Room diverter with expanded neck region

Abstract

The device is provided with a bridge made of a biocompatible material that surrounds a neck region that encloses the shunt. In one embodiment, the bridge is configured to maintain an outer diameter to regulate a fluid flow rate through the device when the diameter of the neck region is reduced or increased within the body. The bridge may have an outer diameter configured to be placed within the enlarged spaced holes. A method for regulating the flow of a fluid and a method of manufacturing such a device are also provided.

Description

CROSS-REFERENCE TO RELATED APPLICATIONS

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

[0002] The 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 making such devices. Background Art

[0003] For many medical conditions, it is beneficial to regulate the flow of fluids in the human body, such as through a channel between two body cavities. Such channels are generally used in catheterization procedures where a catheter is delivered through the patient's vasculature. In certain catheterization procedures, it is beneficial to move from one cavity to another by creating a channel. For example, such a channel can be formed between the right side of the heart and the left side of the heart, for example, between the right atrium toward the left atrium, where a clinical procedure is performed on the left side of the heart using an entry point 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] Additionally, channels can be created and maintained in the heart wall between the two heart chambers to accommodate shunts, which are used to redistribute blood from one heart chamber to the other 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 demands or can only be met at higher filling pressures. HF has multiple potential causes, including MI, coronary artery disease, valvular disease, hypertension (such as PAH), and myocarditis. Chronic heart failure is associated with neurohormonal activation and altered autonomic control. Although these compensatory neurohormonal mechanisms provide valuable support to the heart under normal physiological circumstances, they also play an important role in the development of HF and its subsequent progression.

[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 the blood ejected from the left ventricle (cardiac output) divided by the maximum volume of the left ventricle at the end of diastole, or the relaxation period. A normal ejection fraction is greater than 50%. Systolic heart failure typically results in a reduction in ejection fraction of less than 40%. Such patients suffer from heart failure with reduced ejection fraction ("HFrEF"). Patients with HFrEF typically may have a larger left ventricle due to a phenomenon known as "cardiac remodeling" that occurs subsequent to higher ventricular pressures.

[0006] In DHF, the heart normally contracts normally with a normal ejection fraction, but is stiffer or less compliant than a healthy heart when relaxed and filled with blood. Such patients are said to have heart failure with preserved ejection fraction ("HFpEF"). This stiffness can hinder blood from filling the heart and create backup for the lungs, which can lead to pulmonary venous hypertension and pulmonary edema. HFpEF is more common in patients older than 75 years, especially in hypertensive women.

[0007] Both variants of HF are treated with pharmacological approaches, which generally involve the use of vasodilators to reduce the workload of the heart by reducing systemic vascular resistance, and diuretics, which inhibit fluid accumulation and edema formation and reduce cardiac filling pressures. No drug therapy has been shown to improve the morbidity or mortality of HFpEF, however, several classes of drugs have had a major impact on the management of patients with HFrEF - including renin-angiotensin antagonists, neprilysin inhibitors, beta blockers, mineralocorticoid antagonists, and sodium-glucose co-transporter-2 (SGLT2) inhibitors. Despite this, overall, HF remains a progressive disease and cardiac function and symptoms worsen over time in most patients. In the United States, more than 1 million patients are hospitalized each year for severe HF exacerbations, and the mortality rate is higher than that of most forms of cancer.

[0008] In more severe cases of HFrEF, mechanical circulatory support (MCS) devices (such as mechanical pumps) are used to relieve the heart by performing all or part of the pumping functions normally performed by the heart. Chronic left ventricular assist devices ("LVAD"), total artificial hearts, and heart transplants are used as measures of last resort. However, these assist devices are generally intended to improve the pumping capacity of the heart to increase cardiac output to a level compatible with normal life and support the patient until a donor heart for transplantation is available. This use of MCS is also known as "bridge to transplant therapy" while waiting for a heart transplant. Due to the insufficient supply of donor hearts for transplantation to meet demand, MCS is more often the only treatment option - also known as "destination therapy". Such mechanical devices are able to advance significant blood volumes (liters / minute), but are limited by the need for power supplies, relatively large pumps, and pose risks of hemolysis, thrombosis, and infection. Temporary assist devices, intra-aortic balloons, and pacing devices have also been used.

[0009] Various devices have been developed that utilize stents to alter blood pressure and blood flow between a given vessel or heart chamber. For example, U.S. Pat. No. 6,120,534 to Ruiz relates to an intraluminal stent for regulating fluid flow through a human vessel or organ, such as for regulating blood flow through a pulmonary artery to treat congenital heart defects. The stent may include an expandable mesh having balloon expandable lobed or tapered portions connected by a shape memory contraction region that restricts blood flow through the stent. The contraction region may be adjusted in vivo and may additionally be heated to restore maximum contraction. Ruiz does not mention treatment of HF or reduction of left atrial pressure.

[0010] McNamara's U.S. Patent Publication No. 2013 / 0178784 describes an adjustable decompression shunt that can be expanded, for example, via an inflatable 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 (e.g., an elastic coil spring or a series of ratchets and unidirectional mechanical ramps) in response to measurements of the patient's physiological parameters. A major disadvantage of the method described in the patent is the hysteresis effect, that is, when the shunt is permanently deformed, the underlying crystalline structure is irreversibly changed. Importantly, this plastic deformation may cause stress and fatigue-related fractures of the device. Another disadvantage of the system is that, over time, after tissue has been attached to the shunt, expanding or shrinking the encapsulated shunt can cause trauma to the atrial septum.

[0011] U.S. Patent No. 6,468,303 to Amplatz et al. describes a collapsible medical device and related methods for shunting selected organs and vessels. Amplatz describes that the device can be adapted to shunt a septal defect in a patient's heart, for example, by creating a shunt in the atrial septum of a neonate with hyperplastic left heart syndrome ("HLHS"). The patent also describes that increasing the mixing of pulmonary and systemic venous blood improves oxygen saturation, and that the shunt can then be closed by an occluding device. Amplatz does not mention treatment of HF or reduction of left atrial pressure and methods for regulating the rate at which blood flows through the device.

[0012] Implantable interatrial shunt devices have been successfully used in patients with severe symptomatic heart failure. By diverting or shunting blood from the left atrium ("LA") to the right atrium ("RA"), the pressure in the left atrium is reduced or prevented from rising as high as it would otherwise (left atrial decompression). It is expected that this achievement will prevent, alleviate or limit the symptoms, signs and syndromes associated with 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 fluid dynamics of the shunt device. The latter is generally affected by the geometry and material composition of the shunt. For example, the general flow properties of similar shunt designs have been shown to be related to the atrial pressure gradient and the effective orifice diameter.

[0014] Percutaneous implantation of an interatrial shunt typically requires transseptal catheterization immediately prior to insertion of the shunt device. The transseptal catheterization system is placed from the entry site of the femoral vein, across the atrial septum in the fossa ovalis ("FO") region, which is the center and thinnest region of the atrial septum. The major axis dimension of the adult FO is generally 15-20 mm and the thickness is <3 mm, but in some cases it can be up to 10 mm thick. Access to the LA chamber can be achieved using a variety 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 expanded to facilitate the passage of a shunt device with the desired orifice size. The expansion is usually completed by advancing a tapered sheath / dilator catheter system or an angioplasty-type balloon across the FO. This is the general location where a congenital secondary atrial septal defect ("ASD") will be located.

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

[0016] In addition, after implantation of a shunt device in the heart wall, tissue ingrowth, including an endothelial layer or a neointimal layer, generally forms on the device, thereby inhibiting thrombosis of the shunt device. Adjustment of the shunt after tissue ingrowth at the atrial septum may be risky if the shunt is pulled away from the tissue ingrowth at the atrial septum during adjustment. Additionally, or alternatively, anchoring in an atrial septum with a relatively large hole (e.g., by a device such as The presence of a shunt in the interventricular cavity (caused by the transseptal access during insertion) can present challenges to shunt fixation. Therefore, there is a need for an interatrial shunt that is securely adjustable and / or can be securely anchored within a relatively large septal cavity. Summary of the invention

[0017] Devices and methods for making and using improved interatrial shunts are described herein to improve treatment and outcomes for patients with cardiovascular and cardiopulmonary disorders, including heart failure (HF) and pulmonary arterial hypertension (PAH). Interatrial shunts are particularly suitable for in vivo regulation of the size of a flow path through the shunt and / or for use in relatively large holes in the septum (e.g., following transseptal interventions).

[0018] The device may be designed to shunt blood between a first atrium and a second atrium of a patient to treat a medical condition. The device may include an encapsulated shunt including a first flared end region, a second flared end region, and a neck region disposed therebetween. The encapsulated shunt may be formed by a frame encapsulated in a biocompatible material such as ePTFE so that it defines a passage to allow blood to flow from the first atrium to the second atrium via the passage. The device may also include a bridge extending from a first outer surface of the first flared end region to a second outer surface of the second flared end region. The bridge may be made of a biocompatible material and configured to engage the patient's atrial septum.

[0019] The bridge can be configured to engage the patient's atrial septum rather than the encapsulated shunt itself engaging the patient's atrial septum, so that when the device is adjusted in vivo, the bridge can be configured to maintain the outer diameter unchanged, while only the inner diameter of the shunt is changed. The bridge is preferably designed to prevent cracking and reduce tissue trauma that may be caused by such adjustment. The bridge can also be configured to maintain contact with the septal tissue so that leakage or bypass flow around the outer surface of the shunt is minimized. In addition, the bridge can also be used in conjunction with an encapsulated shunt that is not adjustable in vivo. For example, incorporating a biocompatible material bridge into the encapsulated shunt increases the outer diameter of the device, thereby allowing the device to be implanted in an enlarged septal hole without affecting the inner diameter of the device and the fluid flow rate through the device.

[0020] According to another aspect of the present invention, a shunt is provided for implantation in the atrial septum to treat a cardiac condition. The shunt may include a single metal frame including a proximal region, a distal region, and an intermediate region disposed therebetween. The proximal region and the distal region may each be configured to self-expand when deployed so that the intermediate region is positioned in the opening of the atrial septum. In addition, the single metal frame may be heat treated so that the intermediate region is adjustable in vivo between a first diameter and a second diameter greater than the first diameter, and the intermediate region is selectively capable of heat shrinking 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 martensitic finish temperature and an austenitic finish temperature greater than the martensitic finish temperature, so that the intermediate region is selectively capable of heat shrinking by being heated to a temperature between the martensitic finish temperature and the austenitic finish temperature.

[0021] The middle region may be ductile at body temperature and may include Nitinol having an austenite finish temperature between 45-60°C. The proximal region and the distal region may be superelastic and may include Nitinol having an austenite finish temperature between 5-20°C. In addition, the middle region may also be heat treated to exhibit shape memory properties that are different from the proximal region and the distal region. In addition, the middle region may also include a plastically deformable material such that the middle region may be configured to expand in vivo via mechanical expansion. The proximal region and the distal region and the middle region may define a diabolo-shaped shunt.

[0022] The system may also include a biocompatible material coating the proximal region, the distal region, and the middle region to define a passageway to allow blood to flow through the atrial septum via the shunt. The cross-sectional flow area of ​​the passageway at the middle region may be minimal. In addition, the passageway may also be sized and shaped to allow a sufficient amount of blood to flow through the atrial septum via the shunt to treat pulmonary hypertension. Additionally or optionally, the passageway may also 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 also include a bridge extending from a first outer surface of the proximal region to a second outer surface of the distal region. The bridge may be made of a biocompatible material and configured to engage the atrial septum. In addition, the bridge may define an outer diameter greater than the second diameter. For example, the bridge may extend to form a gap between the inner surface of the bridge and the outer surface at the middle region, such that when the middle region is expanded in vivo, the gap may be configured to decrease in size, while the outer diameter may be configured to maintain a constant diameter. Additionally or alternatively, the bridge may extend to form a gap between the inner surface of the bridge and the outer surface at the middle region, such that when the middle region is contracted in vivo, the gap is configured to increase in size, while the outer diameter is configured to maintain contact with the septum.

[0023] According to another aspect of the present disclosure, a system for treating a cardiac condition is provided, the system comprising the shunt. For example, the system may further include a catheter configured to inject a heated fluid into the shunt in vivo to heat the shunt in vivo, thereby selectively shrinking the middle region from the second diameter to the third diameter. For example, the heated fluid may have an initial fluid temperature, the initial fluid temperature being 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 at the shunt in vivo, thereby heating at least the middle region of the shunt to a target shunt temperature. In addition, the catheter may further include a temperature sensor disposed at a distal tip of the catheter, the temperature sensor being configured to generate a signal indicating the temperature of the heated fluid delivered within the shunt. The catheter may be sized and shaped to inject the heated fluid at an injection rate selected to minimize dilution and washout of the heated fluid in vivo. In some embodiments, the system may include an inflatable balloon configured to be disposed adjacent the distal region in an expanded state to occlude blood flow through the shunt during injection of the heated fluid, thereby minimizing dilution and loss of the heated fluid in vivo.

[0024] In addition, the system may also include a power injector that is fluidly coupled to the conduit to inject a heated fluid through the conduit. The power injector may be operably coupled to the temperature sensor and programmed to adjust at least one of a rate or duration of injection of the heated fluid through the conduit in response to the signal. The power injector may be programmed to automatically terminate injection of the heated fluid through the conduit when the temperature of the heated fluid delivered within the shunt reaches a predetermined temperature sufficient to selectively heat shrink the intermediate region of the shunt to the third diameter. In addition, the power injector may be programmed to adjust the rate of injection of the heated fluid through the conduit so as to maintain the temperature of the heated fluid delivered within the shunt for a predetermined period of time.

[0025] According to yet another aspect of the present invention, a method for treating a cardiac condition is provided. The method may include: delivering a shunt into an opening of the atrial septum in a contracted state, the shunt comprising a frame having a proximal region and a distal region and an intermediate region disposed therebetween, the frame being heat treated so that the intermediate region is adjustable in vivo between a first diameter and a second diameter greater than the first diameter; deploying the shunt within the opening to an expanded state, wherein the proximal region self-expands within a first atrium, the distal region self-expands within a second atrium, and the intermediate region is disposed at the atrial septum; expanding the intermediate region of the shunt in vivo from the first diameter to the second diameter; and selectively heat shrinking the intermediate region of the shunt in vivo from the second diameter to a third diameter greater than the first diameter.

[0026] Expanding the middle region of the shunt may include mechanically expanding the middle region of the shunt. In addition, selectively heat shrinking the middle region of the shunt may include heating the middle 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 middle region of the shunt. The method may also include measuring the temperature of the heated fluid delivered within the shunt via a temperature sensor disposed at the distal tip of the catheter, and adjusting at least one of the rate or duration of injection of the heated fluid through the catheter based on the measured temperature of the heated fluid delivered within the shunt. In addition, the method may also include automatically terminating the injection of the heated fluid through the catheter when the temperature of the heated fluid delivered within the shunt reaches a predetermined temperature sufficient to selectively heat shrink the middle region of the shunt to the third diameter, and / or adjusting the rate of injection of the heated fluid through the catheter to maintain the temperature of the heated fluid delivered within the shunt for a predetermined period of time. The method may further include expanding an inflatable balloon adjacent to the distal region of the shunt to an expanded state to block blood flow through the shunt during injection of the heated fluid, thereby minimizing dilution and loss of the heated fluid in vivo. In addition, the method may further include allowing a sufficient amount of blood to flow through the atrial septum via the shunt to treat pulmonary hypertension and / or heart failure. BRIEF DESCRIPTION OF THE DRAWINGS

[0027] These and other features, aspects, and advantages of the present invention will become apparent from the following description, appended claims, and accompanying exemplary embodiments as illustrated in the accompanying drawings, the following of which is briefly described.

