Apparatus and methods are described including a left-ventricular assist device that includes a pump-outlet tube shaped to define one or more blood-outlet openings and configured for insertion into a subject's left ventricle, such that the blood-outlet openings are disposed within the subject's aorta and a distal portion of the pump-outlet tube is disposed within the left ventricle. An impeller is disposed within the pump-outlet tube and is configured to pump blood of the subject proximally through the pump-outlet tube. A delivery tube extends from outside the subject's body to the distal portion of the pump-outlet tube. An expandable element surrounds the delivery tube proximally to the blood-outlet openings, a length of the delivery tube between the expandable element and the blood-outlet openings being less than 30 mm. Other applications are also described.

FIELD OF EMBODIMENTS OF THE INVENTION

Some applications of the present invention generally relate to medical apparatus. Specifically, some applications of the present invention relate to a ventricular assist device and methods of use thereof.

BACKGROUND

Ventricular assist devices are mechanical circulatory support devices designed to assist and unload cardiac chambers in order to maintain or augment cardiac output. They are used in patients suffering from a failing heart and in patients at risk for deterioration of cardiac function during percutaneous coronary interventions. Most commonly, a left-ventricular assist device is applied to a defective heart in order to assist left-ventricular functioning. In some cases, a right-ventricular assist device is used in order to assist right-ventricular functioning. Such ventricular assist devices are either designed to be permanently implanted or mounted on a catheter for temporary placement.

SUMMARY OF EMBODIMENTS

In accordance with some applications of the present invention, an expandable element surrounds a delivery tube of a left-ventricular assist device. For example, the expandable element may include an expandable stent, an expandable braided element, or an inflatable element (e.g., a balloon). Typically, the left-ventricular assist device includes a pump-outlet tube shaped to define one or more blood-outlet openings and configured for insertion, through a subject's aorta, into the left ventricle of the subject's heart such that the blood-outlet openings are disposed within the aorta and a distal portion of the pump-outlet tube is disposed within the left ventricle. An impeller is disposed within the distal portion of the pump-outlet tube and is configured to pump blood of the subject proximally through the pump-outlet tube. The delivery tube extends from outside the subject's body, through the pump-outlet tube to the distal portion of the pump-outlet tube.

Typically, the expandable element surrounds the delivery tube proximally to the blood-outlet openings, with the length of the delivery tube between the expandable element and the blood-outlet openings being less than 30 mm. For example, the expandable element may be disposed in the vicinity of the interface between the delivery tube and the region at which the proximal end of the pump-outlet tube is coupled to the delivery tube. In some embodiments, the expandable element is entirely proximal to the pump-outlet tube. In other embodiments, the expandable element is an inflatable element (e.g., a balloon), which when inflated, is disposed at least partly within, e.g., entirely within, the pump-outlet tube.

Typically, the expandable element is configured to center a portion of the ventricular assist device (e.g., a portion of the delivery tube and, in particular, the portion of the delivery tube near the pump-outlet tube) within the aorta, by contacting the aorta wall or, if the expandable element is within the pump-outlet tube, by pushing the wall of the pump-outlet tube against the aorta wall. In some embodiments, the expandable element is an inflatable element that is shaped to direct the blood through the blood-outlet openings. For example, the distal end of the inflatable element may have a width that decreases moving distally, e.g., the distal end may be frustoconical, such that the blood is directed by the distal end of the inflatable element, at an angle, through the blood-outlet openings.

For some applications, the inflatable element (e.g., balloon) is inflated using a fluid, which is pumped through the ventricular-assist device. This fluid may include purging fluid, which, distally to openings, purges an interface between an axial shaft and radial bearings of the device.

For some applications, the pump-outlet tube defines a blood-flow chamber at its proximal end. For some applications, the blood-flow chamber is defined by an internal membrane that is disposed within the proximal end of the pump-outlet tube and that defines holes therethrough. Blood flows into the blood-flow chamber via the holes. Subsequently, the blood flows out of the blood-flow-chamber and into the subject's aorta via blood-outlet openings. Typically, by virtue of the blood flowing through the blood-flow chamber, the blood-flow chamber inflates such as to center a portion of the left-ventricular assist device (e.g., the delivery tube, and in particular, the portion of the delivery tube near the pump-outlet tube) within the aorta, by contacting the aorta wall. Typically, the internal membrane is shaped so as to direct the blood flow out of the blood-outlet openings.

