used to encapsulate the frame, or may be a different material. Preferably, the biocompatible material of the bridge is one that encourages tissue attachment so that contact with the septal wall is maintained when the inner diameter of the shunt is reduced. Maintaining contact with the septal wall helps prevent any fluid bypass around the outside of the device. The biocompatible material may be configured to encourage tissue ingrowth over the entire bridge or over only a portion of the bridge. For example, holes may be placed at locations of the bridge 240 that are configured to engage the atrial septum, while the remainder of the bridge 240 remains intact such that tissue ingrowth is not encouraged over the flared end regions. This configuration may be beneficial with respect to the embodiment shown in Figures 12A-12D, in which the biocompatible material of the bridge 240 extends over the flared end regions and along the interior of the encapsulated shunt.

[0068] Additionally or alternatively, the bridge 240 may be made from a different biocompatible material than the biocompatible material used to encapsulate the shunt. For example, the shunt may be encapsulated with a biocompatible material, such as ePTFE, that has a pore size small enough so that tissue ingrowth is mitigated, and the bridge may be made from a biocompatible material with a larger pore size designed to encourage tissue ingrowth. Generally, the larger the pore size of the biocompatible material, the greater the attachment of tissue to the biocompatible material. In addition to encouraging tissue ingrowth, the greater porosity allows for the exchange of fluids in and out of the gap between the necked down region 104 and the outer surface of the bridge 240. For example, the bridge 240 may be made from ePTFE that has a larger internodal distance (e.g., about 60-200 μm) than the ePTFE that encapsulates the shunt.

[0069] Alternatively, the bridge 240 may be made from woven Dacron to further encourage tissue ingrowth. The Dacron may be securely attached to the encapsulated shunt using stitches rather than the method described below. Because Dacron is bulkier than ePTFE, the cross-section of the device in a collapsed or crimped configuration may be increased, which may mean that a larger diameter sheath may be required for delivery of the device. Additional materials that may be used to encourage tissue ingrowth include using a mesh-like structure, electrospun fabric, or silicone.

[0070] The bridge 240 may have a first end 241 and a second end 242, and may be shaped and sized such that the first end 241 is disposed approximately midway through the first diverging end region 102 and the second end 242 is disposed approximately midway through the second diverging end region 106. Alternatively, the first end 241 and the second end 242 may extend further up the first diverging end region 102 and the second diverging end region 106 or may be attached closer to the necking down region 104. The bridge 240 may be elongated such that a gap is created between the outer surface of the necking down region 104 and the inner surface of the bridge 240. The gap may be widest at the narrowest point of the outer surface of the necking down region 104.

[0071] As explained above, the encapsulated shunt may be adjusted in vivo to increase or decrease the necking dimension, thereby adjusting the fluid flow rate through the shunt. The encapsulated shunt may be designed to encourage tissue ingrowth and endothelialization, so that tissue may attach to the shunt over time. Adjustment of the encapsulated shunt to increase or decrease the dimension may therefore result in trauma to the atrial septum. The bridge 240 is designed to prevent dehiscence and reduce tissue trauma that may result from such adjustment. The bridge 240 is configured to engage the atrial septum and defines an outer diameter D4. Preferably, the outer diameter D4 is greater than the necking dimension D3. In one embodiment, the outer diameter D4 may be 7-9 mm, and the necking dimension D3 may be 4.5-5.5 mm. When the device is adjusted in vivo, the bridge 240 may be configured to remain the same outer diameter D4, while only the necking dimension D3 and the size of the gap of the shunt are modified. Due to the creation of a gap between the narrowing region 104 and the bridge 240, the narrowing dimension D3 can be reduced or increased to an outer diameter D4, causing an increase or decrease in the size of the gap, while maintaining contact with the septal tissue so that leakage or bypass flow around the outer surface of the shunt is minimized without disturbing the septal tissue that contacts and surrounds the bridge 240.

[0072] The bridge 240 may also be used with encapsulated shunts that are not in vivo adjustable. In particular, incorporating the bridge 240 into the encapsulated shunt may be beneficial for patients who have a pre-dilated septal defect, e.g., from a prior transseptal procedure, that has an enlarged hole prior to implantation of the device, or that is larger than the delivery system required to implant the device described herein. For example, for patients with severe mitral regurgitation and poor left ventricular function, it may be clinically desirable to first perform a repair procedure on the mitral valve, e.g., MitraClip® mitral valvuloplasty via a percutaneous transseptal approach, followed by interatrial shunt placement. These mitral valve procedures currently use a 23 Fr ID (approximately 8 mm outer diameter) guiding catheter to cross the foramen ovale. After mitral valve repair, a shunt with a minimum outer diameter that matches the larger opening defect caused by the prior procedure may be implanted, with the conduit being the smaller diameter (e.g., 5.0-6.5 mm) desired for the shunt. Similarly, such a shunt may be advantageously used in cases where the fossa ovalis is torn during a transseptal procedure, thus creating a larger opening defect than that required for the embodiment shown in FIG. 1A.

[0073] Incorporating a bridge of biocompatible material into the encapsulated shunt increases the outer diameter of the device, thus allowing implantation of the device in an enlarged septal hole without affecting the inner diameter and fluid flow rate throughout the device. Additionally, the bridge 240 allows the inner diameter of the encapsulated shunt to be temporarily increased, for example, during a separate transseptal procedure after implantation of the device, without disturbing the outer diameter of the narrowing region, thus minimizing the risk of tears to the septal tissue.

[0074] 1C, 1D, 1E, and 1F, additional side, front, side cross-sectional, and perspective views, respectively, are shown. FIG 1C shows a shunt 100 that is encapsulated with a biocompatible material 117 to produce a shunt-graft assembly 210 and define a flow path through the shunt-graft assembly 210. As shown in FIG 1E, the necking region 104 defines the inner diameter of the passageway through which blood flows, i.e., necking dimension D3. A bridge 240 surrounds the entire necking region 104 and defines an outer diameter D4.

[0075] Reference is now made to Figures 2A and 2B, which are cross-sectional side views of the device of Figure 1B, showing the unfilled and filled gaps between the bridge and the encapsulated shunt. The bridge 240 is attached at a first end 241 to the first diverging end region 102 and at a second end 242 to the second diverging end region 106. An exemplary method of attaching the bridge 240 to the encapsulated shunt is further described below with respect to Figures 11A-11D and 12A-12D. Preferably, the bridge 240 is stretched such that a gap is created between the narrowing region 104 and the bridge 240. The gap 243 may be filled with a flexible biocompatible material or a liquid biocompatible material such as a hydrogel. Alternatively, the gap 243 may be filled with bodily fluids upon delivery and implantation of the shunt-graft 210. This embodiment is advantageous when the dimensions of the shunt are configured to be adjusted within the living body. Specifically, gap 243 may be increased or decreased during such adjustment, while the outer diameter of the constriction region (e.g., bridge) that contacts the atrial septum remains the same, thus minimizing any tissue trauma. For example, bridge 240 may be made from ePTFE having holes such that a biocompatible material disposed within gap 243 is configured to permeate through bridge 240 when shunt-graft assembly 210 expands.

[0076] As shown in FIG. 2B, the gap between the bridge 240 and the encapsulated shunt may be filled with a biocompatible material, such as a polymer or a non-flexible natural material, so that the gap remains the same size even if the dimensions of the shunt-graft assembly 210 are adjusted. This embodiment may be beneficial when the device is designed to be implanted into an enlarged septal hole and the device is not configured to adjust in vivo. In particular, since the inner diameter of the device cannot be adjusted, the dimensions of the gap 244 do not need to be adjustable. Thus, the gap 244 may be filled with a solid material or a low durometer material, such as a hydrogel, which may increase the stability of the device and make the bridge more robust.

[0077] 3A-3E diagrammatically illustrate exemplary steps for using the device of FIG. 1B within the human body. The shunt-graft assembly 210 may be crimped into a cylindrical shape, for example, by pushing it through a conical loading device. In one non-limiting example, the shunt-graft assembly 210 may be crimped to an outer dimension of about 4.6 mm, which is the inner dimension of a 14F Cook sheath. For example, FIG. 3A illustrates the shunt-graft assembly 210 disposed within a sheath 300. As would be understood by one of ordinary skill in the art, the shunt-graft assembly 210 may be crimped to a smaller or larger outer dimension if a different size Cook sheath is used. In addition, a layer of hydrogel may be placed with a gap 244 between the bridge 240 and the shunt-graft assembly 210, which may affect the crimping dimensions. The delivery catheter and sheath 300 may be designed as described in commonly assigned U.S. Pat. No. 9,713,696 to Yacoby et al. and U.S. Patent Publication No. 2020 / 0315599 to Nae et al., entitled "Systems and Methods for Delivering Implantable Devices across an Atrial Septum," each of which is incorporated herein by reference in its entirety. The sheath may be placed percutaneously through a blood vessel to a desired location within the body. As the crimped shunt is pushed out of the sheath 300, the self-expanding superelastic gradually widening end regions spring open to their set configuration while the malleable shape memory central neck region remains constrained at or near its crimped dimensions, e.g., in the manner illustrated in FIG. 3B, where bridges 240 disposed across neck region 104 (designated "B" and corresponding to second component 120) engage an opening within the body.Depending on the desired direction of blood flow through the shunt-graft assembly 210, the first diverging end region 102 and the second diverging end region 106 (designated "A" or "C" and corresponding to the first component 110 or the third component 130) provide an inlet and the other of the diverging ends (designated "C" or "A" and corresponding to the third component 130 or the first component 110) provides an outlet. For example, the bridge 240 may engage an opening created through the fossa ovalis of the interatrial septum between the right and left atria, with one of the diverging ends extending into the right atrium and the other diverging end extending into the left atrium. In some configurations, the diverging end in the right atrium is the inlet and the diverging end in the left atrium is the outlet, while in other configurations, the diverging end in the left atrium is the inlet and the diverging end in the right atrium is the outlet. As used herein, "inlet" refers to a component with inflow of blood flow, and "outlet" refers to a component with outflow of blood flow. The specific components that can be used to provide inflow and outflow of blood flow, respectively, may be selected based on the condition being treated. For example, in HF, the inlet may be on the left atrium (LA) side, and blood flow from LA to right atrium (RA) and LA decompression are desired. In contrast, in PAH, the interatrial pressure gradient is reversed, causing R to L flow and RA decompression, and the inlet is on the RA side.

[0078] The cross-sectional area (and size) of the orifice provided by the malleable shape memory central constriction region may be increased or decreased to regulate the flow of fluid through the shunt-graft assembly 210. For example, in the manner illustrated in FIG. 3C, the constriction region may be expanded by balloon expansion using a balloon 301 (e.g., a 12 mm diameter balloon), which may be pumped through the orifice using a wire 302. Preferably, the balloon 301 expands the constriction region only to a threshold outer diameter defined by the bridges 240, such that expansion of the constriction region does not affect the outer diameter of the device or disturb the septal tissue surrounding the bridges 240. For example, the bridges 240 may be sized and shaped to define an outer diameter of 7-14 mm, and the balloon 301 may be configured to expand the constriction region to 9 mm.