[0028] Figure 1A and Figure 1B 2 is a side view of an hourglass shaped flow divider with and without a bridge constructed according to the method of the present invention.

[0029] Figure 1C-1F yes Figure 1B Side, cross-sectional and perspective views of the device.

[0030] Figure 2A and Figure 2B yes Figure 1B Cross-sectional side view of the device.

[0031] Figure 3A-3E Shown for use in humans Figure 1B Example steps of the device.

[0032] Figure 4A-4D The structure of a shunt-graft assembly having a double layer region and a triple layer region is shown.

[0033] Figure 5A-5E are side views sequentially illustrating the technique of depositing a first graft portion on a shunt.

[0034] Fig. 6A and Figure 6B are side views sequentially illustrating the technique of attaching a second graft portion to the shunt.

[0035] Fig. 7A and Figure 7B Shown include Fig. 7A The first spindle portion and Figure 7B The mandrel assembly of the second mandrel section is shown.

[0036] Figure 8A-8C are side and close-up views showing the technique for attaching a third graft portion over the shunt.

[0037] Fig. 9A and Fig. 9B A perspective view of a flexible sleeve and a side view of the flexible sleeve mounted on a shunt-graft assembly are shown.

[0038] Fig. 10A and Fig. 10B A perspective view of the compression shell and a side view of the compression shell mounted on a flexible sleeve and shunt-graft assembly are shown.

[0039] Figures 11A-11F Shown is a bridge having a Figure 4A-4D A first embodiment of a shunt-graft assembly.

[0040] Figures 12A-12D Shown is a bridge having a Figure 4A-4D A second embodiment of a shunt-graft assembly.

[0041] Fig.13 Another embodiment of a shunt-graft assembly having a bridge is shown.

[0042] Figure 14A-14B Shows Figures 12A-12D Image of the encapsulated stent.

[0043] Fig.15 is a graph showing incremental thermal contraction of a diverter orifice according to the principles of the present disclosure.

[0044] Fig.16 is a flow chart illustrating exemplary method steps for treating a cardiac condition according to principles of the present disclosure. DETAILED DESCRIPTION

[0045] Devices and methods for making and using improved interatrial shunts to improve treatment and outcomes for patients with cardiovascular and cardiopulmonary disorders such as pulmonary arterial hypertension (PAH) or heart failure (HF) are described herein. In some aspects, the device has a size that can be reduced and increased in vivo.

[0046] The device of the present invention can be implanted in the human body permanently or temporarily and includes one or more components that can be adjusted in size (larger or smaller) after implantation. For example, such an adjustable device may be needed in the treatment of pulmonary arterial hypertension (PAH) or heart failure (HF). In PAH, a shunt is placed in the atrial septum, allowing the excessive blood pressure in the right atrium to be relieved by allowing a certain amount of blood to flow from the right atrium through an orifice to the left atrium. In HF, a shunt is placed in the atrial septum, allowing the excessive blood pressure in the left atrium to be relieved by allowing a certain amount of blood to flow from the left atrium through an orifice to the right atrium. In both PAH and HF, atrial shunts have been shown to be effective in reducing symptoms and increasing exercise tolerance. Atrial shunts can also reduce the need for hospitalization and even increase life expectancy.

[0047] However, if the orifice of the interatrial shunt is too small, too little blood can be diverted and the shunt may be relatively ineffective and provide little clinical benefit. In contrast, shunting too much blood through an orifice that is too large ("over-shunting") may lead to serious or even fatal complications over time. For example, in patients with PAH, over-shunting may lead to systemic oxygen desaturation and its sequelae, including cyanosis, polycythemia with increased blood viscosity, end-organ ischemia, and potential death. In patients with HF, over-shunting may lead to pulmonary hypertension, right ventricular failure, and potential death.

[0048] Currently, there is no known method to predict the response of a given patient to a particular shunt orifice size. As previously known, the shunt orifice can be enlarged in vivo, for example, by expanding the shunt through an inflatable balloon catheter or other similar mechanical expansion device within a properly designed shunt, provided, however, that the shunt is made of a ductile material and will remain expanded due to plastic deformation or some other physical property, whereby when the balloon or other expansion device is removed, the amount of elastic rebound or rebound will be low enough to achieve the desired orifice size increment. A disadvantage of this method is that only the orifice size can be increased. If the shunt starts out too large or becomes too large due to balloon expansion, but the patient needs a smaller shunt, there is no way to return to a smaller shunt except to provide a smaller shunt or place a smaller shunt in the cavity of the original shunt. This technology is called "shunt-in-shunt". Thus, finding the appropriate shunt orifice size for a given patient is a process of trial and error, wherein the shunt orifice size is selected based on the patient's response, which may be observed over a period of time, ranging from a few minutes to a few months, and the shunt orifice size is increased (e.g., by balloon expansion) or decreased (by providing a new, smaller shunt) based on the patient's response. Thus, the opportunities to increase or decrease the size of the shunt are very limited and may not be repeatable. In addition, the extent to which an expandable balloon catheter can expand the shunt orifice may be limited by the maximum size of the balloon. Therefore, what is needed is a means to repeatedly and non-invasively adjust the orifice size of shunts and other implantable devices in vivo and in both directions - either larger or smaller.

[0049] In some examples, the device provided herein may incorporate adjustable cross-sectional flow area technology, and the adjustable cross-sectional flow area can be easily reduced and / or expanded in vivo in any order according to clinical needs. Examples of room shunts with adjustable cross-sectional flow areas are described in U.S. Patent No. 9,724,499 to Rottenberg et al., U.S. Patent 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, and the entire contents of each article are incorporated herein by reference. Some examples of the present device include self-expanding superelastic (austenite phase) materials and ductile shape memory (martensite phase) materials. When the device is implanted in the human body-for example, by transporting the device in a compressed state to the desired position in the sheath and then removing the sheath-the self-expanding superelastic material can be automatically deployed to its desired size, and the ductile shape memory material can initially remain in a reduced size state. The cross-sectional area of ​​the ductile shape memory material can then be expanded and reduced in vivo as needed (e.g., by providing an appropriate fluid flow rate) to obtain a cross-sectional area suitable for treating a patient, or to properly secure the device in the patient's body while allowing repositioning to improve the effectiveness of the treatment. A wide variety of devices, such as those exemplified herein, can be prepared using components that include a self-expanding superelastic material and a ductile shape memory material, respectively.

[0050] When the size of the shunt is adjusted over time, one complication that may occur is tissue trauma to the atrial septum. The encapsulated shunt can be designed to promote tissue ingrowth and endothelialization, so that over time, after the tissue is attached to the encapsulated shunt, the encapsulated shunt is expanded or reduced to cause trauma to the atrial septum. This article provides a device for adjusting the size of the shunt without interfering with the septal tissue surrounding the device. Specifically, the device may include a bridge formed of a biocompatible material, which extends between the outer surface of the first flared end region and the second flared end region, thereby creating a gap between the bridge and the neck region of the encapsulated shunt. The bridge can be configured to engage the patient's atrial septum, while the non-encapsulated shunt itself engages the patient's atrial septum, so that when the device is adjusted in vivo, the bridge can be configured to keep the outer diameter unchanged, and only the inner diameter of the shunt is changed. Therefore, the bridge prevents cracking that may be caused when the device is adjusted in vivo. In addition, after the diameter of the neck region is reduced in vivo, the bridge can mitigate any bypass flow that can flow around the outside of the device.

[0051] The above-described bridges may also be used in conjunction with encapsulated shunts that are not adjustable in vivo. Patients who may benefit from an interatrial shunt may also have required or will require a prior interatrial septal procedure to create a hole in the interatrial septum. Alternatively, the patient may have a septal defect that was previously dilated to be larger than the delivery system required for implantation of the devices described herein. Incorporating a biocompatible material bridge into an encapsulated shunt increases the outer diameter of the device, thereby allowing the device to be implanted into the enlarged septal hole without affecting the inner diameter of the device and the rate of fluid flow through the device.

[0052] In some examples, the device can be or include an hourglass-shaped or "diabolo"-shaped shunt, which is optionally encapsulated with a biocompatible material and can be used to treat subjects suffering from disorders (such as CHF or PAH) in which regulating fluid flow may be useful. In some examples, the hourglass-shaped shunt can be specifically configured to be securely stuck in the atrial septum, for example, in an opening through the fossa ovalis, thereby allowing blood to flow from the left atrium to the right atrium when the blood pressure in the left atrium exceeds the blood pressure in the right atrium, or allowing blood to flow from the right atrium to the left atrium when the blood pressure in the right atrium exceeds the blood pressure in the left atrium. As provided herein and described in more detail in PCT application WO 2021 / 224736 incorporated above, the internal dimensions of the hourglass-shaped shunt can be appropriately adjusted in vivo, for example, to regulate the flow of fluid therethrough, for example, to regulate the flow of fluid through the atrial septum between the left atrium and the right atrium.

[0053] Reference now Figure 1A and Figure 1B , showing a shunt 100 with and without a bridge. The shunt 100 is hourglass-shaped or "diabolo"-shaped and may include a first component 110, a second component 120, and a third component 130, which are fluidically coupled to each other. In some embodiments, the shunt 100 has an internal dimension 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 ductile shape memory material, and the third component 130 may include a second self-expanding superelastic material. The ductile shape memory material of the second component 120 may have a first cross-sectional area (or diameter) that allows a first rate of fluid to flow through the second component, may be capable of expanding to a second cross-sectional area (or diameter) that allows a second rate of fluid to flow through the second component, and may be capable of contracting to a third cross-sectional area (or diameter) that allows a third rate of fluid to flow through the second component. The total rate of fluid flow through the device 100 may also depend on the cross-sectional area (or diameter) of the first component 110 and the third component 130.

[0054] The flow diverter 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, and the entire contents of each are incorporated herein by reference.

[0055] The first component 110 may include any suitable number of rings (e.g., rings 112, 113) formed of or including a first self-expanding material, and may optionally be sinusoidal. The second component 120 may include any suitable number of rings (e.g., rings 114) formed of or including a ductile shape memory material, and may optionally be sinusoidal. The third component 130 may include any suitable number of rings (e.g., rings 115, 116) formed of or including 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.

[0056] The first component 110 may provide a first flared end region 102, the third component 130 may provide a second flared end region 106, and the second component 120 may provide a neck region 104 disposed between the first flared end region and the second flared end region. The inlet and outlet of the device 100 may include flanges 102, 106, and the neck 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 when deployed; and the sinusoidal ring 114 may have sufficient strength to maintain its diameter when the balloon is expanded or heat-contracted.

[0057] exist Figure 1A In the non-limiting example shown, the first flared end region 102 has a first end region dimension D1, the second flared end region 106 has a second end region dimension D2, and the neck region 104 has a neck dimension D3 - which may be expressed as, for example, with reference to Figure 3A-3D The second component 120 shown is enlarged or reduced in the manner described. Figure 1A As shown, the neck region 104 of the flow splitter 100 can be significantly narrower than the flared end regions 102 and 106, for example, can have a smaller cross-sectional area and smaller dimensions than the flared end regions 102 and 106. Figure 1AAs shown, the shunt 100 can be asymmetric. For example, the shunt 100 can be asymmetric to take advantage of the natural features of the atrial septum and the left and right atrial chambers of the heart. Optionally, the hourglass-shaped shunt 100 can be symmetrical, wherein the first end region dimension D1 is equal to the second end region dimension D2. The first flared end region 102 and the second flared end region 106 can also have straight or curved profiles, or both. For example, the support 111 has a straight profile, while the support 108 has a curved profile. In addition, the first flared end region 102 and the second flared end region 106 can present any angular position consistent with the hourglass configuration.

[0058] The device 100 may optionally be made from a single tube of material that is laser cut to define a plurality of struts and connecting members, such as a plurality of sinusoidal rings connected by longitudinally extending struts. The sinusoidal rings and longitudinal struts may be laser cut to form a unitary, integral piece of construction, and different regions of the piece may be heat treated differently from one another to produce components having different austenite finish temperatures (Af) from one another in a manner as described elsewhere herein. Alternatively, the sinusoidal rings of the first component 110, the second component 120, and the third component 130 may be separately defined to form different pieces of material having suitable Af that are subsequently joined together to form the device 100. The device 100 may also be electropolished to reduce thrombosis.

[0059] In some examples, the first self-expanding superelastic material of the first component 110, the ductile shape memory material of the second component 120, and the second self-expanding superelastic material of the third component 130 may include materials that are different from each other, or may include 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 include one or more materials selected from nickel titanium alloy (NiTi) (also known as Nitinol), other shape memory alloys, self-expanding materials, superelastic materials, polymers, etc. In a non-limiting example, the first component 110 and the third component 130 may each include a Nitinol alloy that has an austenite finish temperature (Af) that is sufficiently lower than body temperature so that the material is in an austenite superelastic phase in the human body. In a non-limiting example, the self-expanding superelastic material of the first component 110 and the third component 130 includes Nitinol having an Af lower than 37°C. For example, the Af of the Nitinol of the self-expanding superelastic material may be between 5-20°C. The first component 110, the second component 120, and the third component 130 may optionally be integrally formed with one another from a common frame. For example, the first component 110, the second component 120, and the third component 130 may initially be cut and processed into a single unit from the same tube, sheet, or other suitable frame configuration as one another. Portions of the common frame may be heat treated differently from one another 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 the PCT application WO 2021 / 224736 incorporated above.

[0060] The second component 120 may include a nitinol alloy having an austenitic phase transition temperature Af slightly above body temperature such that the material remains in its martensitic ductile shape memory phase while in vivo unless and until it is heated to its Af, such as by injecting warm or hot saline (or other fluid) into or flowing through the second component 120, or applying heat by electrical energy (such as with RF energy). In one non-limiting example, the ductile shape memory material of the second component 120 includes nitinol having an austenitic finish temperature (Af) above 37°C. For example, the Af of the nitinol of the ductile shape memory material of the second component 120 may be between 45-65°C, such as between 50-55°C. In some examples, using a side hole catheter positioned through the device 100, warm or hot saline (or other fluid) may be injected sufficiently close to the second component 120 to heat the component to or above its Af. Optionally, an expandable balloon may be disposed at the distal end of the side hole catheter and expanded at the distal opening of the shunt so that blood flow in the shunt is blocked during saline delivery. In a similar example, an expandable balloon may be disposed near the distal end of the side hole catheter and expanded at the proximal opening of the shunt so that blood flow in the shunt is blocked during saline delivery. In yet another example, warm or hot saline may be injected through the distal end of the central lumen of the catheter located adjacent to or near the second component 120. Preferably, the distal end of the catheter includes a hole larger than the above-mentioned side hole, so that heated saline can be delivered more quickly. Optionally, a separate balloon catheter may be inserted through the device 100 so that the balloon is located at the distal side of the catheter for saline delivery. Before injecting saline through the distal end of the catheter, the balloon may be expanded at the distal opening of the device 100 to block blood flow during saline delivery, which may increase the effectiveness of heating. Using a separate catheter for delivering saline and blocking blood flow may allow saline to heat the stent more quickly while keeping the expanded balloon in place.