For some applications, the internal membrane is a continuation of the pump-outlet tube, and the internal membrane is covered with an external membrane, which defines the blood outlet openings, and which forms the external surface of blood-flow chamber. For such applications, the proximal end of the blood-outlet tube is shaped so as to direct the blood flow out of the blood-outlet openings.

Typically, the combination of the proximal end of the blood-outlet tube and an additional membrane (whether an internal membrane or an external membrane) is configured to define the blood-flow chamber, which typically functions as described above. In general, the scope of the present disclosure includes any structure that provides a blood-flow chamber disposed at a proximal end of the pump-outlet tube, the blood-flow chamber defining (a) holes via which blood is pumped into the blood-flow chamber and (b) blood-outlet openings configured to be disposed with within the aorta via which the blood flows out of the blood-flow chamber and into the aorta.

For some applications, the proximal end of the pump-outlet tube is folded inwardly so as to define one or more surfaces configured to direct the blood through blood-outlet openings by virtue of being oblique with respect to the longitudinal axis of the pump-outlet tube. In some embodiments, the surfaces define a projection, such as a frustoconical projection, having a width that decreases moving distally. Blood is directed by the projection, at an angle, through the blood-outlet openings. In such embodiments, typically, the projection is distally coupled to the delivery tube.

In general, in the specification and in the claims of the present application, the term “proximal” and related terms, when used with reference to a device or a portion thereof, should be interpreted to mean an end of the device or the portion thereof that, when inserted into a subject's body, is typically closer to a location through which the device is inserted into the subject's body. The term “distal” and related terms, when used with reference to a device or a portion thereof, should be interpreted to mean an end of the device or the portion thereof that, when inserted into a subject's body, is typically further from the location through which the device is inserted into the subject's body.

There is therefore provided, in accordance with some embodiments of the present invention, an apparatus, including:a left-ventricular assist device, including:a pump-outlet tube shaped to define one or more blood-outlet openings and configured for insertion, through an aorta of a subject, into a left ventricle of a heart of the subject such that the blood-outlet openings are disposed within the aorta and a distal portion of the pump-outlet tube is disposed within the left ventricle;an impeller disposed within the distal portion of the pump-outlet tube and configured to pump blood of the subject proximally through the pump-outlet tube;a delivery tube configured to extend, from outside a body of a subject, through the pump-outlet tube to the distal portion of the pump-outlet tube;a drive cable passing through the delivery tube and configured to rotate the impeller; andan expandable element surrounding the delivery tube at least partially proximally to the blood-outlet openings,a length of the delivery tube between the expandable element and the blood-outlet openings being less than 30 mm.

In some embodiments, the expandable element is configured to center the delivery tube within the aorta.

In some embodiments, the pump-outlet tube includes a cylindrical portion, and the cylindrical portion of the pump-outlet tube is shaped to define the blood-outlet openings, such that the blood-outlet openings are laterally facing.

In some embodiments, the pump-outlet tube includes:a narrower section, which includes the distal portion of the pump-outlet tube; anda wider section, which is proximal to and wider than the narrower section, and which is shaped to define at least a portion of each of the blood-outlet openings such that a normal vector to the portion has a distally-facing component.

In some embodiments, the expandable element is entirely proximal to the pump-outlet tube.

In some embodiments, the expandable element includes an inflatable element.

In some embodiments, the inflatable element, when inflated, is disposed at least partly within the pump-outlet tube.

In some embodiments, the inflatable element is disposed entirely within the pump-outlet tube.

In some embodiments, the inflatable element is shaped to direct the blood through the blood-outlet openings.

In some embodiments, a distal end of the inflatable element has a width that decreases moving distally.

In some embodiments, the distal end of the inflatable element is frustoconical.