[0079] Additionally, in the manner illustrated in FIG. 3D and described above, the necked-down region may be contracted by injecting a bolus of hot saline through the distal end of catheter 303, the saline having a temperature above the Af of the malleable shape memory material (e.g., at 45-65° C.), which may cause the necked-down region to return to its heat-set dimension, which may differ from its crimped dimension, preferably 4.5-5.5 mm. FIG. 3D illustrates one method of heating the necked-down region, but other methods may also be used. For example, saline may be injected through a side hole in catheter 303. Additionally, a second balloon catheter may be inserted through shunt-graft assembly 210 such that the balloon is distal to where saline is delivered and the balloon may be expanded to block blood flow through shunt-graft assembly 210 during delivery of saline. In other embodiments, a pair of electrodes may be positioned in contact with the shunt-graft assembly, for example, via catheter 303, and operated at an appropriate voltage and frequency to heat the necked-down region to or above its Af. In still other embodiments, other suitable means of locally applying heat to device 100 may be used, such as a laser, magnetic inductance, electrical resistance, or the like. Preferably, the return of the necked-down region to its heat-set dimension does not affect the outer diameter of bridge 240, as illustrated in FIG. 3E.

[0080] For example, heat from the saline may transition the malleable shape memory material to the austenitic phase and cause the necked down region to shrink back to its crimped (or otherwise heat set) dimension, followed by cooling the necked down region to body temperature and transitioning back to its martensite phase. The saline may be delivered in any suitable manner, such as by a flexible catheter having one or more openings (e.g., an end hole, a side hole, or multiple side holes) through which the hot saline may flow and which may be placed through the necked down region, for example, over a guidewire. In one non-limiting example, the necked down region may initially have its crimped inner dimension, typically 1-2 mm, a first time, such as when deployed in the manner illustrated in FIG. 3B. The necked down region may then be expanded a second time using balloon expansion to any desired larger dimension between the crimped dimension and the outer diameter of the bridge 240. The neck region may then be shrunk a third time using hot saline to its heat set dimension. The heat set dimension is determined by the size of the jig used in the heat set step during manufacture and may be approximately the same as, smaller than, or larger than the dimension of the neck region in the crimped state. The heat set dimension may be greater than the dimension of the catheter used to deliver the hot saline and greater than the reduced dimension of the dilated balloon, but less than or equal to the smallest anticipated desired final shunt dimension, e.g., 4 mm. The neck region may then be expanded again a second time using balloon expansion to any larger dimension desired between 4 mm and the outer diameter of the bridge 240. Any suitable number of expansions and contractions may be applied to the neck region at any desired time or at times separate from one another to provide suitable and customized fluid flow through the device for each given patient.

[0081] It should be understood that what constitutes a suitable fluid flow for a given patient may also change over time, and the device may be suitably adjusted to provide that fluid flow or to suitably secure the device within the lumen as appropriate. It should also be understood that the self-expanding superelastic components are not affected by the injection of hot saline, and thus will retain their initial fully expanded dimensions while the shape memory components (in this example, the neck region) are being adjusted. Additionally, any suitable method for heating the shape memory material, such as the use of RF heating or lasers, magnetic inductance, electrical resistance, or the like, may be used in the manner as described with reference to FIG. 1A, besides or in addition to hot saline.

[0082] 4A-D, a method of encapsulating a shunt within a biocompatible material is described and further described in U.S. Pat. No. 11,304,831. Preferably, the shunt 100 is at least partially coated with a biocompatible material to produce a shunt-graft assembly 210. For example, the shunt 100 may be coated with a single tube of biocompatible material as shown in FIG. 4A to produce a shunt-graft assembly 210 having various layers of biocompatible material (e.g., two to three layers of biocompatible material). A tube of biocompatible material, the graft tube 216, has a first graft portion 211, a second graft portion 212, and a third graft portion 213. The graft tube 216 has a length greater than the length of the shunt 100, and preferably greater than twice the length of the shunt 100. The first graft portion 211 begins at a first end 214 of the graft tube 216 and extends to the second graft portion 212. The second graft portion 212 extends between and is continuously joined to the first graft portion 211 and the third graft portion 213. The third graft portion 213 terminates at a second end 215 of the graft tube 216. Figures 5A-10B illustrate an exemplary approach for depositing the graft tube 216 onto the shunt 100 in the configuration shown in Figures 4A-D.

[0083] Cross sections (cross sections B, C, and D) of the shunt-graft assembly 210 illustrated in FIG. 4A are illustrated in FIGS. 4B, 4C, and 4D, respectively. The thickness of the graft material is increased in the cross sections to better illustrate the different layers of graft material across the shunt 100. As would be understood by one skilled in the art, the different cross sections of the diverging end region and the constriction region may be greater than or less than the cross sections illustrated in FIGS. 4B-4D. Preferably, each graft layer is less than the thickness of the shunt 100, such that the encapsulated shunt 100 only minimally increases the dimensions of the diverging end region and the constriction region. Now referring to FIG. 4B, cross section B of the shunt-graft assembly 210 is illustrated. As shown in FIG. 4B, the first diverging end region 102 of the shunt 100 may be covered on the outer and inner surfaces by the graft tube 216. Specifically, the first diverging end region 102 may be covered on an outer surface by the third graft portion 213 and on an inner surface (i.e., on the interior of the first diverging end region 102) by the second graft portion 212. Thus, the shunt 100 may be covered at the first diverging end region 102 by two layers of biocompatible material.

[0084] 4C, cross section C of the shunt-graft assembly 210 is illustrated. As shown in FIG. 4C, the necking down region 104 of the shunt 100 may be covered by two layers of biocompatible material on the outer surface of the necking down region 104 and one layer of biocompatible material on its inner surface. Specifically, the necking down region 104 may first be covered on the outer surface by the first graft portion 211, which is covered by the third graft portion 213. On the inner surface (i.e., on the interior of the necking down region 104), the shunt 100 may be covered by the second graft portion 212. Thus, the shunt 100 of the shunt-graft assembly 210 may be covered by three layers of biocompatible material at the necking down region 104.

[0085] 4D, cross-section D of the shunt-graft assembly 210 is illustrated. As shown in FIG. 4D, the second diverging end region 106 of the shunt 100 is covered on the outer surface and the inner surface by the graft tube 216. Specifically, the second diverging end region 106 may be covered on the outer surface by the first graft portion 211 and on the inner surface (i.e., on the interior of the second diverging end region 106) by the second graft portion 212. Thus, the shunt 100 of the shunt-graft assembly 210 may be covered by two layers of biocompatible material at the second diverging end region 106.

[0086] The layers of biocompatible material may be securely bonded together to form a monolithic layer of biocompatible material. For example, first graft portion 211, second graft portion 212, and third graft portion 213 may be sintered together to form a strong, smooth, substantially continuous coating that covers the inner and outer surfaces of the stent. Portions of the coating may then be removed from selected portions of the stent as desired, for example, using laser cutting or mechanical cutting.

[0087] 5A-10D generally illustrate an exemplary method of making a shunt-graft assembly 210 as depicted in FIGS. 4A-4D. Referring now to FIG. 5A, the process may begin by crimping the shunt 100 to deposit a first graft portion onto the neckdown region and the second diverging end region. For example, the shunt 100 may be placed into a funnel 207 and advanced within the funnel 207 toward a reduced section of the funnel 207. The reduced section is preferably the diameter of the neckdown region of the stent or slightly larger. However, it should be understood that the shunt 100 may be reduced to a different diameter. The shunt 100 may be advanced by a dedicated pusher tool such as that described in U.S. Pat. No. 9,713,696 to Yacoby. The shunt 100 may be constructed in a manner such that, in response to the reduction caused by the funnel 207, the shape of the shunt 100 changes such that the first and second gradually expanding end regions become tapered and ultimately directed inwardly toward the longitudinal axis of the shunt 100, resulting in the shunt 100 having a substantially reduced cross-sectional diameter. It should be understood that the shunt 100 may alternatively be compressed into a compressed state using any well-known compression or crimping technique.

[0088] As shown in Figures 5B and 5C, the funnel 207 may have or be coupled to an introducer tube 208 extending from a narrow side of the funnel 207 that may receive the shunt 100 after the shunt 100 is fully restricted by the funnel 207. The introducer tube 208 may have a diameter smaller than that of the graft tube. Alternatively, the first end of the graft tube may be expanded to a diameter larger than that of the introducer tube 208 using well-known expansion techniques (e.g., applying heat to the graft tube). The introducer tube 208 may thus be inserted into the first end 214 of the graft tube 216, as illustrated in Figure 5C, and the shunt 100, having a reduced diameter, may be advanced out of the introducer tube 208 and partially into the graft tube 216 such that the second gradually expanding end region 106 and the reduced diameter region 104 are advanced into the graft tube 216.

[0089] 5D, the second diverging end region 106 and the narrowing region 104 of the shunt 100 are illustrated after being advanced from the introducer tube 208 into the graft tube 216, thereby releasing the compressive force on the second diverging end region 106 and the narrowing region 104, if applicable. Upon removal of the inward radial force from the introducer tube 208, at least the second diverging end region 106 of the shunt 100 may radially expand to a diameter greater than the diameter of the graft tube 216, thereby engaging the graft tube 216 along the outer surface of the second diverging end region 106 and the narrowing region 104 in a manner that causes the graft tube 216 to expand in an unstressed, wrinkle-free manner. Heat may be applied to both the graft tube 216 and the shunt 100 to increase the elasticity of the graft tube 216 and allow the shunt 100 to expand to its expanded state. For example, heated air may be directed at the graft tube 216 and shunt 100. In an alternative approach, the shunt 100 may be cooled below its martensite-austenite transformation temperature so that it becomes martensite. The graft tube 216 may be loaded onto the second diverging end region 106 and the necked-down region 104 in this contracted state and allowed to slowly expand as the shunt 100 warms to room or elevated temperatures in a manner that causes the graft tube 216 to expand in an unstressed, wrinkle-free manner.

[0090] 5E, after depositing the first graft portion 211 of the graft tube 216 over the second diverging end region 106 and the necked down region 104, the shunt 100 may be fully ejected from the introducer tube 208, thereby relieving any compressive forces on the first diverging end region 102. In response to being ejected from the introducer tube 208 and / or in response to the expandable object, the shunt 100 may expand and return to the expanded state illustrated in FIG.