[0061] In other examples, a pair of electrodes may be contacted with the device 100, for example, via a catheter, and actuated at an appropriate voltage and frequency to heat the component 120 to its Af or above. In yet other examples, any other suitable means may be used to locally apply heat to the device 100, such as laser, magnetic induction, resistance, or the like. Heating the device 100 using resistance may include, for example, contacting the device with a pair of electrodes via a catheter and passing an electric current through the device, thereby heating the device. Heating the device 100 using a laser may include irradiating the device with light from a laser, which may be introduced through a catheter. Heating the device 100 using magnetic induction may include passing an alternating magnetic field through the device, which induces eddy currents in the device that heat the device. Note that in blood vessels (such as the aorta or internal iliac artery) with particularly high blood flow rates (e.g., 2-5 L / min), it may be useful to use direct heating methods such as using electrical energy (e.g., direct current (DC), radio frequency (RF)), laser, magnetic induction, resistance, non-contact radio frequency (RF), dielectric heating, or conductive heating (such as local probe heating) rather than using saline to heat the device 100 before fully heating the device.

[0062] Alternatively, the device 100 may include a single Nitinol alloy that has been heat treated to produce a lower Af in regions corresponding to the first component 110 and the third component 130, respectively, and has been heat treated to produce a higher Af in a region corresponding to the second component 120. The ductile shape memory material of the second component 120 may be expandable and contractible using any suitable technique. For example, the ductile shape memory material of the second component 120 may be mechanically expandable, such as using a balloon as is known in the art. Additionally, or alternatively, the ductile shape memory material of the second component 120 may be heat contracted (e.g., using saline at a temperature at or above the Af of the material), or heated in a manner such as described above (such as with RF energy or using lasers, magnetic induction, electrical resistance, or the like).

[0063] The first component, the second component, and the third component may be coupled to one another using any suitable coupling means (one or more), such as a fluid coupling. For example, any ductile shape memory material may be optionally and independently coupled to any self-expanding superelastic material by welding. Additionally or alternatively, any ductile shape memory material may be optionally and independently coupled to any self-expanding superelastic material using an encapsulant, which may cover at least a portion of at least one of these components and may couple such components to one another. Additionally or alternatively, any shape memory material and any self-expanding superelastic material may be integrally formed with one another by a common frame.

[0064] The encapsulant may include any suitable biocompatible material, such as a polymer or a natural material. Examples of polymers suitable for use as encapsulants include expanded polytetrafluoroethylene (ePTFE), silicone, polycarbonate polyurethane, DACRON (polyethylene terephthalate), ultra-high molecular weight polyethylene (UHMWPE) and polyurethane. Examples of natural materials suitable for use as encapsulants include pericardial tissue (e.g., from horse, cattle or pig sources) or human tissue (such as human placenta or other human tissue). The biocompatible material is preferably smooth to inhibit thrombosis, and optionally can be impregnated with carbon to promote tissue ingrowth. Optionally, in order to promote tissue ingrowth and endothelialization, the biocompatible material can form a mesh structure. The present device can be encapsulated with a biocompatible material in a manner similar to that described in U.S. Patent No. 11,304,831 (the entire contents of which are incorporated herein by reference) entitled "Systems and Methods for Making Encapsulated Hourglass Shaped Stents" by Nae et al. Other methods of encapsulating shunts are described in US Pat. Nos. 10,835,394 to Nae et al., 11,109,988 to Rosen et al., 9,034,034 to Nitzan et al., 9,980,815 to Nitzan et al., and 10,076,403 to Eigler.

[0065] In one example, the device is encapsulated with ePTFE. Those skilled in the art will appreciate that the ePTFE material has a characteristic microstructure consisting of nodes and fibrils, wherein the fibrils are oriented substantially parallel to the longitudinal expansion axis. The expanded polytetrafluoroethylene material can be prepared by ram extruding a compressed billet of granular polytetrafluoroethylene and an extrusion lubricant through an extrusion die to form a sheet or tubular extrudate. The extrudate is then longitudinally expanded to form a node-fibril microstructure and heated to a temperature at or above the crystalline melting point of polytetrafluoroethylene (i.e., 327° C.) for a period of time sufficient to sinter the ePTFE material. The heating can be performed in a vacuum chamber to prevent or inhibit oxidation of the device. Alternatively, the heating can be performed in a nitrogen-rich environment. A heating furnace can be used to heat the encapsulated device. Alternatively, or in addition, a mandrel on which the encapsulated device is placed can be used to heat the encapsulated device.

[0066] Reference now Figure 1B , Figure 1AThe device shown in is modified to add a bridge at the neck region 104, which is configured to engage the patient's atrial septum. The device 100 is preferably encapsulated with a graft material to create a shunt-graft assembly and a pathway to allow blood flow; to better illustrate the location of the bridge 240, Figure 1B Such graft material is omitted from the device. The bridge 240 can be made of a biocompatible material, such as a polymer or natural material as described above, and can be the same material as the material used to encapsulate the frame, or it can be a different material. Preferably, the biocompatible material of the bridge is one that promotes tissue attachment so that if the inner diameter of the shunt is reduced, contact with the septal wall is maintained. Maintaining contact with the septal wall helps prevent any fluid from bypassing the outside of the device. The biocompatible material can be configured to promote tissue ingrowth over the entire bridge or only over a portion of the bridge. For example, holes can be placed at locations on the bridge 240 that are configured to engage the atrial septum, while the remainder of the bridge 240 remains intact so as not to promote tissue ingrowth over the flared end area. This configuration is beneficial to Figures 12A-12D The embodiment shown in is beneficial, wherein the biocompatible material of the bridge 240 extends up to the flared end region and along the interior of the encapsulated shunt.

[0067] Additionally, or alternatively, the bridge 240 may be made of a biocompatible material different from the biocompatible material used to encapsulate the shunt. For example, the shunt may be encapsulated with a biocompatible material (such as ePTFE) with a pore size small enough to slow tissue ingrowth, while the bridge may be made of a biocompatible material with a larger pore size designed to promote tissue ingrowth. In general, the larger the pore size of the biocompatible material, the greater the adhesion of the tissue to the biocompatible material. In addition to promoting tissue growth, the greater porosity allows fluid to be exchanged in and out of the gap between the outer surface of the neck region 104 and the bridge 240. For example, the bridge 240 may be made of ePTFE having an intermodal distance (e.g., approximately 60-200 μm) greater than the ePTFE encapsulating the shunt.

[0068] Alternatively, the bridge 240 may be made of braided Dacron to further promote tissue ingrowth. Stitches may be used rather than the following method to securely attach the Dacron to the encapsulated shunt. Since Dacron is thicker than ePTFE, the cross-section of the device may increase in a crimped or curled configuration, which may mean that a larger diameter sheath may be required for the delivery device. Other materials that may be used to promote tissue ingrowth include the use of mesh structures, electrospun fabrics, or silicones.

[0069] 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 halfway along the first flared end region 102, and the second end 242 is disposed approximately halfway along the second flared end region 106. Optionally, the first end 241 and the second end 242 may extend further along the first flared end region 102 and the second flared end region 106, and may also be attached closer to the neck region 104. The bridge 240 may be stretched such that a gap is created between the outer surface of the neck 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 neck region 104.

[0070] As described above, the encapsulated shunt can be adjusted in vivo to increase or decrease the neck size, thereby adjusting the fluid flow rate through the shunt. Because the encapsulated shunt can be designed to promote tissue ingrowth and endothelialization, tissue may attach to the shunt over time. Therefore, adjusting the encapsulated shunt to increase or decrease the size can cause trauma to the atrial septum. The bridge 240 is designed to prevent cracking and reduce tissue trauma that may be caused by such adjustment. The bridge 240 is configured to engage with the atrial septum and define an outer diameter D4. Preferably, the outer diameter D4 is greater than the neck size D3. In one embodiment, the outer diameter D4 can be 7-9 mm, and the neck size D3 can be 4.5-5.5 mm. When the device is adjusted in vivo, the bridge 240 can be configured to keep the outer diameter D4 unchanged, and only the neck size D3 of the shunt and the size of the gap are modified. As a result of creating a gap between the neck region 104 and the bridge 240, the neck dimension D3 can be reduced or increased up to the outer diameter D4, thereby causing an increase or decrease in the size of the gap without interfering with the spacer tissue contacting and surrounding the bridge 240, and while maintaining contact with the spacer tissue so that leakage or bypass flow around the outer surface of the shunt is minimized.

[0071] The bridge 240 may also be used with an encapsulated shunt that is not adjustable in vivo. Specifically, incorporating the bridge 240 into the encapsulated shunt may be beneficial for patients whose septal defect is enlarged prior to implantation of the device, such as due to a previous transseptal procedure, or for patients with a septal defect that is pre-dilated to be larger than the delivery system required for implantation of the device described herein. For example, for patients with severe mitral regurgitation and poor left ventricular function, it may be clinically necessary to first undergo a repair procedure on the mitral valve, such as mitral annuloplasty via a percutaneous transseptal approach. An interatrial shunt placement is then performed. These mitral valve procedures currently use a 23Fr ID (approximately 8mm outer diameter) guide catheter through the fossa ovalis. After the mitral valve repair, a shunt with the smallest outer diameter that matches the larger aperture defect caused by the previous procedure can be implanted, with the catheter having the smaller diameter required for shunting (e.g., 5.0 to 6.5 mm). Also, during the transseptal procedure, the fossa ovalis is torn, creating a smaller pore size than the fossa. Figure 1A Such a flow diverter may be advantageously used in the case of larger orifice defects as required by the illustrated embodiment.

[0072] Incorporating a biocompatible material bridge into the encapsulated shunt increases the outer diameter of the device, thereby allowing the device to be implanted in an enlarged septal hole without affecting the inner diameter and the rate of fluid flow through the device. In addition, the bridge 240 also allows the inner diameter of the encapsulated shunt to be temporarily increased - for example, during a separate transseptal procedure after device implantation - without disturbing the outer diameter of the neck region, thereby minimizing the risk of septal tissue tearing.

[0073] Reference now Figure 1C , Figure 1D , Figure 1E and Figure 1F , showing additional side views, front views, side cross-sectional views and stereoscopic views, respectively. Figure 1C The shunt 100 is shown encapsulated with a biocompatible material 117 to form a shunt-graft assembly 210 and define a flow path through the shunt-graft assembly 210. Figure 1E As shown, the neck region 104 defines an inner diameter of a passageway through which blood flows, ie, a neck dimension D3. The bridge member 240 surrounds the entire neck region 104 to define an outer diameter D4.

[0074] Reference now Figure 2A and Figure 2B , Figure 1B The cross-sectional side view of the device shows the unfilled and filled gaps between the bridge and the encapsulated shunt. The bridge 240 is attached to the first flared end region 102 at a first end 241 and to the second flared end region 106 at a second end 242. Figures 11A-11D and Figures 12A-12DAn exemplary method of attaching the bridge 240 to the encapsulated shunt is further described. Preferably, the bridge 240 is stretched so that a gap is created between the neck 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. Optionally, the gap 243 may be filled with a body fluid when the shunt-graft 210 is delivered and implanted. This embodiment is advantageous in the case where the size of the shunt is configured to be adjusted in vivo. Specifically, the gap 243 may increase or decrease during such adjustment, while the outer diameter of the neck region (e.g., the bridge) that contacts the atrial septum remains unchanged, thereby minimizing any tissue trauma. For example, the bridge 240 may be made of ePTFE with holes so that when the shunt-graft assembly 210 expands, the biocompatible material disposed in the gap 243 is configured to penetrate the bridge 240.

[0075] like Figure 2B As shown, the gap between the bridge 240 and the encapsulated shunt can 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 size of the shunt-graft assembly 210 is adjusted. This embodiment can be beneficial in the case where the device is designed to be implanted into an enlarged septal hole and the device is not configured to be adjusted in vivo. Specifically, since the inner diameter of the device may not be adjusted, the size of the gap 244 does not need to be adjustable. Therefore, the gap 244 can be filled with a solid material or a low hardness material, such as a hydrogel, which can improve the stability of the device and make the bridge more stable.

[0076] Figure 3A-3E Schematic diagram of the use of Figure 1B Example steps of the device. The shunt-graft assembly 210 can 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 can be crimped to an outer dimension of approximately 4.6 mm, which is the inner dimension of a 14F Cook sheath. For example, Figure 3AA shunt-graft assembly 210 is shown disposed within a sheath 300. As will be appreciated by those skilled in the art, if a Cook sheath of different sizes is used, the shunt-graft assembly 210 can be curled to a smaller or larger outer dimension. In addition, a hydrogel layer can be placed within a gap 244 between the bridge 240 and the shunt-graft assembly 210, which can affect the curling size. The delivery catheter and sheath 300 can be designed as described in U.S. Patent 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 assigned to the assignee of the present application, and the entire contents of each are incorporated herein by reference. The sheath can be placed percutaneously through a blood vessel to a desired location in the human body. When the crimped shunt is pushed out of the sheath 300, the self-expanding super-elastic flared end region springs open to its set configuration, while the ductile shape memory central neck region, for example, expands to a desired shape. Figure 3B The illustrated approach, in which the bridge 240 disposed on the neck region 104 (labeled "B" and corresponding to the second component 120) engages an opening in the human body, remains constrained at or near its crimped size. Depending on the desired direction of blood flow through the shunt-graft assembly 210, the first flared end region 102 and the second flared end region 106 (labeled "A" or "C" and corresponding to the first component 110 or the third component 130) provide an inlet, while the other of the flared ends (labeled "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 atrial septum between the right and left atria, with one of the flared ends extending into the right atrium and the other flared end extending into the left atrium. In some configurations, the flared end in the right atrium is the inlet and the flared end in the left atrium is the outlet, while in other configurations, the flared end in the left atrium is the inlet and the flared end in the right atrium is the outlet. As used herein, "inlet" means the component where blood flows in, and "outlet" means the component where blood flows out (outflow). The specific components that can be used to provide blood flow in and out (outflow), respectively, can be selected based on the condition being treated. For example, in HF, the inlet can be on the left atrium (LA) side, when blood flows from the LA to the right atrium (RA), and LA decompression is ideal. In contrast, in PAH, the intra-atrial pressure gradient is reversed, resulting in flow from R to L and RA decompression, and the inlet is on the RA side.

[0077] The cross-sectional area (and size) of the orifice provided by the malleable shape memory central neck region may be increased or decreased to adjust the flow of fluid through the shunt-graft assembly 210. For example, in a manner such as Figure 3C In the manner shown, the neck region can be expanded by balloon dilation using a balloon 301 (e.g., a 12 mm diameter balloon) that can be fed through the orifice using a wire 302. Preferably, the balloon 301 only expands the neck region to a threshold outer diameter defined by the bridge 240 so that expansion of the neck region does not affect the outer diameter of the device or interfere with the intervening tissue around the bridge 240. For example, the bridge 240 can be sized and shaped to define an outer diameter of 7-14 mm, and the balloon 301 can be configured to expand the neck region up to 9 mm.