In some embodiments, a wall of the delivery tube is shaped to define one or more openings, and the inflatable element surrounds the openings such that a fluid flowing, via the openings, from the delivery tube into the inflatable element inflates the inflatable element.

In some embodiments,the left-ventricular assist device further includes:an axial shaft coupled to the impeller and configured to rotate such that the impeller pumps the blood; andat least one bearing configured not to rotate with the axial shaft, anddistally to the openings, the fluid purges an interface between the axial shaft and the bearing.

In some embodiments, the pump-outlet tube is configured to curve proximally to the impeller.

In some embodiments, the pump-outlet tube is configured to curve by virtue of being pre-shaped.

In some embodiments, the blood-outlet openings are arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of the blood flowing through the blood-outlet openings.

In some embodiments, the pump-outlet tube is further shaped to define one or more blood-inlet openings arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of the blood flowing through the blood-inlet openings.

In some embodiments, the left-ventricular assist device further includes one or more bands bonded to an outer wall of the pump-outlet tube, and the pump-outlet tube is configured to curve by virtue of the bands being bonded to the outer wall.

There is further provided, in accordance with some embodiments of the present invention, an apparatus, including:a left-ventricular assist device, including:a pump-outlet tube configured for insertion, through an aorta of a subject, into a left ventricle of a heart of the subject such that the pump-outlet tube traverses an aortic valve of the subject;an impeller disposed at least partly within a distal portion of the pump-outlet tube and configured to pump blood of the subject proximally through the pump-outlet tube; anda blood-flow chamber disposed at a proximal end of the pump-outlet tube, the blood-flow chamber defining (a) holes, via which blood is pumped into the blood-flow chamber and (b) blood-outlet openings, configured to be disposed with within the aorta, via which the blood flows out of the blood-flow chamber and into the aorta.

In some embodiments, the left-ventricular assist device includes an external membrane that defines the blood-outlet openings and forms an external surface of the blood-flow chamber, and the proximal end of the pump-outlet tube is disposed within the external membrane and defines the holes.

In some embodiments, the left-ventricular assist device includes an internal membrane that defines the holes and forms an internal surface of the blood-flow chamber, and the proximal end of the pump-outlet tube is disposed outside the external membrane and defines the blood-outlet openings.

In some embodiments, by virtue of the blood flowing through the blood-flow chamber, the blood-flow chamber is configured to inflate such as to center a portion of the left-ventricular assist device within the aorta.

In some embodiments, the pump-outlet tube includes a cylindrical portion, and the cylindrical portion of the pump-outlet tube is shaped to define the blood-outlet openings, such that the blood-outlet openings are laterally facing.

In some embodiments, the pump-outlet tube is configured to curve proximally to the impeller.

In some embodiments, the pump-outlet tube is configured to curve by virtue of being pre-shaped.

In some embodiments, the blood-outlet openings are arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of the blood flowing through the blood-outlet openings.

In some embodiments, the pump-outlet tube is further shaped to define one or more blood-inlet openings arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of the blood flowing through the blood-inlet openings.

In some embodiments, the left-ventricular assist device further includes one or more bands bonded to an outer wall of the pump-outlet tube, and the pump-outlet tube is configured to curve by virtue of the bands being bonded to the outer wall.

There is further provided, in accordance with some embodiments of the present invention, an apparatus, including:a left-ventricular assist device, including:a pump-outlet tube shaped to define one or more blood-outlet openings and configured for insertion, through an aorta of a subject, into a left ventricle of a heart of the subject such that the pump-outlet tube traverses an aortic valve of the subject with the blood-outlet openings being disposed within the aorta; anda blood pump disposed at least partly within a distal portion of the pump-outlet tube and configured to pump blood of the subject proximally through the pump-outlet tube,a proximal end of the pump-outlet tube being folded inwardly so as to define one or more surfaces configured to direct the blood through the blood-outlet openings by virtue of being oblique with respect to a longitudinal axis of the pump-outlet tube.

In some embodiments, the pump-outlet tube includes a cylindrical portion, and the cylindrical portion of the pump-outlet tube is shaped to define the blood-outlet openings, such that the blood-outlet openings are laterally facing.