[0091] The graft tube 216 may be cut or otherwise manufactured to the required length to extend along the shunt 100, beginning at the outer surface of the narrowing region 104 adjacent the first diverging end region 102, along the outer surfaces of the narrowing region 104 and the second diverging end region 106, along the inner surface of the shunt 100, and across the outer surfaces of the first diverging end region 102 and the narrowing region 104, terminating at the narrowing region adjacent the second diverging end region 106. Alternatively, the graft tube 216 may be longer than desired and cut using well known cutting techniques (e.g., microscissors, a material cutting guillotine, or a laser cutting machine) to achieve the desired length after the approach described with respect to Figures 5A-10D has been performed.

[0092] 6A, to deposit the graft tube 216 along the inner surface of the shunt 100, the graft tube 216 may be guided through the interior of the shunt as shown in FIG. 6A. As shown in FIG. 6A, the graft tube 216 may be everted at one end and guided through the interior of the shunt 100 such that the everted portion of the graft tube 216 is positioned within the interior of the shunt 100. To facilitate this process, a plunger 217, which may be any tool having a long shaft and a width or diameter less than the inner diameter of the necked down region 104 in the expanded state, may be used to push the graft tube 216 through the interior of the shunt 100. In this manner, the first graft portion 211 may extend along the outer surface of the second diverging end region 106, curve around the end of the second diverging end region 106, and proceed along the interior of the shunt 100 and out the opposite side of the shunt 100.

[0093] 6B, a first mandrel portion 218 may be engaged with the shunt 100 and the second graft portion 212 to engage the remaining portion of the second graft portion 212 with the inner surface of the shunt 100. Having an end shape formed to correspond to the interior of the first diverging end region 102, the first mandrel portion 218 may be gently advanced within the second end 215 of the graft tube 216 and the shunt 100 until the first mandrel portion 218 occupies substantially the entire space within the first diverging end region 102. The process may involve simultaneously guiding the first diverging end region 102 over the first mandrel portion 218 such that the second end 215 of the graft tube 216 extends along the first mandrel portion 218 in an interference-fitting, wrinkle-free manner. When the first diverging end region 102 is properly mounted on the first mandrel portion 218, the second end 215 of the graft tube 216 will extend beyond the first diverging end region 102, as shown in FIG. 6B. In this manner, the second graft portion 212 may be partially engaged with the shunt 100 along an inner surface of the shunt 100.

[0094] 7A and 7B, a first mandrel portion 218 (FIG. 7A) and a second mandrel portion 219 (FIG. 7B) are illustrated. As illustrated in FIGS. 8A and 8B, the first mandrel portion 218 and the second mandrel portion 219 may be removably coupled to form a mandrel assembly 220. Referring now to FIG. 7A, a side view and a front view of the first mandrel portion 218 are illustrated. The first mandrel portion 218 may have a first retention portion 222 and a first body portion 223, the first retention portion 222 extending from the first body portion 223. The first retention portion 222 engages the shunt-graft assembly 210 and is designed to have a shape similar to the first diverging end region and / or the constriction region, but with a slightly smaller dimension. The first body portion 223 may have a cylindrical shape. It should be understood that the first mandrel portion may alternatively have only the first retaining portion 222. The first mandrel portion 218 may have a receiving portion 221 that is sized and configured to receive the protruding portion 228 of the second mandrel portion 219. The receiving portion 221 may extend the entire length of the first mandrel portion 218 or alternatively may extend only a portion of the first mandrel portion 218.

[0095] First mandrel portion 218 may also optionally have one or more ventilation holes 224 in first retention portion 222. Ventilation holes 224 may extend through the exterior surface of first retention portion 222 and through the interior of first retention portion 222 and first body portion 223 to ventilation inlets 225, which may extend through the surface of first body portion 223. Ventilation holes 224 are preferably in the range of 0.1-2 mm in size, although it should be understood that ventilation holes of different sizes may also be beneficial. Ventilation holes 224 may facilitate the release of stent-graft assembly 210 after a heat treatment is applied, as discussed below with respect to FIGS. 10A and 10B. Specifically, after the encapsulated stent is tightly compressed against mandrel assembly 220 and the air between the encapsulated stent and mandrel assembly 220 is released, suction forces may be present that make it difficult to remove the encapsulated stent. The ventilation inlet 225 may allow air to flow through the ventilation inlet to the ventilation holes 224, eliminating or reducing the suction effect. It should be appreciated that multiple ventilation holes may be in communication with one or more ventilation inlets.

[0096] 7B, the second mandrel portion 219 is illustrated. The second mandrel portion 219 may have a second retention portion 226 and a second body portion 227, with the second retention portion 226 extending from the second body portion 227. The second retention portion 226 is designed to engage the shunt-graft assembly 210 and may have a shape similar to the second flared end region and / or the constriction region, but with a slightly smaller dimension. The second body portion 227 may have a cylindrical shape. It should be understood that the second mandrel portion may alternatively have only the second retention portion 226.

[0097] The second mandrel portion 219 has a protruding portion 228 that is sized and shaped to be received by the receiving portion 221 of the first mandrel portion 218. The protruding portion 228 may be, for example, a shaft extending from the second retaining portion 226. The protruding portion may be coaxial with the second mandrel portion 219 and may be designed to extend a portion of, the entire length of, or more than the length of the first mandrel portion 218. Like the first mandrel portion 218, the second mandrel portion 219 may optionally include one or more ventilation holes 229 and one or more ventilation inlets 230.

[0098] 8A , to restrain the shunt-graft assembly 210 on the mandrel assembly 220, the second mandrel portion 219 is removably coupled to the first mandrel portion 218 by engaging the protruding portion 228 with the receiving portion 221. The second mandrel portion 219 may be gently advanced toward the first mandrel portion 218 within the second diverging end region 106 of the shunt-graft assembly 210 until the second mandrel portion 219 occupies substantially the entire space within the second diverging end region 106 and the protruding portion 228 is fully received by the receiving portion 221. In this manner, the second graft portion 212 may be fully engaged with the second diverging end region 106.

[0099] The protruding portion 228 may be designed to engage with the receiving portion 221 such that the protruding portion 228 and the engagement portion are releasably locked together. Alternatively, the protruding portion 228 may be designed to frictionally fit within the receiving portion 221. For example, the protruding portion may be designed with a gradually increasing diameter that may provide a frictional fit with the receiving portion 221. In this example, the first mandrel portion 218 and the second mandrel portion 219 may be released by forcing the first mandrel portion 218 and the second mandrel portion 219 apart. It should be understood that the first mandrel portion 218 and the second mandrel portion 219 may be releasably locked together or otherwise frictionally fit together using a variety of other well-known techniques. It should further be appreciated that the protruding portion 228 could instead extend from the first mandrel portion 218 and the receiving portion 221 could instead be formed within the second mandrel portion 219 .

[0100] 8A, after engaging the second mandrel portion 219 with the first mandrel portion 218, thus restraining the shunt-graft assembly 210 on the mandrel assembly 220, the third graft portion 213 may be separated from the surface of the first mandrel portion 218. For example, forceps may be used to grasp the second end 215 of the graft tube 216 and gently pull the second end 215 over the first diverging end region 102 and over the necking down region 104. In this manner, the graft tube 216 may be everted and guided over the first diverging end region 102 and over the necking down region 104. Alternatively, other well-known techniques may be used to separate the third graft portion 213 from the first mandrel portion 218 and deposit the third graft portion 213 over the first diverging end region 102 and the narrowing region 104.

[0101] 8B, the third graft portion 213 may be gently compressed against the first diverging end region 102 and the constriction region 104 in a manner that reduces or eliminates wrinkles. For example, the second end 215 may be manually stretched toward the constriction region 104 using forceps and allowed to gently contact the portions of the first graft portion 211 that contact the first diverging end region 102 and the constriction region 104.

[0102] By placing the third graft portion 213 over the first diverging end region 102 and the necking down region 104, the graft tube 216 will be deposited over the shunt 100 such that the graft tube 216 covers the shunt 100 in the manner depicted in Figures 4A-4D. As shown in Figure 8C, depositing the graft tube 216 over the shunt 100 in the manner described above results in three biocompatible layers of graft material covering the necking down region 104 of the shunt 100. Specifically, as shown in Figure 8C, the necking down region 104 of the shunt 100 is covered on the inner surface by the second graft portion 212 and on the outer surface by the first graft portion 211 and the third graft portion 213. In Figure 8C, the third graft portion 213 overlaps the first graft portion 211 at the necking down region 104. 4B and 4D, the first and second diverging end regions 102, 106 of the shunt 100 may be covered by only two layers of biocompatible material. Thus, the second graft portion 212 may extend through the lumen of the shunt 100, through the first end of the first diverging end region 102, through the narrowing region 104, to the end of the second diverging end region 106, with the first graft portion 211 extending along the outer surface of the second diverging end region 106 and narrowing region 104, and the third graft portion 213 extending along the outer surface of the first diverging end region 102 and narrowing region 104, at least partially overlapping and joining with the first graft portion 211.

[0103] It should be appreciated that the graft tube 216 may be deposited onto the shunt 100 using approaches different from those detailed in Figures 5A-8C to form a shunt-graft assembly 210 having the same three-layer structure in the narrowing region 104 and two-layer structure in the first and second narrowing end regions 102, 106. It should be further appreciated that a three-layer structure may be deposited onto the shunt 100 in a similar manner, but with the result that the three-layer structure occurs across the first and second narrowing end regions 102, 106, and / or narrowing region 104. In another example, a similar approach may be used to produce a shunt-graft assembly having one region with four layers of biocompatible material and another region with two layers of biocompatible material. These and other methods and embodiments are further described in U.S. Pat. No. 11,304,831.

[0104] 9A-9D, to securely bond the first graft portion 211, the second graft portion 212, and the third graft portion 213 to the shunt 100 and to each other, pressure and heat may be applied to the shunt-graft assembly 210 to achieve sintering. As explained above, sintering results in a strong, smooth, and generally continuous coating covering the inner and outer surfaces of the stent. Sintering may be achieved by first covering the shunt-graft assembly 210 with a flexible sleeve (e.g., flexible clamshell 231 shown in FIG. 9A), applying a compressor 232 (shown in FIG. 10A), and / or applying heat. In this manner, the first graft portion, the second graft portion, and the third graft portion may be bonded to each other through the through-wall openings in the shunt 100. Thus, the third graft portion and the second graft portion may be sintered together and joined.

[0105] The flexible sleeve may be tubular and may be elastic and biocompatible. For example, the flexible sleeve may be a flexible clamshell 231 as illustrated in FIG. 9A. The flexible clamshell 231 is preferably made of biocompatible silicone, but may alternatively be other biocompatible materials with elastic properties. The flexible clamshell 231 may be hollow and have a consistent thickness. Alternatively, the flexible clamshell 231 may have various thicknesses at different points or within different sections. The flexible clamshell 231 has a first end 235 and a second end 236 that may come together in a neutral position. Alternatively, there may be a vertical gap between the first end 235 and the second end 236 in the neutral position. Flexible clamshell 231 is designed such that first end 235 and second end 236 can be separated by pulling first end 235 and second end 236 in opposite directions.