[0078] In addition, such as Figure 3D In the manner shown and described above, the neck region can be shrunk by injecting a bolus of hot saline having a temperature above the Af of the stretchable shape memory material (e.g., at 45-65°C) through the distal end of the catheter 303, which can cause the neck region to return to its heat-set size, which can be different from its curled size and is preferably 4.5-5.5 mm. Figure 3D One method of heating the neck region is shown, but other methods may be used. For example, saline may be injected through a side hole in catheter 303. In addition, a second balloon catheter may be inserted through shunt-graft assembly 210 such that the balloon is distal to the saline delivery location and the balloon may be expanded to block blood flow through shunt-graft assembly 210 during saline delivery. In other examples, a pair of electrodes may be positioned, for example, via catheter 303, to contact the shunt-graft assembly and actuated at an appropriate voltage and frequency to heat the neck region to a temperature at or above its Af. In yet other examples, other suitable means (such as laser, magnetic induction, resistance, or the like) may be used to locally apply heat to device 100. Preferably, the neck region recovers its heat-set dimensions without affecting the outer diameter of bridge 240, such as Figure 3E shown.

[0079] For example, heat from the saline can cause the ductile shape memory material to transform into an austenite phase, thereby contracting the neck region back to its curled (or otherwise heat-set) size, after which the neck region cools to body temperature and transforms back to its martensite phase. The saline can be delivered in any suitable manner, such as through a flexible catheter having one or more holes (e.g., an end hole, a side hole, or multiple side holes) through which the hot saline can flow, and the flexible catheter can be placed in the neck region, such as over a guidewire that passes through the neck region. In one non-limiting example, initially, the neck region can have its curled inner dimensions, typically 1-2 mm, such as when initially in Figure 3BWhen deployed in the manner shown. Then, for the second time, the neck region can be expanded to any desired larger size between the curled size and the outer diameter of the bridge 240 using balloon expansion. Then, for the third time, the neck can be shrunk to its heat-set size using hot saline. The heat-set size is determined by the size of the jig used in the heat-setting step of the manufacturing process, and can be approximately the same as the size of the neck region in the curled state, can be smaller than or larger than the size of the neck region in the curled state. The heat-set size can be larger than the size of the catheter used to deliver hot saline, and larger than the deflated size of the expansion balloon, but less than or equal to the minimum expected final shunt size required, such as 4mm. Then, the neck region can be expanded again using balloon expansion to any desired larger size between 4mm and the outer diameter of the bridge 240. Any suitable number of expansions and contractions can be applied to the neck region at any desired time or at separate times to provide a suitable and customized fluid flow through the device for each given patient.

[0080] It will be appreciated that what constitutes an appropriate fluid flow for a given patient may also change over time, and the device may be appropriately adjusted to provide fluid flow, or to properly secure the device within the cavity, as appropriate. It will also be appreciated that the self-expanding superelastic member is not affected by the hot saline injection, and thus will maintain its initial fully expanded size while the shape memory member (in this example, the neck region) is adjusted. Furthermore, alternatively or in addition to the hot saline, alternatives such as those described in reference Figure 1A The described approach uses any suitable method for heating the shape memory material, for example, RF heating or using lasers, magnetic induction, electrical resistance or the like.

[0081] Reference now Figure 4A-4D , and further described in U.S. Pat. No. 11,304,831. Preferably, shunt 100 is at least partially covered with a biocompatible material to create shunt-graft assembly 210. For example, shunt 100 may be covered with a single tube of biocompatible material (e.g., Figure 4A100) to create a shunt-graft assembly 210 having different biocompatible material layers (e.g., two to three biocompatible material layers). Graft tube 216 is a biocompatible material tube having a first graft portion 211, a second graft portion 212, and a third graft portion 213. The length of graft tube 216 is longer than the length of shunt 100, and preferably greater than twice the length of shunt 100. The first graft portion 211 begins at a first end 214 of graft tube 216 and extends to the second graft portion 212. The second graft portion 212 extends between the first graft portion 211 and the third graft portion 213 and is continuously connected to the first graft portion 211 and the third graft portion 213. The third graft portion 213 terminates at a second end 215 of graft tube 216. Figure 5A-Figure 10B The graft tube 216 is shown attached to Figure 4A-4D An exemplary method is shown on a flow splitter 100 of the configuration shown.

[0082] Figure 4B , Figure 4C and Figure 4D They are shown respectively Figure 4A 1 and 2. Cross-sections (cross-sections B, C, and D) of a shunt-graft assembly 210 are shown. The thickness of the graft material is exaggerated in these cross-sectional views to better illustrate the different graft material layers on the shunt 100. As will be appreciated by those skilled in the art, the different cross-sections of the flared end region and the neck region may be greater or less than Figure 4B-4D Preferably, each graft layer is smaller than the thickness of the shunt 100 so that encapsulating the shunt 100 will only minimally increase the size of the flared end region and the neck region. Figure 4B , shows a cross-section B of the shunt-graft assembly 210. Figure 4B As shown, the graft tube 216 may cover the outer surface and the inner surface of the first flared end region 102 of the shunt 100. Specifically, the third graft portion 213 may cover the outer surface of the first flared end region 102, and the second graft portion 212 may cover the inner surface of the first flared end region 102 (i.e., cover the inside of the first flared end region 102). Therefore, two biocompatible material layers may cover the first flared end region 102 of the shunt 100.

[0083] Reference now Figure 4C , shows a cross-section C of the shunt-graft assembly 210. Figure 4CAs shown, the neck region 104 of the shunt 100 may be covered by two biocompatible material layers on the outer surface of the neck region 104 and may be covered by one biocompatible material layer on the inner surface of the neck region 104. Specifically, the first graft portion 211 may first be covered on the outer surface of the neck region 104 and then covered by the third graft portion 213. The second graft portion 212 may be covered on the inner surface of the shunt 100 (i.e., on the inside of the neck region 104). Therefore, three biocompatible material layers may be covered at the neck region 104 of the shunt 100 of the shunt-graft assembly 210.

[0084] Reference now Figure 4D , shows a cross-section D of the shunt-graft assembly 210. Figure 4D As shown, the graft tube 216 covers the outer surface and the inner surface of the second flared end region 106 of the shunt 100. Specifically, the first graft portion 211 can cover the outer surface of the second flared end region 106 and the second graft portion 212 can cover the inner surface of the second flared end region 106 (i.e., cover the interior of the second flared end region 106). Therefore, two biocompatible material layers can cover the second flared end region 106 of the shunt 100 of the shunt-graft assembly 210.

[0085] The biocompatible material layers can be firmly bonded together to form a monolithic biocompatible material layer. For example, the first graft portion 211, the second graft portion 212, and the third graft portion 213 can be sintered together to form a strong, smooth, substantially continuous coating covering the inner and outer surfaces of the stent. Then, as needed, the coating portion can be removed from the selected portion of the stent using, for example, laser cutting or mechanical cutting.

[0086] Figure 5A - Figure 10D generally illustrates the manufacture of Figure 4A-4D An exemplary method of implementing a shunt-graft assembly 210 is shown. Referring now to Figure 5AIn order to attach the first graft portion to the neck region and the second flared end region, the method can start with curling the shunt 100. For example, the shunt 100 can be placed in the funnel 207 and pushed toward the reduced section of the funnel 207 in the funnel 207. The reduced section is preferably the diameter of the neck region of the stent or slightly larger. However, it should be understood that the shunt 100 can be reduced to a different diameter. The shunt 100 can be pushed by a dedicated pusher tool (such as the pusher tool described in Yacoby's U.S. Patent No. 9,713,696). The shunt 100 can be constructed in such a way that when the funnel 207 causes the reduction, the shape of the shunt 100 is deformed so that the first flared end region and the second flared end region are gradually reduced and finally turned inward to the longitudinal axis of the shunt 100, resulting in the shunt 100 having a greatly reduced cross-sectional diameter. It should be understood that the shunt 100 can be optionally compressed into a compressed state using any known compression or curling technology.

[0087] like Figure 5B and Figure 5C As shown, the funnel 207 can have or can be coupled to an introducer tube 208 extending from a narrow side of the funnel 207, which can receive the shunt 100 after the shunt 100 is fully constrained by the funnel 207. The diameter of the introducer tube 208 can be smaller than the diameter of the graft tube. Alternatively, the first end of the graft tube can be expanded to a diameter larger than the diameter of the introducer tube 208 using known expansion techniques (e.g., applying heat to the graft tube). Thus, the introducer tube 208 can be inserted into the first end 214 of the graft tube 216, as shown in FIG. Figure 5C As shown, the reduced diameter shunt 100 can be partially pushed out of the introducer tube 208 and into the graft tube 216 such that the second flared end region 106 and the neck region 104 are pushed into the graft tube 216 .

[0088] Reference now Figure 5D, shows the second flared end region 106 and neck region 104 of the shunt 100 after the introducer tube 208 has been advanced and pushed into the graft tube 216, thereby releasing the compressive force (if any) on the second flared end region 106 and neck region 104. After the inward radial force of the introducer tube 208 is removed, at least the second flared end region 106 of the shunt 100 can radially expand to a diameter greater than the diameter of the graft tube 216, thereby engaging the graft tube 216 along the outer surfaces of the second flared end region 106 and neck region 104 in a manner that allows the graft tube 216 to expand in a stress-free and wrinkle-free manner. In order to increase the elasticity of the graft tube 216 to allow the shunt 100 to expand to its expanded state, heat can be applied to both the graft tube 216 and the shunt 100. For example, hot air can be directed at the graft tube 216 and the shunt 100. In an alternative approach, the shunt 100 may be cooled below its martensite to austenite transformation temperature so that it becomes martensite. The graft tube 216 may be loaded onto the second flared end region 106 and the neck region 104 in this contracted state and allowed to slowly expand as the shunt 100 warms to room temperature or an elevated temperature in a manner that causes the graft tube 216 to expand in a stress-free and wrinkle-free manner.

[0089] Reference now Figure 5E After attaching the first graft portion 211 of the graft tube 216 to the second flared end region 106 and the neck region 104, the shunt 100 can be fully ejected from the introducer tube 208, thereby releasing any compressive forces on the first flared end region 102. Upon ejection from the introducer tube 208, and / or in response to the expansion, the shunt 100 can expand and recover. Figure 5E The expanded state is shown.

[0090] The graft tube 216 may be cut or otherwise manufactured to extend along the shunt 100 to a desired length, starting at the outer surface of the neck region 104 adjacent to the first flared end region 102, along the outer surfaces of the neck region 104 and the second flared end region 106, along the inner surface of the shunt 100 and across the outer surfaces of the first flared end region 102 and the neck region 104, and terminating at the neck region adjacent to the second flared end region 106. Alternatively, the graft tube 216 may be longer than desired and may be cut using known cutting techniques (e.g., micro-scissors, material cutting guillotines, or laser cutters) to extend along the shunt 100 after the reference has been performed. Figure 5A - The method described in Figure 10D is then followed to obtain the desired length.

[0091] Reference now Fig. 6A To attach 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, such as Fig. 6A As shown. Fig. 6A As shown, one end of the graft tube 216 can be everted and guided through the interior of the shunt 100 so that the everted portion of the graft tube 216 is positioned inside the shunt 100. To facilitate this process, a plunger 217 can be used to push the graft tube 216 through the interior of the shunt 100. The plunger 217 can be any tool having a major axis and a width or diameter that is less than the inner diameter of the neck region 104 in the expanded state. In this way, the first graft portion 211 can extend along the outer surface of the second flared end region 106, bend around the end of the second flared end region 106, and travel along the interior of the shunt 100, exiting the opposite side of the shunt 100.

[0092] refer to Figure 6B , in order to engage the remainder of the second graft portion 212 with the inner surface of the shunt 100, the first mandrel portion 218 can be engaged with the shunt 100 and the second graft portion 212. After forming an end shape corresponding to the interior of the first flared end region 102, the first mandrel portion 218 can be gently advanced within the second end 215 of the graft tube 216 and the shunt 100 until the first mandrel portion 218 occupies nearly all of the space within the first flared end region 102. This process can involve guiding the first flared end region 102 onto the first mandrel portion 218 while guiding the second end 215 of the graft tube 216 onto the first mandrel portion 218 so that the second end 215 of the graft tube 216 extends along the first mandrel portion 218 in a close-fitting and wrinkle-free manner. When the first flared end region 102 is properly installed on the first mandrel portion 218, the second end 215 of the graft tube 216 will extend beyond the first flared end region 102, as shown in FIG. Figure 6B In this manner, the second graft portion 212 can be partially engaged with the shunt 100 along the inner surface of the shunt 100 .

[0093] Reference now Fig. 7A and Figure 7B , showing the first mandrel portion 218 ( Fig. 7A ) and the second mandrel portion 219 ( Figure 7B ).like Fig. 8A and Figure 8B As shown, the first mandrel portion 218 and the second mandrel portion 219 can be removably coupled to form a mandrel assembly 220. Referring now to Fig. 7A, showing a side view and a front view of the first spindle portion 218. The first spindle portion 218 can have a first retaining portion 222 and a first body portion 223, wherein the first retaining portion 222 extends from the first body portion 223. The first retaining portion 222 is designed to engage with the shunt-graft assembly 210 and has a shape similar to the first flared end region and / or the neck region, but with slightly smaller dimensions. The first body portion 223 can have a cylindrical shape. It should be understood that the first spindle portion can optionally have only the first retaining portion 222. The first spindle portion 218 can have a receiving portion 221, which is sized and configured to receive the protruding portion 228 of the second spindle portion 219. The receiving portion 221 can extend the entire length of the first spindle portion 218, or can optionally extend only a portion of the first spindle portion 218.

[0094] The first mandrel portion 218 may also optionally have one or more ventilation holes 224 in the first retaining portion 222. The ventilation holes 224 may extend through the outer surface of the first retaining portion 222 and may tunnel through the interior of the first retaining portion 222 and the first body portion 223 to a ventilation inlet 225, which may extend through the surface of the first body portion 223. The size of the ventilation holes 224 is preferably in the range of 0.1-2 mm, but it should be understood that ventilation holes of different sizes may be beneficial. The ventilation holes 224 may facilitate release of the stent-graft assembly 210 after a heat treatment is applied, as described below with reference to Fig. 10A and Fig. 10B Specifically, after the encapsulated support is tightly pressed against the mandrel assembly 220 and the air between the encapsulated support and the mandrel assembly 220 is evacuated, there may be a suction force, making it difficult to remove the encapsulated support. The exhaust inlet 225 can allow air to flow through the exhaust inlet to the exhaust hole 224 to eliminate or reduce the suction effect. It should be understood that a plurality of exhaust holes can be connected to one or more exhaust inlets.

[0095] Reference now Figure 7B , a second mandrel portion 219 is shown. The second mandrel portion 219 can have a second retaining portion 226 and a second body portion 227, wherein the second retaining portion 226 extends from the second body portion 227. The second retaining portion 226 is designed to engage with the shunt-graft assembly 210 and can have a shape similar to the second flared end region and / or the neck region, but with slightly smaller dimensions. The second body portion 227 can have a cylindrical shape. It should be understood that the second mandrel portion can optionally have only the second retaining portion 226.

[0096] 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 can be, for example, a shaft extending from the second retaining portion 226. The protruding portion can be coaxial with the second mandrel portion 219 and can be designed to extend a portion of the length, the entire length, or longer than the length of the first mandrel portion 218. Like the first mandrel portion 218, the second mandrel portion 219 can optionally include one or more drain holes 229 and one or more drain inlets 230.