In some embodiments, the pump-outlet tube includes:a narrower section, which includes the distal portion of the pump-outlet tube; anda wider section, which is proximal to and wider than the narrower section, and which is shaped to define at least a portion of each of the blood-outlet openings such that a normal vector to the portion has a distally-facing component.

In some embodiments, the surfaces define a projection having a width that decreases moving distally.

In some embodiments, the projection is frustoconical.

In some embodiments,the blood pump includes an impeller,the left-ventricular assist device further includes:a delivery tube passing through the pump-outlet tube to the blood pump; anda drive cable passing through the delivery tube and configured to rotate the impeller, andthe projection is distally coupled to the delivery tube.

In some embodiments, the pump-outlet tube is configured to curve proximally to the blood pump.

In some embodiments, the pump-outlet tube is configured to curve by virtue of being pre-shaped.

In some embodiments, the blood-outlet openings are arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of the blood flowing through the blood-outlet openings.

In some embodiments, the pump-outlet tube is further shaped to define one or more blood-inlet openings arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of the blood flowing through the blood-inlet openings.

In some embodiments, the left-ventricular assist device further includes one or more bands bonded to an outer wall of the pump-outlet tube, and the pump-outlet tube is configured to curve by virtue of the bands being bonded to the outer wall.

There is further provided, in accordance with some embodiments of the present invention, an apparatus, including:a left-ventricular assist device, including:a pump-outlet tube shaped to define one or more laterally-facing blood-outlet openings and including a first section and a second section that overlap one another between the laterally-facing blood-outlet openings,the pump-outlet tube being configured for insertion, through an aorta of a subject, into a left ventricle of a heart of the subject such that the pump-outlet tube traverses an aortic valve of the subject with the lateral blood-outlet openings being disposed within the aorta; andan impeller disposed within the first section of the pump-outlet tube and configured to pump blood of the subject, through the laterally-facing blood-outlet openings, from the left ventricle into the aorta.

In some embodiments, the first section and second section overlap one another by 0.25-1 mm.

In some embodiments, the blood-outlet openings occupy 20-80% of a circumference of the pump-outlet tube.

In some embodiments, the pump-outlet tube is configured to curve proximally to the impeller.

In some embodiments, the pump-outlet tube is configured to curve by virtue of being pre-shaped.

In some embodiments, the blood-outlet openings are arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of the blood flowing through the blood-outlet openings.

In some embodiments, the pump-outlet tube is further shaped to define one or more blood-inlet openings arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of the blood flowing through the blood-inlet openings.

In some embodiments, the left-ventricular assist device further includes one or more bands bonded to an outer wall of the pump-outlet tube, and the pump-outlet tube is configured to curve by virtue of the bands being bonded to the outer wall.

In some embodiments, the left-ventricular assist device further includes:a delivery tube configured to extend, from outside a body of a subject, through the pump-outlet tube to the first section;a drive cable passing through the delivery tube and configured to rotate the impeller; andan expandable element surrounding the delivery tube proximally to the blood-outlet openings,a length of the delivery tube between the expandable element and the blood-outlet openings being less than 30 mm.

In some embodiments, the expandable element is configured to center the delivery tube within the aorta.

In some embodiments, the expandable element includes an inflatable element.

In some embodiments, the inflatable element, when inflated, is disposed at least partly within the pump-outlet tube.

In some embodiments, the inflatable element is disposed entirely within the pump-outlet tube.

In some embodiments, the inflatable element is shaped to direct the blood through the blood-outlet openings.

In some embodiments, the expandable element is entirely proximal to the pump-outlet tube.

There is further provided, in accordance with some embodiments of the present invention, a method for assembling a left-ventricular assist device, the method including:passing an impeller, which is configured to pump blood of a subject, through a proximal end of a first section of a pump-outlet tube, such that the impeller is disposed within the first section,the proximal end of the first section including multiple first-section tabs and being shaped to define respective first-section gaps between successive ones of the first-section tabs; andbonding a second section of the pump-outlet tube, which includes multiple second-section tabs and is shaped to define respective second-section gaps between successive ones of the second-section tabs, to the first section such that the second-section tabs overlap the first-section tabs and the second-section gaps are continuous with the first-section gaps so as to define one or more laterally-facing blood-outlet openings.