[0106] 9B, first end 235 and second end 236 may be pulled in opposite directions to create a longitudinal gap such that first end 235 and second end 236 may be pulled over shunt-graft assembly 210 while shunt-graft assembly 210 is positioned on mandrel assembly 220. In this manner, flexible clamshell 231 either entirely or nearly entirely covers shunt-graft assembly 210. Flexible clamshell 231 may be sized such that it fits tightly over shunt-graft assembly 210, as illustrated in FIG. It should be appreciated that a thin metal layer may be deposited over the shunt-graft assembly 210 prior to coating the shunt-graft assembly 210 in the clamshell 231 such that the thin metal layer may be sandwiched between the shunt-graft assembly 210 and the flexible clamshell 231 and may act as a barrier between the shunt-graft assembly 210 and the flexible clamshell 231. This barrier may help prevent contaminants from the clamshell 231 from being transferred to the shunt-graft assembly 210. In one embodiment, the thin metal layer may be aluminum foil. In other embodiments, the thin metal layer may be titanium, tantalum, or stainless steel, although it should be appreciated that other metals or alloys may also be used. This process has been observed to enhance the compaction and sintering process. If the stent-graft assembly is wrapped with a tape, such as a TFE or ePTFE tape, the tape may similarly prevent contaminants from being transferred to the shunt-graft assembly.

[0107] The flexible clamshell 231 may be sized such that when positioned over the shunt-graft assembly 210, the flexible clamshell 231 applies a compressive force against the shunt-graft assembly 210. The flexible clamshell 231 may be sized and configured to optimize the fit of the graft tube 216 to the shunt 100 and minimize gaps between layers of the graft tube 216 adjacent the struts of the shunt 100. The degree of pressure that the flexible clamshell 231 applies to the shunt-graft assembly 210 may modify the internodal distance (IND) of the graft material once sintered, as described in more detail below. The extent to which the flexible clamshell 231 covers or does not cover the shunt-graft assembly 210 may also modify the internodal distance. It should be understood that the internodal distance is related to tissue ingrowth and the compressive force applied by the flexible clamshell 231 may be modified to achieve a desired internodal distance. Alternatively, any compressive force applied by flexible clamshell 231 may be negligible. Additional compressive force on shunt-graft assembly 210 may optionally be achieved by first wrapping shunt-graft assembly 210 and / or flexible clamshell 231 with a tape, such as a TFE or ePTFE tape. For example, shunt-graft assembly 210 covered with flexible clamshell 231 may be placed into a spiral winding wrapping machine that tension-wraps shunt-graft assembly 210 and flexible clamshell 231 with at least one overlapping layer of tape, as described in more detail above.

[0108] 10A and 10B, a compressor 232 is illustrated. The compressor 232 has a first half 233 and a second half 234, and further includes a coupler 237 for removably coupling the first half 233 to the second half 234. The coupler 237 may involve a screw with a locking nut or any other well-known technique for removably locking two components together. The first half 233 and the second half 234 may include a receiving portion 238 that is sized and configured to receive the coupler 237. First half 233 and second half 234 have internal indentations configured to receive flexible clamshell covered shunt-graft assembly 210 such that when removably coupled together, the internal indentations of each of first half 233 and second half 234 have the shape of shunt-graft assembly 210 covered by flexible clamshell 231. Alternatively, when removably coupled together, the internal indentations of each of first half 233 and second half 234 have the shape of mandrel assembly 220, except with a slightly larger radial dimension.

[0109] The first and second halves 233, 234 are rigid and preferably stainless steel, although it should be understood that the first and second halves 233, 234 can be other rigid materials. The first and second halves 233, 234 may be designed such that when the first and second halves 233, 234 are bonded together, the first and second halves 233, 234 are positioned a fixed distance from the shunt-graft assembly 210. Alternatively, the distance from the shunt-graft assembly 210 or mandrel assembly 220 may vary in different regions of the first and second halves 233, 234. The first and second halves 233, 234 may be designed with a wall thickness between the interior and exterior surfaces of the first and second halves 233, 234 that allows for a desired degree of heat transfer. For example, the first and second halves 233, 234 may have a thin wall thickness to increase the amount of heat transfer to the stent-graft assembly. Thinner walls may result in shorter sintering times, which may improve production rates. Additionally, shorter sintering times reduce the effect of sintering on the transformation temperature (e.g., austenite finish (Af)) of the Nitinol frame.

[0110] 10B, compressor 232 may be attached to a clamshell 231 that covers shunt-graft assembly 210 while shunt-graft assembly 210 is loaded onto mandrel assembly 220. Specifically, first half 233 may be positioned over shunt-graft assembly 210 and a first half portion of clamshell 231, and second half 234 may be positioned over clamshell 231 and a second half portion of shunt-graft assembly 210, such that first half 233 and second half 234 entirely or nearly entirely cover flexible clamshell 231. Once the receiving portion 238 of the first half 233 and second half 234 are aligned, the coupler 237 may be inserted into the receiving portion 238 to removably couple the first half 233 to the second half 234 across the clamshell 231. The coupler 237 may be tightened to increase the compressive force applied to the shunt-graft assembly 210. It is understood that the compressive force applied by the coupler 237 may modify the internodal distance of the graft material once sintered. It is further understood that the coupler 237 may be tightened to a certain extent to achieve a desired internodal distance.

[0111] In response to coupling the first half 233 to the second half 234 around the flexible clamshell 231, the compressor 232 will be positioned over the flexible clamshell 231, which will be positioned over the shunt-graft assembly 210, which will be positioned over the mandrel assembly 220, forming a sintered assembly 239, as illustrated in FIG. 10B. In response to coupling the first half 233 to the second half 234, the compressor 232 applies a compressive force to the flexible clamshell 231 and the shunt-graft assembly 210, thereby compressing the shunt-graft assembly 210 against the mandrel assembly 220. The flexible clamshell 231 may facilitate an even distribution of the compressive force to the shunt-graft assembly 210 due to its elastic nature. The shunt 100 may include multiple through-wall openings. The force applied by the wrapping tape and / or compressor 232 compresses the shunt-graft assembly 210 against the mandrel assembly 220, thereby forcing the graft layers into intimate contact through interstices of the shunt 100.

[0112] It may be desirable for the compressor 232 to apply a uniform compressive force. Alternatively, it may be desirable to vary the compressive force applied to the shunt-graft assembly 210 at certain points along the shunt-graft assembly 210. For example, the flexible clamshell 231 may have a varying thickness and / or length, allowing the compressor 232 to distribute varying degrees of compressive force to the shunt-graft assembly according to the wall thickness and geometry of the flexible clamshell 231. In addition, the distance from the interior wall to the surface of the shunt-graft assembly 210 may vary at certain points along the first half 233 and / or second half 234. For example, an area of ​​the interior wall of the first half 233 that has a distance to the shunt-graft assembly 210 less than the remainder of the first half 233 may apply more compressive force to the shunt-graft assembly 210. Varying the compressive force applied to the shunt-graft assembly 210 may reduce or increase the compatibility between the first graft portion 211, the second graft portion 212, and the third graft portion 213. In an alternative embodiment, the compressor 232 may be designed such that it applies a compressive force only at the neckdown region 104.

[0113] To form a monolithic layer of biocompatible material, the first graft portion 211, the second graft portion 212, and the third graft portion 213 of the graft tube 216 may be securely bonded together by applying heat to the sintering assembly 239. For example, the sintering assembly 239 may be heated by placing the sintering assembly 239 in a radiant heat furnace, which may be preheated. Sintering may be performed as discussed in more detail in U.S. Pat. No. 11,304,831. The heated assembly may then be allowed to cool for a period of time sufficient to permit manual handling of the assembly. After cooling, the first and second halves 233, 234 of the compressor 232 may be uncoupled and removed from the flexible clamshell 231. The spiral wrap may then be unwound and discarded, if applicable. The flexible clamshell 231 may be removed, and the encapsulated stent may then be rotated concentrically about the axis of the mandrel to release any adhesion between the second graft portion 212 and the mandrel assembly 220. The encapsulated stent, still on the mandrel assembly 220, may then be placed into a laser trimming fixture to trim away excess graft material, if applicable. Additionally, the graft material of the encapsulated stent may be trimmed at various locations along the stent, such as near one of the stent ends, to allow for coupling to a delivery device, as shown in FIG. 1C.

[0114] The resulting structure produced using the approach shown in Figures 4A-4D and described with respect to Figures 5A-10B is beneficial in that the biocompatible material does not terminate at both ends of the shunt-graft assembly 210, but instead terminates at the necked-down region 104 of the shunt 100. The inventors have discovered that biocompatible material extending beyond the stent at both ends is known to result in thrombus formation when implanted and can also cause interference with attachment to the delivery catheter. The reduction of graft overhang at the edges also improves the fluid dynamics at the inlet and outlet of the encapsulated stent. As an added benefit, the inventors have observed that the resulting triple layer structure at the necked-down region 104 further inhibits tissue ingrowth. It will be appreciated that the foregoing method described with respect to Figures 5A-10B helps to prevent gaps between biocompatible layers in the flared end regions that can result in extensive tissue ingrowth, and also helps to minimize microscopic surface thinning defects (MSTDs) that can result in attracting platelet thrombi. Applicants have appreciated that the approach described with respect to Figures 5A-10B provides a high yield and highly reproducible manufacturing process.

[0115] Applicants have further observed that heating the sintering assembly 239, which includes a flexible clamshell of silicone, as described herein, results in small fragments and / or molecular moieties of silicone becoming deposited on and / or impregnated within the graft tube 216. Applicants have observed that the fragments and / or molecular moieties of silicone deposited on and / or impregnated within the graft tube 216 may further reduce tissue ingrowth when the encapsulated stent is implanted.

[0116] 11A-11D, a first embodiment of the shunt-graft assembly of FIGS. 4A-4D is illustrated, further including a bridge as described above. The shunt 100 may be covered with a single tube of biocompatible material, as shown in FIG. 4A, to produce a shunt-graft assembly 210 having various layers of biocompatible material (e.g., 2-3 layers of biocompatible material). For example, the graft tube 216 may be deposited over the shunt 100 such that the second graft portion 212 extends through the lumen of the shunt 100, through a first end of the first diverging end region 102, through the necking down region 104, to the end of the second diverging end region 106, the first graft portion 211 extends along the outer surface of the second diverging end region 106 and the necking down region 104, and the third graft portion 213 extends along the outer surface of the first diverging end region 102 and the necking down region 104 and at least partially overlaps and joins with the first graft portion 211. After depositing the graft tube 216 onto the shunt 100, a bridge 240 may be deposited onto the shunt-graft assembly 210.