[0097] Reference now Fig. 8A To constrain the shunt-graft assembly 210 on the spindle assembly 220, the second spindle portion 219 is removably coupled to the first spindle portion 218 by engaging the protruding portion 228 with the receiving portion 221. The second spindle portion 219 can be gently advanced within the second flared end region 106 of the shunt-graft assembly 210 toward the first spindle portion 218 until the second spindle portion 219 occupies nearly all of the space within the second flared end region 106 and the protruding portion 228 is fully received by the receiving portion 221. In this manner, the second graft portion 212 can be fully engaged with the second flared end region 106.

[0098] The protruding portion 228 can be designed to engage with the receiving portion 221 so that the protruding portion 228 and the engaging portion are releasably locked together. Alternatively, the protruding portion 228 can be designed to be frictionally fitted in the receiving portion 221. For example, the protruding portion can be designed to have a gradually increasing diameter, which can result in a friction fit with the receiving portion 221. In this example, the first mandrel portion 218 and the second mandrel portion 219 can be released by forcibly pulling 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 can be releasably locked together, or can be frictionally fitted together using various other well-known techniques. It should also be understood that the protruding portion 228 can instead extend from the first mandrel portion 218, and the receiving portion 221 can instead be formed in the second mandrel portion 219.

[0099] like Fig. 8AAs shown, after the second mandrel portion 219 is engaged with the first mandrel portion 218, thereby constraining the shunt-graft assembly 210 on the mandrel assembly 220, the third graft portion 213 can be separated from the surface of the first mandrel portion 218. For example, the second end 215 of the graft tube 216 can be grasped with a clamp and the second end 215 can be gently pulled over the first flared end region 102 and the neck region 104. In this way, the graft tube 216 can be everted and guided over the first flared end region 102 and the neck region 104. Alternatively, other well-known techniques can be used to separate the third graft portion 213 from the first mandrel portion 218 and attach the third graft portion 213 to the first flared end region 102 and the neck region 104.

[0100] Reference now Figure 8B , the third graft portion 213 can be gently pressed against the first flared end region 102 and the neck region 104 in a manner that reduces or eliminates wrinkles. For example, the second end 215 can be manually stretched toward the neck region 104 using a clamp and allowed to gently contact the first flared end region 102 and the portion of the first graft portion 211 that contacts the neck region 104.

[0101] By placing the third graft portion 213 over the first flared end region 102 and the neck region 104, the graft tube 216 will be attached to the shunt 100 so that the graft tube 216 is Figure 4A-4D The shunt 100 is covered in the manner shown. Figure 8C As shown, attaching the graft tube 216 to the shunt 100 in the manner described above results in three layers of biocompatible graft material covering the neck 104 of the shunt 100. Specifically, as shown in FIG. Figure 8C As shown, the second graft portion 212 covers the inner surface of the neck region 104 of the shunt 100, while the first graft portion 211 and the third graft portion 213 cover the outer surface of the neck region 104 of the shunt 100. Figure 8C In the embodiment of the present invention, the third graft portion 213 overlaps the first graft portion 211 at the neck region 104. Figure 4B and Figure 4D As shown, the first flared end region 102 and the second flared end region 106 of the shunt 100 can be covered by only two layers of biocompatible material. In this way, the second graft portion 212 can extend through the lumen of the shunt 100 through the first end of the first flared end region 102, through the neck region 104, and to the end of the second flared end region 106; the first graft portion 211 extends along the outer surface of the second flared end region 106 and the neck region 104; and the third graft portion 213 extends along the outer surface of the first flared end region 102 and the neck region 104, at least partially overlapping and connected to the first graft portion 211.

[0102] It should be understood that the Figures 5A-8C The graft tube 216 is attached to the shunt 100 to form a shunt-graft assembly 210 having the same three-layer structure at the neck region 104 and a two-layer structure at the first flared end region 102 and the second flared end region 106. It should also be understood that a three-layer structure can be attached to the shunt 100 in a similar manner, but resulting in a three-layer structure appearing at the first flared end region 102, the second flared end region 106, and / or the neck region 104. In another example, a similar method can be used to create a shunt-graft assembly having four layers of biocompatible material in one area and two layers of biocompatible material in other areas. These and other methods and embodiments are further described in U.S. Patent No. 11,304,831.

[0103] Reference Fig. 9A - FIG. 9D , in order to firmly 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 described above, sintering produces a strong, smooth, substantially continuous coating covering the inner and outer surfaces of the stent. This can be accomplished by first applying a flexible sleeve (e.g., Fig. 9A A flexible clamshell 231 is shown covering the shunt-graft assembly 210, applying a compressor 232 ( Fig. 10A In this way, the first graft portion, the second graft portion, and the third graft portion can be joined to each other through the through-wall opening in the shunt 100. Thus, the third graft portion and the second graft portion can be sintered together and joined.

[0104] The flexible sleeve may be tubular and may also be elastic and biocompatible. For example, the flexible sleeve may be Fig. 9A 231 is shown. The flexible clamshell member 231 is preferably made of biocompatible silicone, but may alternatively be made of other biocompatible materials that are elastic. The flexible clamshell member 231 is hollow and may have a uniform thickness. Optionally, the flexible clamshell member 231 may have different thicknesses at different points or in different sections. The flexible clamshell member 231 has a first end 235 and a second end 236, which may meet in a neutral position. Optionally, in the neutral position, there may be a longitudinal gap between the first end 235 and the second end 236. The flexible clamshell member 231 is designed so that the first end 235 and the second end 236 can be separated by pulling the first end 235 and the second end 236 in opposite directions.

[0105] Reference now Fig. 9B , the first end 235 and the second end 236 can be pulled in opposite directions to create a longitudinal gap so that when the shunt-graft assembly 210 is positioned on the mandrel assembly 220, the first end 235 and the second end 236 can be pulled onto the shunt-graft assembly 210. In this way, the flexible clamshell member 231 completely or nearly completely covers the shunt-graft assembly 210. Fig. 9B As shown, the flexible clamshell 231 can be sized so that the flexible clamshell 231 fits tightly over the shunt-graft assembly 210. It should be understood that before the shunt-graft assembly 210 is covered in the clamshell 231, a thin metal layer can be attached to the shunt-graft assembly 210 so that the thin metal layer can be sandwiched between the shunt-graft assembly 210 and the flexible clamshell 231, and can act as a barrier between the shunt-graft assembly 210 and the flexible clamshell 231. This barrier helps prevent contaminants in the clamshell 231 from being transferred into the shunt-graft assembly 210. In one example, the thin metal layer can be aluminum foil. In other examples, the thin metal layer can be titanium, tantalum, or stainless steel, but it should be understood that other metals or alloys can be used. This process has been observed to enhance the compression and sintering process. When the stent-graft assembly is wrapped with a tape (eg, a TFE tape or an ePTFE tape), the tape can also prevent contaminants from being transferred to the shunt-graft assembly.

[0106] The flexible clamshell 231 can be sized so that when it is positioned on the shunt-graft assembly 210, the flexible clamshell 231 applies a compressive force on the shunt-graft assembly 210. The flexible clamshell 231 can be sized and configured to optimize the conformance of the graft tube 216 with the shunt 100, thereby minimizing the gap between the graft tube 216 and the adjacent layers of the struts of the shunt 100. The degree of pressure applied by the flexible clamshell 231 to the shunt-graft assembly 210 can change the internodal distance (IND) of the graft material after sintering, as described in more detail below. The degree to which the flexible clamshell 231 covers or does not cover the shunt-graft assembly 210 can also change 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 can be changed to obtain the desired internodal distance. Optionally, any compressive forces applied by the flexible clamshell 231 may be negligible. Additional compressive forces on the shunt-graft assembly 210 may optionally be achieved by first wrapping the shunt-graft assembly 210 and / or the flexible clamshell 231 with a tape (e.g., a TFE tape or an ePTFE tape). For example, the shunt-graft assembly 210 covered by the flexible clamshell 231 may be placed in a spiral wrapping machine that tension wraps the shunt-graft assembly 210 and the flexible clamshell 231 with at least one overlapping layer of tape, as described in more detail above.

[0107] Reference now Fig. 10A and Fig. 10B , compressor 232 is shown. Compressor 232 has a first half 233 and a second half 234, and also includes a coupler 237 for removably coupling first half 233 to second half 234. Coupler 237 may include a screw with a locking nut or involve any other known technique for removably locking two components together. First half 233 and second half 234 may include a receiving portion 238, which is sized and configured to receive coupler 237. First half 233 and second half 234 have internal recesses configured to receive shunt-graft assembly 210 covered by flexible clamshell member so that when first half 233 and second half 234 are removably coupled together, their respective internal recesses have the shape of shunt-graft assembly 210 covered by flexible clamshell member 231. Optionally, when the first half 233 and the second half 234 are detachably coupled together, their respective internal recesses have the shape of the spindle assembly 220, except that the radial dimension is slightly larger.

[0108] The first half 233 and the second half 234 are rigid and preferably stainless steel, but it should be understood that the first half 233 and the second half 234 can be other rigid materials. The first half 233 and the second half 234 can be designed so that when the first half 233 and the second half 234 are coupled together, the first half 233 and the second half 234 are positioned at a constant distance from the shunt-graft assembly 210. Optionally, the distance from the shunt-graft assembly 210 or the mandrel assembly 220 at different areas of the first half 233 and the second half 234 can vary. The first half 233 and the second half 234 can be designed to have a wall thickness between the inner surface of the first half 233 and the second half 234 and the outer surface of the first half 233 and the second half 234 that allows a desired degree of heat transfer. For example, the first half 233 and the second half 234 can have a thin wall thickness to increase the heat transferred to the stent-graft assembly. Thinner wall thickness can lead to shorter sintering time, which can improve productivity. In addition, shorter sintering time reduces the effect of sintering on the transformation temperature (eg, austenite finish temperature (Af)) of the Nitinol framework.

[0109] Reference now Fig. 10B , when the shunt-graft assembly 210 is mounted on the mandrel assembly 220, the compressor 232 can be attached to the clamshell 231 covering the shunt-graft assembly 210. Specifically, the first half 233 can be positioned on the first half of the shunt-graft assembly 210 and the clamshell 231, and the second half 234 can be positioned on the clamshell 231 and the second half of the shunt-graft assembly 210 so that the first half 233 and the second half 234 completely or nearly completely cover the flexible clamshell 231. Once the receiving portions 238 of the first half 233 and the second half 234 are aligned, the coupler 237 can be inserted into the receiving portion 238 to removably couple the first half 233 to the second half 234 on the clamshell 231. The coupler 237 can be tightened to increase the compressive force applied to the shunt-graft assembly 210. It should be understood that once sintered, the compressive force applied by the coupler 237 can change the inter-nodal distance of the graft material. It should also be understood that the coupler 237 can be tightened to a certain degree to achieve the desired inter-nodal distance.

[0110] After coupling the first half 233 to the second half 234 around the flexible clamshell 231, the compressor 232 will be positioned on the flexible clamshell 231, the flexible clamshell 231 will be positioned on the shunt-graft assembly 210, and the shunt-graft assembly 210 will be positioned on the mandrel assembly 220, as shown in FIG. Fig. 10B23, thereby forming a sintered assembly 239. After the first half 233 is coupled to the second half 234, the compressor 232 applies a compressive force on the flexible clamshell 231 and the shunt-graft assembly 210, thereby pressing the shunt-graft assembly 210 against the mandrel assembly 220. The flexible clamshell 231, due to its elasticity, can promote a uniform distribution of the compressive force on the shunt-graft assembly 210. The shunt 100 may include a plurality of through-wall openings. The force applied by the wrapping tape and / or the compressor 232 presses the shunt-graft assembly 210 against the mandrel assembly 220, thereby bringing the graft layers into close contact through the gaps of the shunt 100.

[0111] It may be desirable for compressor 232 to apply a uniform compressive force. Alternatively, it may be desirable to vary the compressive force applied to shunt-graft assembly 210 at certain points along shunt-graft assembly 210. For example, flexible clamshell 231 may have varying thicknesses and / or lengths, thereby allowing compressor 232 to distribute varying degrees of compressive force on the shunt-graft assembly depending on the wall thickness and geometry of flexible clamshell 231. Additionally, the distance from the inner wall to the surface of shunt-graft assembly 210 at certain points along first half 233 and / or second half 234 may vary. For example, an area where the inner wall of first half 233 is less distant from shunt-graft assembly 210 than the remainder of first half 233 may apply greater compressive force on shunt-graft assembly 210. Varying the compressive force applied to shunt-graft assembly 210 can reduce or increase conformality between first graft portion 211, second graft portion 212, and third graft portion 213. In an alternative embodiment, compressor 232 can be designed such that it applies compressive force only at neck region 104.

[0112] 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 can be firmly bonded together by heating the sintering assembly 239. For example, the sintering assembly 239 can be heated by placing the sintering assembly 239 in a radiant heat heating furnace that can be preheated. Sintering can be performed as discussed in more detail in U.S. Pat. No. 11,304,831. The heated assembly can then be allowed to cool for a period of time sufficient to allow manual manipulation of the assembly. After cooling, the first half 233 and the second half 234 of the compressor 232 can be separated and removed from the flexible clamshell 231. Next, the spiral wrap (if any) can be unwound and discarded. The flexible clamshell 231 can be removed, and the encapsulated stent can 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, can then be placed into a laser trimming fixture to trim away excess graft material, if any. In addition, the graft material encapsulating the stent can also be trimmed at various locations along the stent (such as near one of the ends of the stent) to allow for coupling to a delivery device, such as, for example, Figure 1C shown.

[0113] Figure 4A-4D As shown and using Figure 5A-Figure 10B The resulting structure created by the described method is beneficial because the biocompatible material does not terminate at either end of the shunt-graft assembly 210, but rather terminates at the neck region 104 of the shunt 100. The inventors have discovered that biocompatible material extending beyond the stent at either end can lead to thrombosis when the stent is implanted and can also interfere with attachment to a delivery catheter. The reduction in graft overhang at the edges also improves fluid dynamics at the inlet and outlet of the encapsulated stent. As an additional benefit, the inventors have also observed that the three-layer structure created at the neck region 104 further inhibits tissue ingrowth. It should be understood that with respect to Figure 5A-Figure 10B The foregoing methods described help prevent gaps between the biocompatible layers at the flared end region, which can result in extensive tissue ingrowth, and also help minimize microscopic surface thinning defects (MSTDs), which can result in the attraction of platelet thrombi. Figures 5A-10B The described method provides a high yield and highly reproducible manufacturing process.

[0114] Applicants have also observed that heating a sintered assembly 239 comprising a flexible clamshell composed of silicone as described herein can cause small fragments and / or molecular portions of silicone to adhere to and / or impregnate into the graft tube 216. Applicants have observed that the fragments and / or molecular portions of silicone that adhere to and / or impregnate into the graft tube 216 can further reduce tissue ingrowth when the encapsulated stent is implanted.

[0115] Reference now Figures 11A-11D , showing Figure 4A-4D The first embodiment of the shunt-graft assembly further includes the above-mentioned bridge member. The shunt 100 can be covered with a single biocompatible material tube (such as Figure 4A 106) to create a shunt-graft assembly 210 having different biocompatible material layers (e.g., two to three biocompatible material layers). For example, a graft tube 216 can be attached to the shunt 100 such that the second graft portion 212 extends through the lumen of the shunt 100 through the first end of the first flared end region 102, through the neck region 104, and to the end of the second flared end region 106; the first graft portion 211 extends along the outer surface of the second flared end region 106 and the neck region 104; and the third graft portion 213 extends along the outer surface of the first flared end region 102 and the neck region 104, at least partially overlapping and connected to the first graft portion 211. After the graft tube 216 is attached to the shunt 100, the bridge 240 can be attached to the shunt-graft assembly 210.