In some embodiments, bonding the second section to the first section includes bonding the second section to the first section such that the first section and second section overlap one another by 0.25-1 mm.

There is further provided, in accordance with some embodiments of the present invention, an apparatus, including:a left-ventricular assist device, including:a pump-outlet tube shaped to define one or more blood-outlet openings and including:a narrower section; anda wider section, which is proximal to and wider than the narrower section, and which is shaped to define at least a portion of each of the blood-outlet openings such that a normal vector to the portion has a distally-facing component,the pump-outlet tube being configured for insertion, through an aorta of a subject, into a left ventricle of a heart of the subject such that the pump-outlet tube traverses an aortic valve of the subject with the blood-outlet openings being disposed within the aorta; andan impeller disposed within the narrower section of the pump-outlet tube and configured to pump blood of the subject, through the blood-outlet openings, from the left ventricle into the aorta.

In some embodiments,the pumping of the blood produces a distal thrust on the pump-outlet tube, andby virtue of the distally-facing component, a flow of the blood through the blood-outlet openings produces a proximal thrust on the pump-outlet tube that at least partially cancels the distal thrust.

In some embodiments, an angle between the normal vector and a longitudinal axis of the pump-outlet tube at the wider section is between 20 and 80 degrees.

In some embodiments, the normal vector is parallel to a longitudinal axis of the pump-outlet tube at the wider section.

In some embodiments, each of the blood-outlet openings spans an interface between the narrower section and the wider section.

In some embodiments, the pump-outlet tube is configured to curve proximally to the impeller.

In some embodiments, the pump-outlet tube is configured to curve by virtue of being pre-shaped.

In some embodiments, the blood-outlet openings are arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of a flow of the blood through the blood-outlet openings.

In some embodiments, the pump-outlet tube is further shaped to define one or more blood-inlet openings arranged in a non-axisymmetric arrangement, and the pump-outlet tube is configured to curve by virtue of the blood flowing through the blood-inlet openings.

In some embodiments, the left-ventricular assist device further includes one or more bands bonded to an outer wall of the pump-outlet tube, and the pump-outlet tube is configured to curve by virtue of the bands being bonded to the outer wall.

In some embodiments, the left-ventricular assist device further includes:a delivery tube configured to extend, from outside a body of a subject, through the pump-outlet tube to the narrower section;a drive cable passing through the delivery tube and configured to rotate the impeller; andan expandable element surrounding the delivery tube proximally to the blood-outlet openings,a length of the delivery tube between the expandable element and the blood-outlet openings being less than 30 mm.

In some embodiments, the expandable element is configured to center the delivery tube within the aorta.

In some embodiments, the expandable element is entirely proximal to the pump-outlet tube.

In some embodiments, the expandable element includes an inflatable element.

In some embodiments, the inflatable element, when inflated, is disposed at least partly within the pump-outlet tube.

In some embodiments, the inflatable element is disposed entirely within the pump-outlet tube.

In some embodiments, the inflatable element is shaped to direct the blood through the blood-outlet openings.

There is further provided, in accordance with some embodiments of the present invention, an apparatus, including:a left-ventricular assist device, including:a pump-outlet tube configured for insertion, through an aorta of a subject, into a left ventricle of a heart of the subject such that a distal portion of the pump-outlet tube is disposed within the left ventricle and a proximal portion of the pump-outlet tube is disposed within the aorta;an impeller disposed within the distal portion of the pump-outlet tube and configured to pump blood of the subject proximally through the pump-outlet tube;a delivery tube configured to extend, from outside a body of a subject, through the pump-outlet tube to the distal portion;a drive cable passing through the delivery tube and configured to rotate the impeller; anda porous expandable element surrounding the delivery tube within, or at a proximal end of, the pump-outlet tube, such that the blood is pumped through the porous expandable element.

In some embodiments, the porous expandable element is configured to center the delivery tube within the aorta.