[0117] The bridge 240 may be a single tube of biocompatible material similar to the material of the graft tube 216. As described with reference to FIG. 1B, the biocompatible material of the bridge 240 is preferably one that encourages tissue attachment so that contact with the septal wall is maintained when the inner diameter of the shunt is reduced. The biocompatible material may be configured to encourage tissue ingrowth over the entire bridge or only over a portion of the bridge. For example, holes may be placed at the locations of the bridge 240 that are configured to engage the atrial septum, while the remaining portions of the bridge 240 remain intact such that tissue ingrowth is not encouraged on the flared end regions. Such addition of holes or other treatments to encourage tissue ingrowth or to encourage attachment may be performed on the biocompatible material prior to or after the shunt is assembled. Additionally or alternatively, the bridge 240 may be made of a different biocompatible material than the biocompatible material used to encapsulate the shunt. For example, the shunt may be encapsulated with a biocompatible material such as ePTFE having a pore size small enough so that tissue ingrowth is mitigated, and the bridge may be made of a biocompatible material with a larger pore size designed to encourage tissue ingrowth. Alternatively, the bridge 240 may be made from woven Dacron, a mesh-like structure, an electrospun fabric, or silicone to further encourage tissue ingrowth. As further described above, if woven Dacron is used to form the bridge, the Dacron may be securely attached to the encapsulated shunt using stitches rather than the method described below. In some embodiments, the shunt may be encapsulated with ePTFE having a thickness of 0.002 inches and an internodal distance of ≦30 microns.Clowes et al. "Rapid Transmural Capillary Ingrowth Provides a Source of Intimal Endothelium and Smooth Muscle in Porous PTFE Prostheses," Arterial Graft Failure, Vol. 123, No. 2, pp. 220-230 (May 1986) describes ePTFE with an IND≦30 microns exhibiting low porosity, and applicants' studies have shown that tissue ingrowth is inhibited in shunts encapsulated with ePTFE with an IND=30 microns. On the other hand, the bridge 240 may have a thickness of 0.002 inches or 0.005 inches and a pore size to enhance transmural infiltration of cells and collagen into the potential space between the bridge 240 and other layers of ePTFE. This can be done using larger pore sizes of ePTFE (e.g., with ePTFE material having an IND ranging from 60 to 200 microns). Alternatively, transmural infiltration may be encouraged by creating a similarly sized pattern or multiple perforations in the bridge 240 machined from a conventional low porosity (IND≦30 microns) either before or after its application to the shunt. Such a dedicated perforation process may be performed using an energy source, such as a laser, RF, or a mechanical source, such as a perforator, or any other technique known to those skilled in the art of thin material processing. In some embodiments, the frame encapsulation material is intended to block tissue ingrowth, while the bridge encapsulation material will be more elastic to support significant expansion / contraction in diameter without damaging the fossa ovalis or the frame encapsulation material. In some embodiments, the gap 243 between the bridge 240 and the constriction region 104 increases as the shunt constriction region 104 is constricted. The bridge 240 may be configured to remain engaged with the patient's atrial septum when the constriction region is constricted.In some embodiments, the biocompatible material of the bridge 240 has a porosity (e.g., as measured by its internodal distance) that exceeds the porosity of the biocompatible material of the encapsulation of the shunt frame. Thus, the internodal distance of the bridge material may be selected to allow tissue ingrowth, while the internodal distance of the encapsulation material is selected to prevent tissue ingrowth. In some embodiments, the internodal distance of the bridge is greater than 30 microns (e.g., in the range of 45-200 microns), while the internodal distance of the encapsulation is less than or equal to 30 microns. In one embodiment, the internodal distance of the bridge is 60 microns, while the internodal distance of the encapsulation is 30 microns. The biocompatible material of the bridge and the biocompatible material of the encapsulation may be expanded polytetrafluoroethylene (ePTFE).

[0118] The bridge 240 may have a length that is shorter than the shunt-graft assembly 210 and a diameter that is greater than the diameter of the narrowing region 104. Preferably, the bridge 240 is shaped and sized such that, after depositing the bridge 240 over the narrowing region 104, a first end 241 extends to approximately the middle of the first diverging end region 102, a second end 242 extends to approximately the middle of the second diverging end region 106, and a gap 243 is created between the narrowing region 104 and the bridge 240.

[0119] Cross sections (cross sections B, C, D, E, and F) of the shunt-graft assembly 210 illustrated in FIG. 11A are illustrated in FIGS. 11B, 11C, 11D, 11E, and 11F, respectively. The thickness of the graft material is exaggerated in the cross sections to better illustrate the different layers of graft material across the shunt 100. As would be understood by one of ordinary skill in the art, the different cross sections of the diverging end region and the constriction region may be greater than or less than the cross sections illustrated in FIGS. 11B-11F. Referring now to FIG. 11B, cross section B of the shunt-graft assembly 210 is illustrated. As shown in FIG. 11B, the first diverging end region 102 of the shunt 100 may be covered on the outer and inner surfaces by a graft tube 216. Specifically, the first diverging end region 102 may be covered on an outer surface by the third graft portion 213 and on an inner surface (i.e., on the interior of the first diverging end region 102) by the second graft portion 212. Thus, the shunt 100 may be covered by two layers of biocompatible material at the first diverging end region 102. From where the first end 241 of the bridge 240 is joined to the third graft portion 213 and the necking down region 104, the first diverging end region 102 may be covered by three layers of biocompatible material, for example, as shown in FIG.

[0120] 11D, cross section D of the shunt-graft assembly 210 is illustrated. As shown in FIG. 11D, the necking down region 104 of the shunt 100 may be covered by three layers of biocompatible material on the outer surface of the necking down region 104 and one layer of biocompatible material on its inner surface. Specifically, the necking down region 104 may first be covered on the outer surface by the first graft portion 211, which is covered by the third graft portion 213. Additionally, the necking down region 104 may be covered by a bridge 240, and a gap 243 may extend between the third graft portion 213 and the bridge 240. On the inner surface (i.e., on the interior of the necking down region 104), the shunt 100 may be covered by the second graft portion 212. Thus, the shunt 100 of the shunt-graft assembly 210 may be covered in the region of narrowing 104 with four layers of biocompatible material.

[0121] 11F, a cross section F of the shunt-graft assembly 210 is illustrated. As shown in FIG. 11F, the second diverging end region 106 of the shunt 100 is covered on the outer surface and the inner surface by the graft tube 216. Specifically, the second diverging end region 106 may be covered on the outer surface by the first graft portion 211 and on the inner surface (i.e., on the interior of the second diverging end region 106) by the second graft portion 212. Thus, the shunt 100 of the shunt-graft assembly 210 may be covered by two layers of biocompatible material at the second diverging end region 106. From where the second end 242 of the bridge 240 is joined to the first graft portion 211 and the constriction region 104, the second diverging end region 106 may be covered by three layers of biocompatible material, for example, as shown in FIG. 11E.

[0122] The bridge 240 may be deposited on the shunt-graft assembly 210 in a manner similar to the method described in Figures 5A-10D. Preferably, the bridge 240 is incorporated into the device after the graft tube 216 is sintered to the shunt 100 to form a monolithic layer of biocompatible material, thereby ensuring that the graft tube 216 remains in place while the bridge 240 is deposited on the shunt-graft assembly 210. For example, the shunt-graft assembly 210 may be placed into a funnel and advanced towards a reduced section of the funnel. The shunt-graft assembly 210 may then be inserted into an introducer tube, which may have a diameter smaller than that of the bridge 240. The introducer tube may then be inserted into a first end 241 of the bridge 240 and advanced partially out of a second end 242. The shunt-graft assembly 210 may then be advanced out of the introducer tube such that the second diverging end 106 extends past the second end 242 and the first diverging end 102 extends in the opposite direction past the first end 241. The bridge 240 may then be positioned across the narrowing region 104 such that the first end 241 is disposed approximately to the middle of the first diverging end region 102 and the second end 242 is disposed approximately to the middle of the second diverging end region 106, with a gap 243 disposed between the narrowing region 104 and the bridge 240. The gap 243 may be created by manually stretching (e.g., using forceps) the first end 241 toward the first diverging end region 102 and the second end 242 toward the second diverging end region 106.

[0123] To securely bond the first end 241 to the third graft portion 213 and the second end 242 to the first graft portion 211, pressure may be applied to the shunt-graft assembly 210 to achieve sintering. Additionally or alternatively, the gap 243 may be filled with a biocompatible material, as shown in FIG. 2B. For example, a biocompatible material may be deposited on the shunt-graft assembly 210 in the necked-down region either before or after the bridge 240 is deposited over the necked-down region 104.

[0124] 12A-12D illustrate another embodiment of the shunt-graft assembly of FIGS. 4A-4D further including a bridge as described above. As described above, the shunt 100 may be encapsulated to produce a shunt-graft assembly 210 having various layers of biocompatible materials, such as two layers across the first and second diverging end regions 102, 106 and three layers across the neck region 104. After depositing the graft tube 216 onto the shunt 100, a bridge 240 may be deposited onto the shunt-graft assembly 210.

[0125] The bridge 240 may be created using a second graft tube 245, which may be a single tube of a biocompatible material similar to the material of the graft tube 216. The second graft tube 245 may have a length greater than twice the length of the shunt 100 and a diameter greater than the diameter of the neckdown region 104. The second graft tube 245 may comprise a fourth graft portion 246, a fifth graft portion 247, and a sixth graft portion 248. Preferably, the fourth graft portion 246 extends across the neckdown region 104 to create the bridge 240.

[0126] Cross sections (cross sections B, C, and D) of the shunt-graft assembly 210 illustrated in FIG. 12A are illustrated in FIGS. 12B, 12C, and 12D, respectively. The thickness of the graft material is increased in the cross sections to better illustrate the different layers of graft material across the shunt 100. As would be understood by one of ordinary skill in the art, the different cross sections of the diverging end region and the constriction region may be greater than or less than the cross sections shown in FIGS. 12B-12D. Referring now to FIG. 12B, cross section B of the shunt-graft assembly 210 is illustrated. As shown in FIG. 12B, the first diverging end region 102 of the shunt 100 may be covered on the outer and inner surfaces by a graft tube 216. Specifically, the first diverging end region 102 may be covered on an outer surface by the third graft portion 213 of the graft tube 216 and on an inner surface by the second graft portion 212 of the graft tube 216 (i.e., on the interior of the first diverging end region 102). Additionally, a second graft tube 245 having fourth, fifth, and sixth graft portions 246, 247, and 248 may be deposited over the graft tube 216. Specifically, the first diverging end region 102 may be covered on an outer surface by a fourth graft portion 246 of the second graft tube 245, which may extend over the third graft portion 213 of the graft tube 216, and on an inner surface by a fifth graft portion 247 of the second graft tube 245, which may extend over the second graft portion 212 of the graft tube 216. In some portions of the first diverging end region 102, the sixth graft portion 248 of the second graft tube 245 may be deposited such that it overlaps the fourth graft portion 246. Thus, the shunt 100 may be covered by four or five layers of biocompatible material in the first diverging end region 102. As will be understood by one of ordinary skill in the art, the fourth graft portion 246 may instead overlap the sixth graft portion 248 on the second diverging end region 106 such that the first diverging end region 102 is covered by only four layers of biocompatible material.