[0116] The bridge 240 may be a single tube of biocompatible material similar to the material of the graft tube 216. Figure 1BAs described above, the biocompatible material of the bridge 240 is preferably a biocompatible material that promotes tissue attachment so that if the inner diameter of the shunt is reduced, it will maintain contact with the septal wall. The biocompatible material can be configured to promote tissue ingrowth on the entire bridge or only on a portion of the bridge. For example, holes can be placed in the position of the bridge 240 configured to engage the atrial septum, while the rest of the bridge 240 remains intact so that the tissue is not promoted to grow inward in the flared end area. Holes can be added like this or other treatments can be performed on the biocompatible material before or after the shunt is assembled to promote tissue ingrowth or promote adhesion. Additionally, or optionally, the bridge 240 can be made of a biocompatible material different from the biocompatible material used to encapsulate the shunt. For example, the shunt can be encapsulated with a biocompatible material (such as ePTFE) with a sufficiently small pore size to slow down tissue ingrowth, while the bridge can be made of a biocompatible material with a larger pore size, which is designed to promote tissue ingrowth. Optionally, the bridge 240 may be made of woven Dacron, a mesh structure, an electrospun fabric, or silicone to further promote tissue ingrowth. As further described above, if the bridge is formed with woven Dacron, the Dacron can be securely attached to the encapsulated shunt with stitches rather than the method described below. In some embodiments, the shunt can be encapsulated with ePTFE having a thickness of .002" and an internode distance of <= 30 microns. Rapid Transmural Capillary Ingrowth by Clowes et al. 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 that ePTFE with IND <= 30 microns exhibits low porosity, and applicants' studies have shown that tissue ingrowth is inhibited in shunts encapsulated with ePTFE with IND = 30 microns. On the other hand, the bridge 240 can have a thickness of .002" or .005" and a pore size to increase the potential space between the bridge 240 and the other ePTFE layers for cell and collagen transmural penetration. This can make This can be achieved with larger pore size ePTFE (e.g., ePTFE material with an IND range of 60 to 200 microns). Alternatively, cross-wall penetration can be facilitated by creating patterned or multiple perforations of similar size into a bridge 240 made of conventional low porosity (IND <= 30 microns) before or after it is applied to the diverter. This dedicated perforation process can be performed using, for example, an energy source (such as laser, RF, etc.) 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 resilient to support significant expansion / contraction of the diameter without damaging the fossa ovalis or the frame encapsulation material. In some embodiments, the gap 243 between the bridge 240 and the neck region 104 increases as the shunt neck region 104 contracts. The bridge 240 can be configured to remain engaged with the patient's atrial septum when the neck region contracts. In some embodiments, the porosity of the biocompatible material of the bridge 240 (measured by, for example, its internode distance) is greater than the porosity of the biocompatible material of the shunt frame encapsulation. Thus, the internode distance of the bridge material can be selected to allow tissue ingrowth, while the internode distance of the encapsulation material can be selected to inhibit tissue ingrowth. In some embodiments, the internode distance of the bridge is greater than 30 microns (e.g., in the range of 45-200 microns), while the encapsulation internode distance is less than or equal to 30 microns. In one embodiment, the inter-segment distance of the bridge is 60 microns and the encapsulation inter-segment distance is 30 microns. The biocompatible material of the bridge and the encapsulation biocompatible material may be expanded polytetrafluoroethylene (ePTFE).

[0117] The bridge 240 may be shorter in length than the shunt-graft assembly 210 and longer in diameter than the diameter of the neck region 104. Preferably, the bridge 240 is shaped and sized such that after the bridge 240 is attached to the neck region 104, the first end 241 extends approximately halfway to the first flared end region 102, the second end 242 extends approximately halfway to the second flared end region 106, and a gap 243 is created between the neck region 104 and the bridge 240.

[0118] Fig.11A The cross-sections of the shunt-graft assembly 210 shown (cross-sections B, C, D, E, and F) are respectively Fig. 11B , Fig. 11C , Fig.11D , Fig.11E and Fig.11F The thickness of the graft material is exaggerated in the cross-sectional view to better illustrate the different graft material layers on the shunt 100. Those skilled in the art will appreciate that the different cross-sectional areas of the flared end region and the neck region may be greater or less than Figure 11B-Figure 11F Now refer to Fig. 11B , shows a cross-section B of the shunt-graft assembly 210. Fig. 11BAs shown, the graft tube 216 may cover the outer surface and the inner surface of the first flared end region 102 of the shunt 100. Specifically, the third graft portion 213 may cover the outer surface of the first flared end region 102, and the second graft portion 212 may cover the inner surface of the first flared end region 102 (i.e., cover the inside of the first flared end region 102). Therefore, two layers of biocompatible material may cover the first flared end region 102 of the shunt 100. From the position where the first end 241 of the bridge 240 is connected to the third graft portion 213 and the neck region 104, three layers of biocompatible material may cover the first flared end region 102, for example, as shown in FIG. Fig. 11C shown.

[0119] Reference now Fig.11D , shows a cross-section D of the shunt-graft assembly 210. Fig.11D As shown, three biocompatible material layers may be covered on the outer surface of the neck region 104 of the shunt 100, and one biocompatible material layer may be covered on the inner surface of the neck region 104. Specifically, the first graft portion 211 may be covered on the outer surface of the neck region 104 first, and then covered by the third graft portion 213. In addition, the neck region 104 may also be covered by the bridge 240, and the gap 243 may extend between the third graft portion 213 and the bridge 240. The second graft portion 212 may be covered on the inner surface of the shunt 100 (i.e., covered on the inside of the neck region 104). Therefore, four biocompatible material layers may be covered at the neck region 104 of the shunt 100 of the shunt-graft assembly 210.

[0120] Reference now Fig.11F , shows a cross-section F of the shunt-graft assembly 210. Fig.11F As shown, the graft tube 216 covers the outer surface and the inner surface of the second flared end region 106 of the shunt 100. Specifically, the first graft portion 211 can cover the outer surface of the second flared end region 106, and the second graft portion 212 can cover the inner surface of the second flared end region 106 (i.e., cover the inside of the second flared end region 106). Therefore, two layers of biocompatible material can cover the second flared end region 106 of the shunt 100 of the shunt-graft assembly 210. From the position where the second end 242 of the bridge 240 is connected to the first graft portion 211 and the neck region 104, three layers of biocompatible material can cover the second flared end region 106, for example, Fig.11E shown.

[0121] Can be used with Figure 5A- The bridge 240 is attached to the shunt-graft assembly 210 in a manner similar to the method described in FIG. 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 the correct position when the bridge 240 is attached to the shunt-graft assembly 210. For example, the shunt-graft assembly 210 can be placed within the funnel and advanced toward the narrowing section of the funnel. The shunt-graft assembly 210 can then be inserted into an introducer tube, which can have a smaller diameter than the bridge 240. The introducer tube can then be inserted into the first end 241 of the bridge 240 and can be partially pushed out of the second end 242. The shunt-graft assembly 210 can then be pushed out of the introducer tube such that the second flared end 106 extends past the second end 242, while the first flared end 102 extends in the opposite direction past the first end 241. The bridge 240 can then be positioned over the neck region 104 such that the first end 241 is disposed approximately halfway along the first flared end region 102, the second end 242 is disposed approximately halfway along the second flared end region 106, and a gap 243 is disposed between the neck region 104 and the bridge 240. The gap 243 can be created by manually pulling (e.g., using pliers) the first end 241 toward the first flared end region 102 and the second end 242 toward the second flared end region 106.

[0122] In order to firmly 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, such as Figure 2B For example, before or after the bridge 240 is attached to the neck region 104, a biocompatible material may be attached to the shunt-graft assembly 210 at the neck region.

[0123] Figures 12A-12D Shows Figure 4A-4D Another embodiment of a shunt-graft assembly of the present invention further includes the bridge member described above. As described above, the shunt 100 can be encapsulated to create a shunt-graft assembly 210 having different layers of biocompatible materials, such as two layers on the first flared end region 102 and the second flared end region 106, and three layers on the neck region 104. After the graft tube 216 is attached to the shunt 100, the bridge member 240 can be attached to the shunt-graft assembly 210.

[0124] The bridge 240 may be created using a second graft tube 245, which may be a single tube of 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 neck region 104. The second graft tube 245 may include a fourth graft portion 246, a fifth graft portion 247, and a sixth graft portion 248. Preferably, the fourth graft portion 246 extends over the neck region 104 to create the bridge 240.

[0125] Fig. 12A The cross-sections of the shunt-graft assembly 210 shown (cross-sections B, C, and D) are respectively Fig. 12B , Fig. 12C and Fig.12D The thickness of the graft material is increased in the cross-sectional view to better illustrate the different graft material layers on the shunt 100. Those skilled in the art will appreciate that the different cross-sections of the flared end region and the neck region may be greater or less than Figure 12B-Figure 12D Now refer to Fig. 12B , shows a cross-section B of the shunt-graft assembly 210. Fig. 12B As shown, the graft tube 216 may be covered on the outer surface and the inner surface of the first flared end region 102 of the shunt 100. Specifically, the third graft portion 213 of the graft tube 216 may be covered on the outer surface of the first flared end region 102, and the second graft portion 212 of the graft tube 216 may be covered on the inner surface of the first flared end region 102 (i.e., covered on the inside of the first flared end region 102). In addition, the second graft tube 245 having the fourth graft portion 246, the fifth graft portion 247, and the sixth graft portion 248 may be attached to the graft tube 216. Specifically, the fourth graft portion 246 of the second graft tube 245 may be covered on the outer surface of the first flared end region 102, which may extend on the third graft portion 213 of the graft tube 216, and the fifth graft portion 247 of the second graft tube 245 may be covered on the inner surface of the first flared end region 102, which may extend on the second graft portion 212 of the graft tube 216. At certain portions of the first flared end region 102, the sixth graft portion 248 of the second graft tube 245 may be attached such that it overlaps the fourth graft portion 246. Thus, four or five layers of biocompatible material may be covered on the first flared end region 102 of the shunt 100. As will be appreciated by those skilled in the art, the fourth graft portion 246 may instead overlap the sixth graft portion 248 on the second flared end region 106 such that the first flared end region 102 is covered by only four layers of biocompatible material.

[0126] Reference now Fig. 12C , shows a cross-section C of the shunt-graft assembly 210. Fig. 12C As shown, three layers of biocompatible material may be covered on the outer surface of the neck region 104 of the shunt 100, and two layers of biocompatible material may be covered on the inner surface of the neck region 104. Specifically, the first graft portion 211 of the graft tube 216 may be covered on the outer surface of the neck region 104 first, and then covered by the third graft portion 213 of the graft tube 216. In addition, the neck region 104 may also be covered by the bridge 240 (which may be a part of the fourth graft portion 246 of the second graft tube 245), and the gap 243 may extend between the third graft portion 213 of the graft tube 216 and the bridge 240. The second graft portion 212 of the graft tube 216 and the fifth graft portion 247 of the second graft tube 245 may be covered on the inner surface of the shunt 100 (i.e., on the inside of the neck region 104). Thus, five layers of biocompatible material may cover the neck region 104 of the shunt 100 of the shunt-graft assembly 210 .

[0127] Reference now Fig.12D , shows a cross-section D of the shunt-graft assembly 210. Fig.12D As shown, the graft tube 216 covers the outer surface and the inner surface of the second flared end region 106 of the shunt 100. Specifically, the first graft portion 211 of the graft tube 216 may cover the outer surface of the second flared end region 106, and the second graft portion 212 of the graft tube 216 may cover the inner surface of the second flared end region 106 (i.e., cover the inside of the second flared end region 106). In addition, the second graft tube 245 may be attached to the graft tube 216. Specifically, the fourth graft portion 246 of the second graft tube 245 may cover the outer surface of the second flared end region 106, which may extend on the first graft portion 211 of the graft tube 216, and the fifth graft portion 247 of the second graft tube 245 may cover the inner surface of the second flared end region 106, which may extend on the second graft portion 212 of the graft tube 216. Thus, four layers of biocompatible material may cover second flared end region 106 of shunt 100 of shunt-graft assembly 210 .

[0128] Can be used with Figure 5A- The bridge 240 is attached to the shunt-graft assembly 210 in a manner similar to the method described in FIG. 10D. For example, the shunt-graft assembly 210 can be placed within the funnel and advanced toward the narrowing section of the funnel. The shunt-graft assembly 210 can then be inserted into an introducer tube, which can have a diameter that is smaller than the diameter of the bridge 240. The introducer tube can then be inserted into the first end of the bridge 240 such that all of the second flared end region 106 and the neck region 104 and a portion of the first flared end region 102 are pushed into the second graft tube 245. After the introducer tube is removed, the first flared end region 102 and the second flared end region 106 can be radially expanded to a diameter that is larger than the diameter of the second graft tube 245, thereby engaging the second graft tube 245 along the outer surface of the flared end regions in a manner that allows the second graft tube 245 to expand in a stress-free and wrinkle-free manner. Preferably, the diameter of the second graft tube 245 is greater than the diameter of the neck region 104 such that a bridge 240 is created with a gap 243 between the encapsulated neck region 104 and the second graft tube 245. The bridge 240 can have a first end 241, which can be disposed approximately halfway along the first flared end region 102, and a second end 242, which can be disposed approximately halfway along the second flared end region 106.

[0129] After attaching fourth graft portion 246 of second graft tube 245 to second flared end region 106, neck region 104, and a portion of flared end region 102, one end of second graft tube 245 may be everted and guided through the interior of shunt 100 using a plunger or other tool as described above. Fig. 7A and Figure 7B The illustrated mandrels of the first mandrel portion 217 and the second mandrel portion 219 may be used to ensure that the fifth graft portion 247 is properly engaged with the inner surfaces of the first flared end region 102 , the neck region 104 , and the second flared end region 106 .

[0130] After the fifth graft portion 247 of the second graft tube 245 is attached to the interior of the shunt 100, the second end of the second graft tube 245 can be clamped with a clamp and the second end can be gently pulled onto the first flared end region 102 to attach the sixth graft portion 248. Preferably, the sixth graft portion 248 is attached to the first flared end region 102 so that it partially overlaps the fourth graft portion 246, such as Fig. 12A shown.

[0131] After the sixth graft portion 248 of the second graft tube 245 is attached to the first flared end region, a flexible sleeve similar to the flexible sleeve 231 (described above and as Fig. 9APreferably, the shape of the flexible sleeve is configured to receive the shunt-graft assembly 210 including the bridge 240, and the diameter of the neck region of the flexible sleeve corresponds to the diameter of the bridge 240 so that the gap 243 is not reduced. Fig. 10A The compressor of the compressor 232 in is positioned on the flexible sleeve. Preferably, the compressor has a first half and a second half, each having an internal recess sized and configured to receive the flexible sleeve.