In some embodiments, the porous expandable element includes a structure made of a shape-memory alloy.

DETAILED DESCRIPTION OF EMBODIMENTS

Reference is now made toFIGS.1A,1B, and1C, which are schematic illustrations of a ventricular assist device20, a distal end of which is configured to be disposed in a subject's left ventricle22, in accordance with some applications of the present invention.FIG.1Ashows an overview of the ventricular assist device system including a control console21and a motor unit23.FIG.1Bshows the ventricular assist device being inserted into the subject's left ventricle, andFIG.1Cshows a pump-head portion27of the ventricular assist device in greater detail.

As shown inFIG.1B, the ventricular assist device includes a pump-outlet tube24, which traverses an aortic valve26of the subject, such that a proximal end28of the pump-outlet tube is disposed in an aorta30of the subject and a distal end32of the pump-outlet tube is disposed within left ventricle22. Pump-outlet tube24(which may also be referred to as a “blood-pump tube”) is typically an elongate tube, an axial length of the pump-outlet tube typically being substantially larger than its diameter. The scope of the present invention includes using the apparatus and methods described herein in anatomical locations other than the left ventricle and the aorta. Therefore, the ventricular assist device and/or portions thereof are sometimes referred to herein (in the specification and the claims) as a blood pump.

For some applications, the ventricular assist device is used to assist the functioning of a subject's left ventricle during a percutaneous coronary intervention. In such cases, the ventricular assist device is typically used for a period of up to six hours (e.g., up to ten hours), during a period in which there is risk of developing hemodynamic instability (e.g., during or immediately following the percutaneous coronary intervention). Alternatively or additionally, the ventricular assist device is used to assist the functioning of a subject's left ventricle for a longer period (e.g., 2-20 days, e.g., 4-14 days) upon a patient suffering from cardiogenic shock, which may include any low-cardiac-output state (e.g., acute myocardial infarction, myocarditis, cardiomyopathy, post-partum, etc.). For some applications, the ventricular assist device is used to assist the functioning of a subject's left ventricle for yet a longer period (e.g., several weeks or months), e.g., in a “bridge to recovery” treatment. For some such applications, the ventricular assist device is permanently or semi-permanently implanted, and the impeller of the ventricular assist device is powered transcutaneously, e.g., using an external antenna that is magnetically coupled to the impeller.

As shown inFIG.1B, which shows steps in the deployment of the ventricular assist device in the left ventricle, typically the distal end of the ventricular assist device is guided to the left ventricle over a guidewire10. During the insertion of the distal end of the device into the left ventricle, a delivery catheter143is disposed over the distal end of the device. Once the distal end of the device is disposed in the left ventricle, the delivery catheter is typically retracted to the aorta, and the guidewire is withdrawn from the subject's body. The retraction of the delivery catheter typically causes self-expandable components of the distal end of the device to assume non-radially-constrained configurations, as described in further detail hereinbelow. Typically, the ventricular assist device is inserted into the subject's body in order to provide an acute treatment to the subject. For some applications, in order to withdraw the left ventricular device from the subject's body at the end of the treatment, the delivery catheter is advanced over the distal end of the device, which causes the self-expandable components of the distal end of the device to assume radially-constrained configurations. Alternatively or additionally, the distal end of the device is retracted into the delivery catheter which causes the self-expandable components of the distal end of the device to assume radially-constrained configurations.

For some applications (not shown), the ventricular assist device and/or delivery catheter143includes an ultrasound transducer at its distal end and the ventricular assist device is advanced toward the subject's ventricle under ultrasound guidance.

Reference is made toFIG.1C, which shows pump-head portion27of ventricular assist device20, in accordance with some applications of the present invention, in greater detail. Typically, an impeller50is disposed within a distal portion102of pump-outlet tube24and is configured to pump blood from the left ventricle into the aorta by rotating. The pump-outlet tube typically defines one or more blood-inlet openings108at the distal end of the pump-outlet tube, via which blood flows into the pump-outlet tube from the left ventricle, during operation of the impeller. As shown inFIG.1C, for some applications, the pump-outlet tube defines a single axially-facing blood-inlet opening. Alternatively, the pump-outlet tube defines a plurality of lateral blood-inlet openings (e.g., as shown inFIG.1B), as described in further detail hereinbelow. For some applications, proximal portion106of the pump-outlet tube defines one or more blood-outlet openings109, via which blood flows from the pump-outlet tube into the ascending aorta, during operation of the impeller.