[0127] 12C, cross section C of shunt-graft assembly 210 is illustrated. As shown in FIG. 12C, necking region 104 of shunt 100 may be covered by three layers of biocompatible material on the outer surface of necking region 104 and two layers of biocompatible material on its inner surface. Specifically, necking region 104 may be first covered on the outer surface by first graft portion 211 of graft tube 216, which is covered by third graft portion 213 of graft tube 216. Additionally, necking region 104 may be covered by bridge 240 (which may be a section of fourth graft portion 246 of second graft tube 245), and gap 243 may extend between third graft portion 213 of graft tube 216 and bridge 240. On the inner surface (i.e., on the interior of the necking down region 104), the shunt 100 may be covered by the second graft portion 212 of the graft tube 216 and the fifth graft portion 247 of the second graft tube 245. Thus, the shunt 100 of the shunt-graft assembly 210 may be covered in the necking down region 104 by five layers of biocompatible material.

[0128] 12D, cross-section D of the shunt-graft assembly 210 is illustrated. As shown in FIG. 12D, the second diverging end region 106 of the shunt 100 is covered on the outer surface and on the inner surface by the graft tube 216. Specifically, the second diverging end region 106 may be covered on the outer surface by the first graft portion 211 of the graft tube 216 and on the inner surface (i.e., on the interior of the second diverging end region 106) by the second graft portion 212 of the graft tube 216. Additionally, a second graft tube 245 may be deposited over the graft tube 216. Specifically, the second diverging end region 106 may be covered on an outer surface by a fourth graft portion 246 of the second graft tube 245, which may extend over the first graft portion 211 of the graft tube 216, and on an inner surface by a fifth graft portion 247 of the second graft tube 245, which may extend over the second graft portion 212 of the graft tube 216. Thus, the shunt 100 of the shunt-graft assembly 210 may be covered by four layers of biocompatible material at the second diverging end region 106.

[0129] 5A-10D. For example, the shunt-graft assembly 210 may be placed into a funnel and advanced toward the reduced section of the funnel. The shunt-graft assembly 210 may then be inserted into an introducer tube, which may have a diameter smaller than that of the bridge 240. The introducer tube may then be inserted into a first end of the bridge 240 such that all of the second diverging end region 106 and the constriction region 104 and a portion of the first diverging end region 102 are advanced into the second graft tube 245. Upon removal of the introducer tube, the first diverging end region 102 and the second diverging end region 106 may radially expand to a diameter greater than the diameter of the second graft tube 245, thereby engaging the second graft tube 245 along the outer surface of the diverging end region in a manner that causes the second graft tube 245 to expand in an unstressed and wrinkle-free manner. Preferably, the second graft tube 245 has a diameter greater than the diameter of the necking down region 104 such that a bridge 240 is created with a gap 243 between the encapsulated necking down region 104 and the second graft tube 245. The bridge 240 may have a first end 241 that may be located approximately to the middle of the first diverging end region 102 and a second end 242 that may be located approximately to the middle of the second diverging end region 106.

[0130] After depositing the fourth graft portion 246 of the second graft tube 245 over the second diverging end region 106, the necking region 104, and a portion of the second diverging end region 102, the second graft tube 245 may be everted at one end and guided through the interior of the shunt 100 using a plunger or other tool as described above. A mandrel similar to the first mandrel portion 217 and the second mandrel portion 219 described above and shown in Figures 7A and 7B may be used to ensure that the fifth graft portion 247 is properly engaged with the inner surfaces of the first diverging end region 102, the necking region 104, and the second diverging end region 106.

[0131] After depositing the fifth graft portion 247 of the second graft tube 245 onto the interior of the shunt 100, a forceps may be used to grasp the second end of the second graft tube 245 and gently pull the second end over the first diverging end region 102, thereby depositing the sixth graft portion 248. Preferably, the sixth graft portion 248 is deposited over the first diverging end region 102 such that it partially overlaps the fourth graft portion 246, as shown in FIG.

[0132] After depositing the sixth graft portion 248 of the second graft tube 245 onto the first diverging end region, a flexible sleeve similar to the flexible sleeve 231 described above and shown in FIG. 9A may be positioned over the mandrel. Preferably, the flexible sleeve has a shape configured to receive the shunt-graft assembly 210 including the bridge 240, and the necked-down region of the flexible sleeve has a diameter corresponding to the diameter of the bridge 240 such that the gap 243 is not reduced. A compressor similar to the compressor 232 described above and shown in FIG. 10A may be positioned over the flexible sleeve. Preferably, the compressor has first and second halves, each with internal indentations sized and configured to receive the flexible sleeve.

[0133] To securely bond the second graft tube 245 to the graft tube 216, pressure may be applied to the shunt-graft assembly 210 to effect sintering, as described above. Preferably, heating the sintering assembly causes the fourth graft portion 246, the fifth graft portion 247, and the sixth graft portion 248 to become sintered to the shunt-graft assembly 210, except for the necking down region 104, such that a gap 243 that cannot be bonded is created between the necking down region 104 and the bridge 240. Additionally or alternatively, the gap 243 may be filled with a biocompatible material, as shown in FIG. 2B. For example, a biocompatible material may be deposited on the shunt-graft assembly 210 in the necking down region, either before or after the fourth graft portion 246 of the second graft tube 245 is deposited over the necking down region 104.

[0134] FIG. 13 illustrates another embodiment of a shunt-graft assembly including the bridge described above. In this embodiment, a single tube of biocompatible material may be used to encapsulate the frame and form the bridge. For example, a single tube may be wrapped twice, using three flips, starting at one location, e.g., one end near the end of the neck or one flange of the frame, to lay down two complete encapsulations. In FIG. 13, the first complete encapsulation "first layer" is shown in blue and the second encapsulation "second layer" is shown in pink. In some embodiments, the neck area is sintered prior to the final flip of the second layer. In such an embodiment, after the neck is sintered, there will be three layers of material in the neck, pink and blue on the inside and blue only on the outside, all of which may be sintered together. The final end of the tube may then be flipped back over the shunt. This layer will form the bridge described herein. In some embodiments, the bridge will not be sintered to the other layers at the narrowest diameter of the neck so that the bridge remains free in the neck area. The bridge is sintered at its end, which contacts the first layer of encapsulation, shown in blue. By varying the gap at the sintering in the neck, the length and height of the bridge can be selected.

[0135] 14A and 14B, images of an exemplary encapsulated stent formed using the methods described in Figures 7A-10B and 12A-12D are shown. Specifically, Figure 14A shows a first diverging end region, a second diverging end region, and a constriction region of an encapsulated stent, where a bridge surrounds the constriction region to create a larger outer diameter, and where the bridge is configured to engage the atrial septum. Figure 14B shows the first diverging end region and the constriction region, which define the inner diameter (or constriction dimension) of the passageway through which blood flows.

[0136] As explained above, the neck region of the interatrial shunt may be expanded and / or contracted in vivo to provide suitable and customized fluid flow through the device for each given patient, for example, as further explained in U.S. Patent No. 10,898,698. Specifically, the frame, e.g., a metal frame, of the shunt device described herein may be heat treated during manufacture such that the neck region of the shunt is malleable, e.g., mechanically expandable, at body temperature and capable of contracting when heated above its martensitic finish temperature (Mf). The shunt device may also be heat treated such that the neck region may be contracted from an expanded state by a variable amount, e.g., a selected increment, depending on the temperature to which the frame of the shunt device is raised. For example, a shunt orifice in a region of the reduced diameter of a shunt device may be expanded in vivo from a first diameter to a second diameter greater than the first diameter, e.g., via mechanical expansion with a balloon catheter, and thermally contracted in vivo from the second diameter to a third diameter greater than the first diameter by heating the frame of the shunt device to a predetermined temperature above its martensitic finish temperature (Mf) but below its austenitic finish temperature (Af). Thus, a multiphase shunt device as described herein may be contracted to a desired diameter greater than its heat solidified austenitic configuration, without the need for it to be fully contracted, e.g., by heating the shunt device to a temperature above its austenitic finish temperature (Af), and subsequently mechanically expanded to the desired diameter.

[0137] Referring now to FIG. 15, a graph is provided illustrating an exemplary gradual incremental contraction of the shunt orifice of an intact frame shunt device and an encapsulated shunt device, respectively. Referring to Table 1 reproduced below, as shown in FIG. 15, during the baseline phase of the individual shunt devices, at body temperature, e.g., a water bath at 37° C., the intact frame shunt device has an orifice diameter of 5.00 mm and the encapsulated shunt device has an orifice diameter of 4.80 mm. In this study, the austenite finish temperature (Af) of each shunt device was about 60° C. While the temperature of the shunt devices remained at 37° C., both shunt devices were expanded via a nominal 9 mm angioplasty balloon, such that the intact frame shunt device shunt orifice expanded to 9.10 mm and the encapsulated shunt device shunt orifice expanded to 8.65 mm. Following expansion, the temperature of the water bath was slowly increased until a temperature of 70° C. was reached, such that the individual shunt orifice diameters of each shunt device were recorded in 5° C. increments. [Table 1]

[0138] As shown in Table 1, the data demonstrates that heat-activated shape memory recovery of individual shunt devices to their heat-set baseline shunt orifice diameters can be achieved incrementally, as opposed to an all-or-nothing transition. Thus, by heating the shunt devices to a predetermined temperature above their individual martensite finish temperatures (Mf) but below their individual austenite finish temperatures (Af), the shunt orifices of the shunt devices can be selectively thermally contracted to a desired diameter between their individual baseline diameters and their individual full balloon expansion diameters.

[0139] As shown in FIG. 15, both shunt devices begin their recovery, e.g., shrinkage from their respective 9 mm balloon expanded diameters toward their respective baseline diameters, at about 45° C., partially recover, e.g., to a shunt orifice diameter of about 7 mm, at about 52° C.-55° C., and fully recover, e.g., reach their respective baseline diameters, at about 60° C. Also, after the temperature of the water bath exceeds the austenite finish temperature (Af) of about 60° C., further shrinkage of the individual shunt devices may be minimal since the shunt orifices have essentially reached their respective baseline diameters. As will be appreciated by one of ordinary skill in the art, the shunt devices with bridges described herein may also be heat treated such that they may be selectively thermally shrunk to diameters other than their heat solidified austenitic configurations.