[0132] As described above, in order to firmly bond the second graft tube 245 to the graft tube 216, pressure may be applied to the shunt-graft assembly 210 to achieve sintering. Preferably, the sintering assembly is heated so that the fourth graft portion 246, the fifth graft portion 247, and the sixth graft portion 248 become sintered to the shunt-graft assembly 210, except for the neck region 104, so that a gap 243 is created between the neck region 104 and the bridge 240, which may be unbonded. Additionally or alternatively, the gap 243 may be filled with a biocompatible material, such as Figure 2B For example, before or after fourth graft portion 246 of second graft tube 245 is attached to neck region 104, a biocompatible material may be attached to shunt-graft assembly 210 at the neck region.

[0133] Fig.13 Another embodiment of a shunt-graft assembly including the above-described bridge is shown. In this embodiment, a single tube of biocompatible material can be used to encapsulate the frame and form the bridge. For example, a single tube can be wrapped twice, starting at a location on one end (e.g., the neck of the frame or the end near one flange of the frame) and using three eversions to lay down two complete wraps. Fig.13 In the figure, the first complete wrap "first layer" is shown in blue, and the second wrap "second layer" is shown in pink. In some embodiments, the neck area is sintered before the last eversion of the second layer. In such an embodiment, after the neck is sintered, there will be 3 material layers at the neck, with pink and blue inside and only blue outside, all of which can be sintered together. The final end of the tube can then be everted and pulled back onto the diverter. This layer will form the bridge described herein. In some embodiments, the bridge will not be sintered to the other layers at the smallest diameter portion of the neck so that the bridge remains free in the neck area. The end of the bridge that contacts the first wrapping layer (shown in blue) will be sintered. By changing the gap sintered at the neck, the length and height of the bridge can be selected.

[0134] Reference now Fig.14A and Fig. 14B, showing the use of Figure 7A-Figure 10B and Figures 12A-12D An image of an exemplary encapsulated stent formed by the method described in . Specifically, Fig.14A A first flared end region, a second flared end region, and a neck region of the encapsulating stent are shown, wherein a bridge surrounds the neck region to create a larger outer diameter, the bridge being configured to engage the atrial septum. Fig. 14B A first flared end region and a neck region are shown, which define the inner diameter (or neck size) of the passageway through which blood flows.

[0135] As described above, the neck region of the room shunt can be expanded and / or contracted in vivo to provide a suitable and customized fluid flow through the device for each given patient, such as further described in U.S. Patent No. 10,898,698. Specifically, the frame (e.g., metal frame) of the shunt device described herein can be heat treated during the manufacturing process so that the neck region of the shunt can be ductile (e.g., mechanically expandable) at body temperature and contract when heated to a temperature above its martensite finish temperature (Mf). In addition, the shunt device can also be heat treated so that the neck region can contract from the expanded state to a variable degree depending on the temperature to which the frame of the shunt device is raised - for example, contracting by a selected increment. For example, the shunt orifice at the neck region of the shunt device can be expanded in vivo from a first diameter to a second diameter greater than the first diameter (e.g., mechanical expansion via a balloon catheter), and thermally contracted from the second diameter to a third diameter greater than the first diameter in vivo by heating the frame of the shunt device to a predetermined temperature above its martensite finish temperature (Mf) but below its austenite finish temperature (Af). Thus, the multiphase diverter devices described herein may be contracted to a desired diameter greater than their heat-set austenitic configuration without necessarily being fully contracted—for example, by heating the diverter device to a temperature above its austenite finish temperature (Af) and then mechanically expanding it to the desired diameter.

[0136] Reference now Fig.15 , provide diagrams showing exemplary gradual incremental contraction of the shunt orifice of a bare frame shunt device and an encapsulated shunt device, respectively. Referring to Table 1 provided below, as Fig.15As shown, during the baseline phase of the respective shunt devices, in a water bath at body temperature (e.g., 37°C), the orifice diameter of the bare frame shunt device was 5.00 mm, while the orifice diameter of the encapsulated shunt device was 4.80 mm. In this study, the austenite finish temperature (Af) of each shunt device was approximately 60°C. When the temperature of the shunt device was maintained at 37°C, both shunt devices were expanded via a nominal 9mm angioplasty balloon, so that the shunt orifice of the bare frame shunt device was expanded to 9.10 mm, while the shunt orifice of the encapsulated shunt device was expanded to 8.65 mm. After expansion, the temperature of the water bath was slowly increased until a temperature of 70°C was reached, so that the shunt orifice diameter of each shunt device was recorded in increments of 5°C. Table 1

[0137] As shown in Table 1, the data demonstrate that heat-activated shape memory recovery of the respective shunt devices to their heat-set baseline shunt orifice diameters can be achieved gradually, as opposed to either full transformation or no transformation. Thus, by heating the shunt devices to a predetermined temperature above their respective martensite finish temperature (Mf) but below their respective austenite finish temperature (Af), the shunt orifices of the shunt devices can be selectively heat-contracted to a desired diameter between their respective baseline diameters and their respective fully balloon-expanded diameters.

[0138] like Fig.15 As shown, both shunt devices begin their recovery (e.g., shrinking from their respective 9 mm balloon-expanded diameters to their respective baseline diameters) at about 45°C, reach half recovery (e.g., about 7 mm shunt orifice diameters) between about 52°C and 55°C, and fully recover (e.g., their respective baseline diameters) at about 60°C. In addition, after the temperature of the water bath exceeds the austenite finish temperature (Af) of about 60°C, further shrinkage of the shunt devices can be minimal because the shunt orifices have substantially reached their respective baseline diameters. As will be appreciated by one of ordinary skill in the art, the shunt devices having the bridges described herein can also be heat treated so that they can be selectively heat shrunk to a diameter different from their heat-set austenite configuration.

[0139] Thus, the shunt devices described herein may be temporarily balloon-expanded, for example, in vivo after deployment at an atrial shunt to allow the device (such as, for example, The shunt device is then effectively reconditioned to a smaller desired diameter (greater than its fully recovered diameter) by heating the shunt device to a specified temperature (e.g., according to Table 1 above) below its austenite finish temperature (Af). For example, the shunt device can be heated by injecting a heated fluid (e.g., saline) into the shunt device via a catheter in vivo until the desired temperature of the shunt device is reached. Preferably, the saline can be heated to an initial temperature - the 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 fluid lumen of the catheter, and / or the outer diameter of the catheter - so that when the saline is injected onto the shunt device at the atrial septum, the saline is at a desired temperature to account for any heat loss between the catheter lumen and external factors (such as ambient air and / or body temperature blood surrounding the catheter) when the saline is injected through the catheter.

[0140] In addition, it is preferred to minimize the dilution and loss of heated saline when it is injected into the shunt device in vivo. For example, an expandable balloon (e.g., a hemispherical balloon with a diameter of 10-16 mm) can be delivered and positioned near the distal region of the shunt device (e.g., across the atrial septum), and expanded to its expanded state to block blood flow through the shunt device during the injection of heated saline into the shunt device (e.g., via an 8-14Fr large injection catheter with an outlet / port adjacent to the proximal region of the shunt device), thereby minimizing the dilution and loss of heated saline. The large injection catheter allows for rapid delivery of heated saline, which further minimizes dilution and loss. Therefore, a certain amount of heated saline must be injected through the catheter to first displace the entire volume in the fluid cavity of the injection catheter, and sufficient heated saline must be delivered to heat the shunt device to the desired temperature in vivo. Alternatively, an expandable balloon can be delivered and positioned near the proximal region of the shunt device, and expanded to its expanded state during the injection of heated saline to minimize the dilution and loss of heated saline. For example, a single injection catheter with an inflatable balloon disposed thereon at a location proximate to the injection port of the injection catheter may be used to simultaneously inject heated saline into the shunt device and block blood flow through the shunt device.

[0141] In some embodiments, the distal tip of the injection catheter may include a temperature sensor that is configured to generate one or more signals indicative of the temperature of the heated fluid delivered within the shunt device (e.g., during injection of heated saline). Thus, when the temperature of the heated fluid delivered within the shunt device reaches a desired temperature, and thus the shunt device reaches a desired shunt orifice diameter, the user can terminate the injection of heated saline through the catheter. Additionally or alternatively, injection of heated saline into the shunt device through the catheter can be performed automatically via a power injector that is fluidly coupled to the injection catheter and operably coupled to the temperature sensor. For example, the power injector can be programmed to automatically terminate the heated saline injection when a signal received from the temperature sensor indicates that the temperature of the heated fluid delivered within the shunt device is at a predetermined desired temperature (e.g., for a predetermined period of time).

[0142] In some embodiments, the power injector may include a heated syringe, the temperature of which may be set to be above the austenite finish temperature (Af) of the shunt device at the atrial 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 within the shunt location may be increased by increasing the injection rate, or conversely, may be decreased by decreasing the injection rate because cooling of the heated saline through the catheter is reduced by reducing the transit time of the heated saline through the catheter.

[0143] Thus, the power injector can be programmed to begin injecting heated saline through the catheter at a rapid rate to initially displace the body temperature fluid within the catheter, and then inject at a slower injection rate as the temperature at the distal tip of the catheter measured by the temperature sensor increases. The power injector can also be programmed to adjust the injection rate to maintain the temperature of the heated fluid delivered within the shunt device, and therefore maintain the temperature of at least the middle region of the shunt device for a predetermined period of time (e.g., one or two seconds or less) to ensure that the shunt orifice diameter has heat shrunk to the desired size.

[0144] Although the middle region of the diverter device is described herein as being heat treated so that it can be heated to a target temperature above its martensite finish temperature (Mf) but below its austenite finish temperature (Af) to selectively shrink to a predetermined size, as will be understood by one of ordinary skill in the art, the proximal and / or distal regions of the diverter device may also be similarly heat treated in addition to or in lieu of the middle region to allow for selective heat shrinkage of the corresponding end regions.

[0145] Other studies have shown that the "effective" Austenite finish temperature (Af) of an ePTFE encapsulated shunt - that is, the temperature required for the shunt device to return to its heat-set austenite configuration - can be lower than the Austenite finish temperature (Af) of the nitinol frame, for example, about 5°C lower. This is theoretically due to the ePTFE acting as an insulator for heat transfer to the nitinol frame. Therefore, for an encapsulated shunt, in order to achieve a target effective Austenite finish temperature (Af) of, for example, 45-60°C, the nitinol frame can be heat treated to exhibit an Austenite finish temperature (Af) of 50-65°C before encapsulation. In addition, the thickness of the encapsulation material can be selected to control the effective Austenite finish temperature (Af) of the encapsulated shunt. For example, a first encapsulated shunt having a nitinol frame having a first austenite finish temperature (Af) and a first encapsulating material thickness may have the same effective austenite finish temperature (Af) as a second encapsulated shunt having a nitinol frame having a second austenite finish temperature (Af) higher than the first second austenite finish temperature (Af) of the first encapsulated shunt and a second encapsulating material thickness thicker than the first encapsulating material thickness of the first encapsulated shunt.

[0146] Reference now Fig.16 , an exemplary method 1600 for treating a cardiac condition using a shunt device described herein is provided. At step 1602, the shunt may be delivered to the atrial septum in its contracted state using any of the above-described delivery methods. For example, the shunt device may be delivered through the atrial septum within a delivery sheath. As described above, the shunt device may have a diabolo shape, and may be heat treated during manufacturing so that its proximal and distal regions are self-expandable at body temperature, and its middle / neck region exhibits a martensite finish temperature (Mf), making it ductile at body temperature, and exhibits an austenite finish temperature (Af), above which it recovers its heat-set austenite configuration. In addition, the shunt device may also be any of a bare frame stent, an encapsulated shunt device, and / or may include a bridge, as described above.

[0147] At step 1604, the shunt device can be deployed at the atrial septum to its expanded state such that its distal region self-expands in the first atrium, its proximal region self-expands in the second atrium, and the neck region is disposed within the atrial septum. At step 1606, the neck region of the shunt device can be expanded (e.g., mechanically balloon-expanded) such that the shunt orifice at the neck region expands from a first diameter to a second diameter that is greater than the first diameter. For example, the second diameter can be selected to allow passage of a medical device therethrough, such as Optionally, the second diameter may be selected to allow blood to flow therethrough at a predetermined flow rate to provide initial treatment.

[0148] Thus, if / when the shunt orifice of the shunt device needs to be reduced, for example during a treatment based on a patient's physiological response, at step 1608, the shunt device can be heated to a predetermined temperature below the austenite finish temperature (Af) (e.g., according to Table 1 above) to selectively heat shrink the shunt orifice at the neck region to a predetermined third diameter - for example, a diameter that is smaller than the second diameter but larger than the first diameter. For example, as described above, an injection catheter can be used to inject a heated fluid (e.g., saline) into the shunt device in vivo to heat the shunt device to a desired temperature. When the shunt device reaches the desired temperature, and therefore the shunt orifice at the neck region reaches the desired size, the delivery sheath and injection catheter can be removed from the patient, leaving the shunt device implanted at the atrial septum. Additionally or alternatively, other methods can be used to selectively heat shrink the shunt device in vivo, such as, for example, applying an electric current to the shunt device via one or more electrodes as described above. As will be appreciated by one of ordinary skill in the art, if the desired diverter orifice size is a heat set austenite configuration (eg, first diameter), the diverter device may be heated to a temperature above the austenite finish temperature (Af).

[0149] Although various illustrative embodiments of the present invention have been described above, it will be apparent to those skilled in the art that various changes and modifications may be made herein without departing from the present invention. For example, although examples of the present device are described as having two or three components, it should be understood that the present device may include any suitable number of components that include a self-expanding superelastic material or a ductile shape memory material, respectively. The appended claims are intended to 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 first atrium and a second atrium of a patient to treat a medical condition, the device comprising: an encapsulated shunt including a first flared end region, a second flared end region, and a neck region disposed therebetween, the encapsulated shunt being formed from a frame encapsulated in a biocompatible material, the encapsulated shunt defining a passageway to allow blood to flow from the first atrium to the second atrium via the passageway; and A bridge extends from a first exterior surface of the first flared end region to a second exterior surface of the second flared end region, the bridge being made of a biocompatible material and configured to engage an atrial septum of the patient.

2. The device of claim 1, wherein the neck region is expandable in vivo such that the passageway expands from a first diameter to a second diameter greater than the first diameter.

3. The device of claim 2, wherein the neck region is collapsible in vivo such that the passageway collapses from the second diameter to a third diameter that is smaller than the second diameter.

4. The device of claim 1, wherein the neck region is ductile at body temperature and comprises Nitinol having an austenite finish temperature (Af) between 45-65°C, and Wherein the first flared end region and the second flared end region are superelastic at body temperature and comprise Nitinol having an austenite finish temperature (Af) between 5-20°C.

5. The device of claim 1, wherein the neck region is mechanically expandable.

6. The device of claim 1, wherein the neck region is heat shrinkable.

7. The device of claim 3, wherein the bridge defines an outer diameter that is greater than the first diameter, the second diameter, and the third diameter of the passageway.

8. The device of claim 7, wherein the outer diameter is between 7 and 14 mm.

9. The device of claim 7, wherein the bridge extends to form a gap between an inner surface of the bridge and an outer surface of the neck region.

10. The device according to claim 9, wherein: When the neck region is expanded in vivo, the gap is configured to decrease in size and the outer diameter is configured to remain constant in diameter.

11. The device of claim 1 , wherein there is no gap between the bridge and the neck region.

12. The device of claim 1, wherein the bridge is integrally formed with the biocompatible material encapsulating the encapsulated shunt.