[0140] Thus, the shunt devices described herein, upon deployment in an inter-atrial shunt, may be temporarily balloon-expanded in vivo, e.g., to allow passage of a device, such as, e.g., the MitraClip® system, through the passageway of the shunt device, and subsequently effectively readjusted to a smaller desired diameter (greater than its full recovery diameter) by heating the shunt device to a specific temperature below its austenite finish temperature (Af), e.g., according to Table 1 above. For example, the shunt device may be heated in vivo by injecting a heated fluid, e.g., saline, into the shunt device via a catheter until the desired temperature of the shunt device is achieved. Preferably, the saline may be heated to an initial temperature, calculated based on, for example, the injection rate of the saline through the catheter and / or the size / dimensions of the catheter, including, for example, the length of the catheter, the wall thickness of the catheter, the inner diameter of the catheter's fluid lumen, and / or the outer diameter of the catheter, so that the saline is at a desired temperature when it is injected onto the shunt device in the interatrial septum, to account for any heat loss between the catheter lumen and external factors, such as the ambient air surrounding the catheter and / or body temperature blood, as the saline is injected through the catheter.

[0141] In addition, it is preferred to minimize dilution and loss of the heated saline as it is injected onto the shunt device in vivo. For example, an inflatable balloon, e.g., a hemispherical 10-16 mm diameter balloon, may be delivered and positioned adjacent the distal end region of the shunt device, e.g., across the interatrial septum, and inflated to its expanded state to block blood flow through the shunt device during injection of heated saline into the shunt device, e.g., via a large 8-14 Fr injection catheter with an outlet / port adjacent the proximal end region of the shunt device, thereby minimizing dilution and loss of the heated saline. A large injection catheter allows for rapid delivery of the heated saline, which further minimizes dilution and loss. Thus, a volume of heated saline must first be injected through the catheter to displace the entire volume within the fluid lumen of the injection catheter and deliver enough heated saline to heat the shunt device to the desired temperature in vivo. Alternatively, an inflatable balloon may be delivered and positioned adjacent the proximal region of the shunt device and inflated to its expanded state during injection of the heated saline to minimize dilution and washout of the heated saline. For example, a single infusion catheter having an inflatable balloon positioned thereon at a location proximal to the injection port of the infusion catheter may be used to simultaneously infuse heated saline into the shunt device and block blood flow through the shunt device.

[0142] In some embodiments, the distal tip of the infusion catheter may include a temperature sensor configured to generate one or more signals indicative of the temperature of the heated fluid delivered into the shunt device, for example, during the infusion of heated saline. Thus, the user may terminate the infusion of heated saline through the catheter when the temperature of the heated fluid delivered into the shunt device reaches a desired temperature and, accordingly, the shunt device achieves a desired shunt orifice diameter. Additionally or alternatively, the infusion of heated saline through the catheter into the shunt device may be automated via an automatic injector that is fluidly coupled to the infusion catheter and dynamically coupled to the temperature sensor. For example, the automatic injector may be programmed to automatically terminate the heated saline infusion when a signal received from the temperature sensor indicates that the temperature of the heated fluid delivered into the shunt device is at a predetermined desired temperature, for example, for a predetermined period of time.

[0143] In some embodiments, the power injector may include a heated syringe that may be set to a temperature higher than the austenite finish temperature (Af) of the shunt device at the interatrial septum. Thus, the power injector may be programmed to control the rate and duration of injection of heated saline through the catheter based on a feedback signal from a temperature sensor indicative of the temperature of the shunt device. For example, the temperature of the saline in the shunt location may be increased by increasing the rate of injection, or conversely, decreased by decreasing the injection rate, since cooling of the heated saline through the catheter is reduced by reducing the transit time of the heated saline through the catheter.

[0144] Thus, the power injector may be programmed to begin injection of heated saline through the catheter at a rapid rate to initially displace body temperature fluid within the catheter, followed by a slower injection rate as the temperature at the distal tip of the catheter increases as measured by the temperature sensor. The power injector may further be programmed to modulate the injection rate to maintain the temperature of the heated fluid delivered within the shunt device, and accordingly the temperature of at least a mid-region of the shunt device, over a predetermined period of time, e.g., one or two seconds or less, to ensure that the shunt orifice diameter has thermally contracted to a desired size.

[0145] Although the intermediate region of the shunt device is described herein as being heat treated such that it may be heated to a target temperature above its martensitic finish temperature (Mf) but below its austenitic finish temperature (Af) in order to selectively shrink to a predetermined size, as will be understood by one of ordinary skill in the art, in addition to or instead of the intermediate region, the proximal and / or distal end regions of the shunt device may be similarly heat treated to enable selective thermal shrinkage of the individual end regions.

[0146] Additional studies have shown that the "effective" austenite finish temperature (Af) of an ePTFE encapsulated shunt, i.e., the temperature required to return the shunt device to its heat-solidified austenitic configuration, can be, for example, about 5°C lower than the austenite finish temperature (Af) of the Nitinol frame. This is theorized to be due to the ePTFE acting as an insulator against heat transfer to the Nitinol frame. Thus, to achieve a target effective austenite finish temperature (Af) of, for example, 45-60°C for the encapsulated shunt, the Nitinol frame may be heat treated prior to encapsulation to exhibit an austenite finish temperature (Af) of 50-65°C. Additionally, the thickness of the encapsulating material may be selected to control the effective austenite finish temperature (Af) of the encapsulated shunt. For example, a first encapsulated shunt with a Nitinol frame having a first austenite finish temperature (Af) and a first encapsulation material thickness may have the same effective austenite finish temperature (Af) as a second encapsulated shunt with a Nitinol frame having a second austenite finish temperature (Af) higher than the first austenite finish temperature (Af) of the first encapsulated shunt and a second encapsulation material thickness thicker than the first encapsulation material thickness of the first encapsulated shunt.

[0147] Referring now to FIG. 16, an exemplary method 1600 for treating cardiac disease using a shunt device as described herein is provided. In step 1602, the shunt may be delivered in its contracted state to the interatrial septum using any of the delivery methods described above. For example, the shunt device may be delivered through the interatrial septum in a delivery sheath. As described above, the shunt device may have a diabolo shape, its proximal and distal end regions may be self-expandable at body temperature, and its middle / narrowed region may be heat treated during manufacture to exhibit a martensite finish temperature (Mf) such that it is malleable at body temperature and an austenite finish temperature (Af) above which it reverts to its heat-solidified austenite configuration. Also, the shunt device may be any one of a plain frame stent, an encapsulated shunt device, and / or may include a bridge, as described above.

[0148] In step 1604, the shunt device may be deployed to its expanded state in the interatrial septum such that its distal end region self-expands in the first atrium, its proximal end region self-expands in the second atrium, and the constriction region is disposed in the interatrial septum. In step 1606, the constriction region of the shunt device may be expanded, e.g., mechanically balloon expanded, such that a shunt orifice in the constriction region expands from a first diameter to a second diameter greater than the first diameter. For example, the second diameter may be selected to allow a medical device, e.g., a MitraClip® system, to pass therethrough. Alternatively, the second diameter may be selected to allow blood flow therethrough at a predetermined flow rate to provide an initial therapy.

[0149] Thus, if / when the shunt orifice of the shunt device needs to be reduced during the course of therapy, for example, based on patient physiological response, in step 1608, the shunt device may be heated to a predetermined temperature below the austenite finish temperature (Af), for example, according to Table 1 above, to selectively thermally shrink the shunt orifice in the region of the constriction to a predetermined third diameter, for example, a diameter smaller than the second diameter but larger than the first diameter. For example, as described above, an infusion catheter may be used to inject heated fluid, for example saline, into the shunt device in vivo to heat the shunt device to a desired temperature. Once the shunt device reaches the desired temperature and accordingly the shunt orifice in the region of the constriction delivers to a desired size, the delivery sheath and infusion catheter may be removed from the patient, leaving the shunt device implanted in the interatrial septum. Additionally or alternatively, other methods may be used to selectively thermally shrink the shunt device in vivo, such as, for example, applying an electrical current to the shunt device via one or more electrodes as described above. As will be understood by one of skill in the art, if the desired shunt orifice size is in a heat-solidified austenitic configuration, e.g., a first diameter, the shunt device may be heated to a temperature above the austenitic finish temperature (Af).

[0150] Although various illustrative embodiments of the present invention are described above, it will be apparent to those skilled in the art that various changes and modifications can be made therein without departing from the present invention. For example, although the examples of the device are described as having two or three components, it should be understood that the device can include any suitable number of components, each of which includes a self-expanding superelastic material or a malleable shape memory material. It is intended that the appended claims cover all such changes and modifications that fall within the true spirit and scope of the present invention.

Claims

1. A device for shunting blood between a patient's first and second atria to treat a medical condition, An encapsulated shunt comprising a first gradually widening end region, a second gradually widening end region, and a narrowing end region positioned between them, wherein the encapsulated shunt is formed from a frame encapsulated within a biocompatible material, and the encapsulated shunt defines a passage, allowing blood to flow through the passage from the first atrium to the second atrium, and A bridge extending from the first outer surface of the first gradual-end region to the second outer surface of the second gradual-end region, formed from a biocompatible material and configured to engage with the patient's atrial septum. The device comprising the above.

2. The device according to claim 1, wherein the reduced diameter region is expandable in vivo so that the passage expands from a first diameter to a second diameter larger than the first diameter.

3. The device according to claim 2, wherein the reduced diameter region is contractible in vivo so that the passage shrinks from the second diameter to a third diameter smaller than the second diameter.

4. The reduced diameter region contains nitinol that is malleable at body temperature and has an austenite termination temperature (Af) of 45 to 65°C, and The first and second gradual-end regions contain nitinol that is superelastic at body temperature and has an austenite termination temperature (Af) of 5 to 20°C. The device according to claim 1.

5. The device according to claim 1, wherein the reduced diameter region is mechanically expandable.

6. The device according to claim 1, wherein the reduced diameter region is thermally shrinkable.

7. The device according to claim 3, wherein the bridge defines an outer diameter larger than the first, second, and third diameters of the passage.

8. The device according to claim 7, wherein the outer diameter is 7 to 14 mm.

9. The device according to claim 7, wherein the bridge extends such as to form a gap between the inner surface of the bridge and the outer surface in the reduced diameter region.

10. The device according to claim 9, wherein when the reduced diameter region is expanded in a living organism, the gap is configured to decrease in size, and the outer diameter remains the same.

11. The device according to claim 1, wherein there is no gap between the bridge and the reduced diameter region.

12. The device according to claim 1, wherein the bridge is integrally formed with the biocompatible material that encapsulates the encapsulation shunt.

13. The device according to claim 1, wherein the biocompatible material of the bridge is different from the biocompatible material of the encapsulation.

14. The device according to claim 13, wherein the biocompatible material of the bridge is configured to allow tissue intrinsic growth, and the biocompatible material of the encapsulation is configured to inhibit tissue intrinsic growth.

15. The device according to claim 13, wherein the biocompatible material of the bridge has porosity selected to allow tissue intrinsic growth.

16. The device according to claim 1, wherein the biocompatible material of the bridge has an internode distance that exceeds the internode distance of the biocompatible material of the encapsulation.

17. The device according to claim 16, wherein the internode distance of the bridge material is selected to allow tissue intrinsic growth, while the internode distance of the encapsulation material is selected to inhibit tissue intrinsic growth.

18. The device according to claim 16, wherein the internode distance of the bridge material is greater than 30 microns, while the internode distance of the encapsulation material is less than or equal to 30 microns.

19. The device according to claim 18, wherein the internode distance of the bridge material is in the range of 45 to 200 microns.

20. The device according to claim 16, wherein the internode distance of the bridge material is 60 microns, while the internode distance of the encapsulation material is 30 microns.

21. The device according to claim 16, wherein the biocompatible material of the bridge and the biocompatible material of the encapsulation are expanded polytetrafluoroethylene (ePTFE).

22. The device according to claim 1, wherein the bridge is configured to remain engaged with the patient's atrial septum when the diameter-reducing region is contracted.

23. A method for creating a shunt, To provide a frame comprising a first gradually widening end region, a second gradually widening end region, and a reduced diameter region positioned between them, Encapsulating the frame with a biocompatible material, and Joining the biocompatible bridge to the encapsulated frame such that the bridge extends between the first outer surface of the first gradually widening end region and the second outer surface of the second gradually widening end region. The method, including the method described above.

24. The method according to claim 23, wherein the biocompatible material for encapsulating the frame and the biocompatible material for the bridge are formed from a single graft tube.

25. The method according to claim 23, wherein the biocompatible material of the bridge is different from the biocompatible material of the encapsulation.

26. The method according to claim 25, wherein the biocompatible material of the bridge is configured to allow tissue intrinsic growth, and the biocompatible material of the encapsulation is configured to inhibit tissue intrinsic growth.

27. The method according to claim 25, wherein the biocompatible material of the bridge has porosity selected to allow tissue growth.

28. Encapsulating the frame with a biocompatible material is Position at least the second widening end region of the frame within the first end of the graft tube, thereby placing the first layer of graft material on at least the second widening end region of the frame. Guide the second end of the graft tube into the second flaring end region of the frame and outside the first flaring end region of the frame, thereby arranging a second layer of graft material along the inner portion of the frame, and The second end of the graft tube is guided onto at least the first gradually widening end region of the frame, thereby placing a third layer of graft material on at least the first gradually widening end region of the frame, thereby forming a shunt-graft assembly. The method according to claim 23, including the method described in claim 23.

29. Encapsulating the frame with a biocompatible material further, To guide the first gradually widening end region onto the first portion of the mandrel such that the second end of the graft tube extends beyond the first gradually widening end region, Connecting the second part of the mandrel to the first part of the mandrel, Positioning a flexible sleeve having a shape configured to receive the encapsulated frame across the shunt-graft assembly while the shunt-graft assembly is positioned on the mandrel. The method according to claim 28, including the method described in claim 28.

30. Encapsulating the frame with a biocompatible material further, Positioning a compression jig having a first half and a second half, each having an inner recess sized and configured to receive the flexible sleeve, around the flexible sleeve covering the shunt-graft assembly, and Connecting the first half of the compression fixture to the second half of the compression fixture while the first and second halves of the compression fixture are positioned around the flexible sleeve covering the shunt-graft assembly. The method according to claim 29, comprising, wherein the compression jig, the flexible sleeve, the shunt-graft assembly, and the mandrel form a sintered assembly.

31. The method according to claim 30, wherein encapsulating the frame with a biocompatible material further comprises heating the sintered assembly to sinter the first, second, and third layers of the graft material together to form a monolithic graft material layer, thereby forming the encapsulated frame.

32. The method according to claim 23, further comprising encapsulating the frame, then guiding a second graft tube onto the outer portion of the encapsulated frame, thereby positioning the first end of the second graft tube on at least a portion of the first gradually widening end region of the encapsulated frame, and positioning the second end of the second graft tube on at least a portion of the second gradually widening end region of the encapsulated frame.

33. Joining the bridge of the biocompatible material to the encapsulated frame is Position the second widening end region, the narrowing diameter region, and at least a portion of the first widening end region of the frame within the first end of the second graft tube, thereby arranging the first layer of the second graft material on the second widening end region, the narrowing diameter region, and at least a portion of the first widening end region of the frame to form the bridge. The second end of the second graft tube is guided into the second flaring end region of the frame and out of the first flaring end region of the frame, thereby arranging the second layer of the second graft material along the inner portion of the frame, and The second end of the second graft tube is guided onto at least a portion of the first gradually widening end region of the frame, thereby positioning a third layer of the second graft material on at least a portion of the first gradually widening end region of the frame. The method according to claim 23, including the method described in claim 23.

34. Joining the bridge of the biocompatible material to the encapsulated frame further involves, To guide the first gradually widening end region onto the first portion of the mandrel such that the second end of the second graft tube extends beyond the first gradually widening end region, Connecting the second part of the mandrel to the first part of the mandrel, The positioning of a flexible sleeve having a shape configured to receive the encapsulated frame and the bridge while the shunt-graft assembly is positioned on the mandrel, wherein the reduced diameter region of the flexible sleeve has a diameter corresponding to the diameter of the bridge. The method according to claim 33, including the method described in claim 33.

35. Joining the bridge of the biocompatible material to the encapsulated frame further involves, Positioning a compression jig having a first half and a second half, each having an inner recess sized and configured to receive the flexible sleeve, around the flexible sleeve covering the shunt-graft assembly, and The first half of the compression fixture is connected to the second half of the compression fixture while the first and second halves of the compression fixture are positioned around the flexible sleeve covering the shunt-graft assembly. The method according to claim 34, wherein the compression jig, the flexible sleeve, the shunt-graft assembly, and the mandrel form a sintered assembly.

36. The method according to claim 35, wherein joining the bridge of the biocompatible material to the encapsulated frame further comprises heating the sintered assembly such that a non-joint gap is created between the reduced diameter region and the bridge, wherein the first, second, and third layers of the second graft material are sintered to the encapsulated frame, except for the outer surface of the reduced diameter region.

37. The method according to claim 33, wherein the second graft tube has a diameter exceeding the diameter of the reduced diameter region.

38. The method according to claim 23, wherein the bridge comprises a first end configured to extend to approximately the middle of the first gradually widening end region, and a second end configured to extend to approximately the middle of the second gradually widening end region, such that a gap is created between the reduced diameter region and the bridge.

39. A shunt to be implanted in the atrial septum to treat heart disease, the shunt is A single metal frame comprising a proximal region, a distal region, and an intermediate region positioned between them, wherein the proximal and distal regions are each configured to self-expand in response to deployment so that the intermediate region is positioned within the opening in the atrial septum, the single metal frame comprising the single metal frame having been heat-treated so that the intermediate region is bio-adjustable between a first diameter and a second diameter greater than the first diameter, wherein, The shunt wherein the intermediate region is selectively thermally shrinkable from the second diameter to a third diameter larger than the first diameter.

40. The shunt according to claim 39, wherein the intermediate region is heat-treated to exhibit a martensite termination temperature and an austenite termination temperature exceeding the martensite termination temperature, and the intermediate region is selectively thermally shrinkable by heating to a temperature between the martensite termination temperature and the austenite termination temperature.

41. The shunt according to claim 39, wherein the intermediate region contains nitinol that is malleable at body temperature and has an austenite termination temperature of 45 to 65°C.

42. The shunt according to claim 39, wherein the proximal and distal regions contain nitinol that is superelastic and has an austenite termination temperature of 5 to 20°C.

43. The shunt according to claim 39, wherein the intermediate region is heat-treated to exhibit different shape memory properties from the proximal and distal regions.

44. The shunt according to claim 39, wherein the intermediate region comprises a material that is plastically deformable so that the intermediate region is configured to expand in a living organism via mechanical expansion.

45. The shunt according to claim 39, wherein the proximal region, distal region, and intermediate region define a diabolo-shaped shunt.

46. The shunt according to claim 39, further comprising a biocompatible material coating the proximal region, the distal region, and the intermediate region to define a passage for enabling blood to flow through the shunt across the atrial septum.

47. The shunt according to claim 46, wherein the flow path cross-sectional area of ​​the passage is smallest in the intermediate region.

48. The shunt according to claim 46, wherein the passage is sized and shaped to allow a sufficient amount of blood to flow through the shunt and across the atrial septum in order to treat pulmonary hypertension.

49. The shunt according to claim 46, wherein the passage is sized and shaped to allow a sufficient amount of blood to flow through the shunt and across the atrial septum in order to treat heart failure.

50. The shunt according to claim 39, further comprising a bridge extending from a first outer surface of the proximal end region to a second outer surface of the distal end region, the bridge being formed from a biocompatible material and configured to engage with the atrial septum, The bridge defines the shunt, which has an outer diameter larger than the second diameter.

51. The shunt according to claim 50, wherein the bridge extends to form a gap between the inner surface of the bridge and the outer surface of the intermediate region, and is configured such that when the intermediate region expands in vivo, the size of the gap decreases and the outer diameter remains the same.

52. The shunt according to claim 50, wherein the bridge extends to form a gap between the inner surface of the bridge and the outer surface of the intermediate region, and is configured such that the gap increases in size when the intermediate region is contracted in vivo, and the outer diameter remains in contact with the atrial septum.

53. A system for treating heart disease, comprising a shunt according to claim 39, further comprising a catheter configured to inject a heated fluid into the shunt in vivo, thereby heating the shunt in vivo to selectively thermally contract the intermediate region from the second diameter to the third diameter.

54. The system according to claim 53, wherein the heated fluid has an initial fluid temperature selected to reach a target fluid temperature in the shunt in vivo, based on at least one of the injection rate of the heated fluid or the size of the catheter, thereby heating at least the intermediate region of the shunt to a target shunt temperature.

55. The system according to claim 53, wherein the catheter comprises a temperature sensor located on the distal tip of the catheter, the temperature sensor being configured to generate a signal indicating the temperature of a heated fluid delivered into the shunt.

56. The system according to claim 55, further comprising an automatic injector fluidly coupled to the catheter for injecting a heated fluid through the catheter, the automatic injector being operationally connected to the temperature sensor and programmed to adjust at least one of the injection rate or injection time of the heated fluid through the catheter in response to the signal.

57. The system according to claim 56, wherein the automatic injector is programmed to automatically terminate the injection of the heated fluid through the catheter when the temperature of the heated fluid delivered into the shunt reaches a predetermined temperature sufficient to selectively thermally contract the intermediate region of the shunt to the third diameter.

58. The system according to claim 56, wherein the automatic injector is programmed to regulate the rate of injection of the heated fluid through the catheter to maintain the temperature of the heated fluid delivered into the shunt for a predetermined period of time.

59. The system according to claim 53, wherein the catheter is sized and shaped to inject the heated fluid at an injection rate selected to minimize dilution and leakage of the heated fluid within the body.

60. The system according to claim 59, further comprising an inflatable balloon configured to be positioned in an inflated state adjacent to the distal end region during the injection of the heated fluid, in order to minimize the dilution and leakage of the heated fluid within the living body.

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