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

14. The device of claim 13, wherein the biocompatible material of the bridge is configured to allow tissue ingrowth and the encapsulating biocompatible material is configured to inhibit tissue ingrowth.

15. The device of claim 13, wherein the biocompatible material of the bridge comprises a porosity selected to allow tissue ingrowth.

16. The device of claim 1, wherein the inter-nodal distance of the biocompatible material of the bridge is greater than the inter-nodal distance of the encapsulating biocompatible material.

17. The device of claim 16, wherein the inter-nodal distance of the bridge material is selected to allow tissue ingrowth and the inter-nodal distance of the encapsulating material is selected to inhibit tissue ingrowth.

18. The device of claim 16, wherein the inter-nodal distance of the bridges is greater than 30 microns and the inter-nodal distance of the encapsulation is less than or equal to 30 microns.

19. The device of claim 18, wherein the inter-nodal distance of the bridge material is in the range of 45-200 microns.

20. The device of claim 16, wherein the inter-node distance of the bridge material is 60 microns and the inter-node distance of the encapsulation material is 30 microns.

21. The device of claim 16, wherein the biocompatible material of the bridge and the encapsulating biocompatible material is expanded polytetrafluoroethylene (ePTFE).

22. The device of claim 1, wherein the bridge is configured to remain engaged with the patient's atrial septum when the neck region is retracted.

23. A method for diverting blood between a first atrium and a second atrium of a patient to treat a medical condition, the method comprising: The encapsulated shunt was percutaneously advanced toward the atrial septum in the contracted delivery state; and deploying the encapsulated shunt within the perforation of the atrial septum such that the encapsulated shunt transitions to an expanded state wherein a first flared end region is disposed in the first atrium, a second flared end region is disposed in the second atrium, a neck region is disposed therebetween, and a bridge of biocompatible material extending from a first outer surface of the first flared end region to a second outer surface of the second flared end region engages the patient's atrial septum, Wherein the enclosed shunt defines a passageway to allow blood to flow from the first atrium to the second atrium via the passageway.

24. The method of claim 23, wherein the first flared end region and the second flared end region comprise a self-expanding superelastic material and the neck region comprises a ductile shape memory material having a first diameter.

25. The method of claim 24, further comprising expanding the neck region to a second diameter greater than the first diameter.

26. The method of claim 25, wherein dilating the neck region comprises inflating a balloon within the neck region.

27. The method of claim 25, further comprising contracting the neck region to a third diameter that is smaller than the second diameter.

28. The method of claim 27, wherein shrinking the neck region comprises heating the neck region, and The heating comprises flowing heated saline through the enclosed shunt or applying radio frequency (RF) energy to the enclosed shunt.

29. The method of claim 27, wherein the bridge defines an outer diameter that is greater than the first, second, and third diameters of the neck region.

30. The method of claim 23, further comprising a gap between the bridge and the neck region.

31. The method of claim 30, wherein the bridge is configured such that the outer diameter does not change when the neck region is expanded or contracted.

32. The method of claim 23, wherein the bridge is configured such that the bridge remains in contact with the atrial septum when the neck region is dilated or contracted.

33. A method for manufacturing a flow diverter, the method comprising: providing a frame comprising a first flared end region, a second flared end region, and a neck region disposed therebetween; encapsulating the framework with a biocompatible material; and A biocompatible material bridge is coupled to the encapsulated frame such that the bridge extends between a first exterior surface of the first flared end region and a second exterior surface of the second flared end region.

34. The method of claim 33, wherein the biocompatible material used to encapsulate the framework and the biocompatible material used for the bridge are formed from a single graft tube.

35. The method of claim 33, wherein the biocompatible material of the bridge is different from the biocompatible material of the encapsulation.

36. The method of claim 35, wherein the biocompatible material of the bridge is configured to allow tissue ingrowth and the encapsulating biocompatible material is configured to inhibit tissue ingrowth.

37. The method of claim 35, wherein the biocompatible material of the bridge comprises a porosity selected to allow tissue ingrowth.

38. The method of claim 33, wherein encapsulating the framework with a biocompatible material comprises: positioning at least the second flared end region of the frame within the first end of the graft tube, thereby attaching a first layer of graft material to at least the second flared end region of the frame; directing the second end of the graft tube into the second flared end region of the frame and out of the first flared end region of the frame to thereby attach a second layer of graft material along an interior of the frame; and The second end of the graft tube is directed over at least the first flared end region of the frame, thereby attaching a third layer of graft material over at least the first flared end region of the frame to form a shunt-graft assembly.

39. The method of claim 38, wherein encapsulating the framework with a biocompatible material further comprises: directing the first flared end region over a first portion of a mandrel such that the second end of the graft tube extends beyond the first flared end region; coupling a second portion of the mandrel to the first portion of the mandrel; and When the shunt-graft assembly is positioned on the mandrel, a flexible sleeve shaped to receive the encapsulated frame is positioned over the shunt-graft assembly.

40. The method of claim 39, wherein encapsulating the framework with a biocompatible material further comprises: positioning a compressor having a first half and a second half about the flexible sleeve covering the shunt-graft assembly, the first half and the second half each including an internal recess sized and configured to receive the flexible sleeve; and coupling the first half of the compressor with the second half of the compressor when the first half and the second half of the compressor are positioned around the flexible sleeve covering the shunt-graft assembly, wherein the compressor, the flexible sleeve, the shunt-graft assembly and the mandrel form a sintered assembly.

41. The method of claim 40, wherein encapsulating the framework with the biocompatible material further comprises heating the sintering assembly so that the first graft material layer, the second graft material layer, and the third graft material layer become sintered together to form a single graft material layer, thereby forming the encapsulated framework.

42. The method of claim 33 further comprising, after encapsulating the frame, directing a second graft tube outside of the encapsulated frame so that a first end of the second graft tube is attached to at least a portion of the first flared end region of the encapsulated frame and a second end of the second graft tube is attached to at least a portion of the second flared end region of the encapsulated frame.

43. The method of claim 33, wherein combining the biocompatible material bridge with the encapsulated frame comprises: positioning the second flared end region, the neck region, and at least a portion of the first flared end region of the frame within the first end of the second graft tube, thereby attaching a first second graft material layer to the second flared end region, the neck region, and at least a portion of the first flared end region of the frame and forming the bridge; directing the second end of the second graft tube into the second flared end region of the frame and out of the first flared end region of the frame to thereby attach a second layer of second graft material along an interior of the frame; and The second end of the second graft tube is directed over at least a portion of the first flared end region of the frame, thereby attaching a third second graft material layer over at least a portion of the first flared end region of the frame.

44. The method of claim 43, wherein combining the biocompatible material bridge with the encapsulated frame further comprises: directing the first flared end region onto a first portion of a mandrel such that the second end of the second graft tube extends beyond the first flared end region; coupling a second portion of the mandrel to the first portion of the mandrel; and When the shunt-graft assembly is positioned on the mandrel, the positioning shape is configured to receive the encapsulated frame and the flexible sleeve of the bridge, wherein the diameter of the neck region of the flexible sleeve corresponds to the diameter of the bridge.

45. The method of claim 44, wherein combining the biocompatible material bridge with the encapsulated frame further comprises: positioning a compressor having a first half and a second half about the flexible sleeve covering the shunt-graft assembly, the first half and the second half each including an internal recess sized and configured to receive the flexible sleeve; and coupling the first half of the compressor with the second half of the compressor when the first half and the second half of the compressor are positioned around the flexible sleeve covering the shunt-graft assembly, wherein the compressor, the flexible sleeve, the shunt-graft assembly and the mandrel form a sintered assembly.

46. ​​The method of claim 45, wherein bonding the biocompatible material bridge to the encapsulated framework further comprises heating the sintering assembly so that the first, second, and third second graft material layers become sintered to the encapsulated framework except for an outer surface of the neck region, such that an unbonded gap is created between the neck region and the bridge.

47. The method of claim 43, wherein the diameter of the second graft tube is greater than the diameter of the neck region.

48. A method according to claim 33, wherein the bridge member includes a first end and a second end, the first end is configured to extend to approximately halfway of the first flared end region, and the second end is configured to extend to approximately halfway of the second flared end region, thereby creating a gap between the neck region and the bridge member.

49. A shunt for implantation in the atrial septum to treat a heart condition, the shunt comprising: a single metal frame comprising a proximal region, a distal region, and a middle region disposed therebetween, the proximal region and the distal region each being configured to self-expand after deployment such that the middle region is positioned in the opening of the atrial septum, the single metal frame being heat treated such that the middle region is adjustable in vivo between a first diameter and a second diameter greater than the first diameter, The intermediate region is selectively heat shrinkable from the second diameter to a third diameter greater than the first diameter.

50. A diverter according to claim 49, wherein the intermediate region is heat treated to exhibit a martensite finish temperature and an austenite finish temperature higher than the martensite finish temperature, and wherein the intermediate region is capable of selective heat contraction by being heated to a temperature between the martensite finish temperature and the austenite finish temperature.

51. The shunt of claim 49 wherein the intermediate region is ductile at body temperature and comprises Nitinol having an austenite finish temperature between 45-65°C.

52. The shunt of claim 49 wherein the proximal region and the distal region are superelastic and comprise Nitinol having an austenite finish temperature between 5-20°C.

53. The shunt of claim 49 wherein the intermediate region is heat treated to exhibit shape memory properties that are different from the proximal region and the distal region.

54. The shunt of claim 49, wherein the intermediate region comprises a plastically deformable material such that the intermediate region is configured to be expanded in vivo via mechanical expansion.

55. The shunt of claim 49, wherein the proximal and distal regions and the intermediate region define a diabolo-shaped shunt.

56. The shunt of claim 49, further comprising a biocompatible material coating the proximal region, the distal region, and the intermediate region to define a pathway that allows blood to flow through the atrial septum via the shunt.

57. A flow splitter according to claim 56, wherein the cross-sectional flow area in the passage is smallest at the intermediate region.

58. The shunt of claim 56, wherein the passageway is sized and shaped to allow a sufficient amount of blood to flow through the atrial septum via the shunt to treat pulmonary hypertension.

59. The shunt of claim 56, wherein the passageway is sized and shaped to allow a sufficient amount of blood to flow through the atrial septum via the shunt to treat heart failure.

60. The flow diverter of claim 49, further comprising: a bridge extending from a first outer surface of the proximal region to a second outer surface of the distal region, the bridge formed of a biocompatible material and configured to engage the atrial septum, Wherein the bridge defines an outer diameter greater than the second diameter.

61. A shunt according to claim 60, wherein the bridge extends to form a gap between the inner surface of the bridge and the outer surface at the middle region, so that when the middle region is expanded in vivo, the gap is configured to decrease in size and the outer diameter is configured to maintain a constant diameter.

62. A shunt according to claim 60, wherein the bridge extends to form a gap between the inner surface of the bridge and the outer surface at the middle region, so that when the middle region is contracted in vivo, the gap is configured to increase in size and the outer diameter is configured to maintain contact with the gap.

63. A system for treating a cardiac condition, comprising a shunt according to claim 49, the system further comprising a catheter configured to inject a heated fluid into the shunt in vivo to heat the shunt in vivo, thereby selectively heat shrinking the intermediate region from the second diameter to the third diameter.

64. A system according to claim 63, wherein the heated fluid has an initial fluid temperature, which is selected to reach a target fluid temperature at the shunt in the body based on at least one of the injection rate of the heated fluid or the size of the catheter, thereby heating at least the middle region of the shunt to the target shunt temperature.

65. The system of claim 63, wherein the catheter comprises 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 within the shunt.

66. The system of claim 65 further comprising a power injector coupled to the conduit fluid to inject a heated fluid through the conduit, the power injector being operably coupled to the temperature sensor and programmed to adjust at least one of a rate or duration of injection of the heated fluid through the conduit in response to the signal.

67. A system according to claim 66, wherein the power injector is programmed to automatically terminate injection of heated fluid through the catheter when the temperature of the heated fluid delivered within the shunt reaches a predetermined temperature sufficient to selectively heat shrink the intermediate region of the shunt to the third diameter.

68. The system of claim 66 wherein the power injector is programmed to regulate the rate at which heated fluid is injected through the conduit to maintain the temperature of the heated fluid delivered within the shunt for a predetermined period of time.

69. The system of claim 63, wherein the catheter is sized and shaped to inject the heated fluid at a selected injection rate to minimize dilution and loss of the heated fluid within the body.

70. The system of claim 69, further comprising an inflatable balloon configured to be disposed in an expanded state proximate the distal region during injection of the heated fluid to minimize dilution and loss of the heated fluid within the body.

71. A method for treating a cardiac condition, the method comprising: delivering a shunt into an opening of the atrial septum in a collapsed state, the shunt comprising a frame having a proximal region and a distal region and an intermediate region disposed therebetween, the frame being heat treated so that the intermediate region is adjustable in vivo between a first diameter and a second diameter greater than the first diameter; deploying the shunt within the opening to an expanded state, wherein the proximal region self-expands within the first atrium, the distal region self-expands within the second atrium, and the intermediate region is disposed at the atrial septum; expanding the intermediate region of the shunt from the first diameter to the second diameter in vivo; and The intermediate region of the shunt is selectively heat shrunk in vivo from the second diameter to a third diameter greater than the first diameter.

72. The method of claim 71, wherein expanding the intermediate region of the diverter comprises mechanically expanding the intermediate region of the diverter.

73. A method according to claim 71, wherein the intermediate region is heat treated to exhibit a martensite finish temperature and an austenite finish temperature greater than the martensite finish temperature, and wherein selectively heat shrinking the intermediate region of the diverter includes heating the intermediate region to a temperature between the martensite finish temperature and the austenite finish temperature.

74. A method according to claim 71, wherein selectively thermally shrinking the middle region of the diverter includes injecting a heating fluid into the diverter via a catheter in vivo to heat at least the middle region of the diverter, the heating fluid having an initial fluid temperature, and the initial fluid temperature is selected to reach a target fluid temperature at the diverter in vivo based on at least one of an injection rate of the heating fluid or a size of the catheter, thereby heating the middle region of the diverter to a target diverter temperature.

75. The method of claim 74, further comprising: measuring the temperature of the heated fluid delivered within the shunt via a temperature sensor disposed at the distal tip of the catheter; and At least one of a rate or duration of heated fluid injection through the catheter is adjusted based on a measured temperature of the heated fluid delivered within the shunt.

76. The method of claim 74, further comprising: measuring the temperature of the heated fluid delivered within the shunt via a temperature sensor disposed at the distal tip of the catheter; and Injection of heated fluid through the catheter is automatically terminated when the temperature of the heated fluid delivered within the shunt reaches a predetermined temperature sufficient to selectively heat shrink the intermediate region of the shunt to the third diameter.

77. The method of claim 74, further comprising: measuring the temperature of the heated fluid delivered within the shunt via a temperature sensor disposed at the distal tip of the catheter; and The rate at which the heated fluid is injected through the conduit is adjusted to maintain the temperature of the heated fluid delivered within the shunt for a predetermined period of time.

78. The method of claim 74 further comprising expanding an inflatable balloon adjacent the distal region of the shunt to an expanded state to block blood flow through the shunt during injection of the heated fluid, thereby minimizing dilution and loss of the heated fluid in vivo.

79. The method of claim 71 further comprising allowing a sufficient amount of blood to flow through the atrial septum via the shunt to treat pulmonary hypertension or heart failure.

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