For some applications, control console21(shown inFIG.1A), which typically includes a computer processor25, drives the impeller to rotate. For example, the computer processor may control a motor74(shown inFIG.7B), which is disposed within motor unit23(shown inFIG.1A) and which drives the impeller to rotate via a drive cable130(shown inFIG.12A). For some applications, the computer processor is configured to detect a physiological parameter of the subject (such as left-ventricular pressure, cardiac afterload, rate of change of left-ventricular pressure, etc.) and to control rotation of the impeller in response thereto, as described in further detail hereinbelow. Typically, the operations described herein that are performed by the computer processor, transform the physical state of a memory, which is a real physical article that is in communication with the computer processor, to have a different magnetic polarity, electrical charge, or the like, depending on the technology of the memory that is used. Computer processor25is typically a hardware device programmed with computer program instructions to produce a special-purpose computer. For example, when programmed to perform the techniques described herein, computer processor25typically acts as a special-purpose, ventricular-assist computer processor and/or a special-purpose, blood-pump computer processor.

For some applications, a purging system29(shown inFIG.1A) drives a fluid (e.g., a glucose solution) to pass through portions of ventricular assist device20, for example, in order to cool portions of the device, to purge and/or lubricate interfaces between rotating parts and stationary bearings, and/or in order to wash debris from portions of the device.

Typically, along distal portion102of pump-outlet tube24, a frame34is disposed within the pump-outlet tube around impeller50. The frame is typically made of a shape-memory alloy, such as nitinol. For some applications, the shape-memory alloy of the frame is shape set such that at least a portion of the frame (and thereby distal portion102of tube24) assumes a generally circular, elliptical, or polygonal cross-sectional shape in the absence of any forces being applied to distal portion102of tube24. By assuming its generally circular, elliptical, or polygonal cross-sectional shape, the frame is configured to hold the distal portion of the pump-outlet tube in an open state. Typically, during operation of the ventricular assist device, the distal portion of the pump-outlet tube is configured to be placed within the subject's body, such that the distal portion of the pump-outlet tube is disposed at least partially within the left ventricle.

For some applications, along proximal portion106of pump-outlet tube24, the frame is not disposed within the pump-outlet tube, and the pump-outlet tube is therefore not supported in an open state by frame34. Pump-outlet tube24is typically made of a blood-impermeable collapsible material, such that the pump-outlet tube is collapsible. For example, pump-outlet tube24may include polyurethane, polyester, and/or silicone. Alternatively or additionally, the pump-outlet tube is made of polyethylene terephthalate (PET) and/or polyether block amide (e.g., PEBAX®). For some applications (not shown), the pump-outlet tube is reinforced with a reinforcement structure, e.g., a braided reinforcement structure, such as a braided nitinol tube. Typically, the proximal portion of the pump-outlet tube is configured to be placed such that it is at least partially disposed within the subject's ascending aorta. For some applications, the proximal portion of the pump-outlet tube traverses the subject's aortic valve, passing from the subject's left ventricle into the subject's ascending aorta, as shown inFIG.1B.

As described hereinabove, the pump-outlet tube typically defines one or more blood-inlet openings108at the distal end of the pump-outlet tube, via which blood flows into the pump-outlet tube from the left ventricle, during operation of the impeller. For some applications, the proximal portion of the pump-outlet tube defines one or more blood-outlet openings109, via which blood flows from the pump-outlet tube into the ascending aorta, during operation of the impeller. Typically, the pump-outlet tube defines a plurality of blood-outlet openings109, for example, between two and eight blood-outlet openings (e.g., between two and four blood-outlet openings). During operation of the impeller, the pressure of the blood flow through the pump-outlet tube typically maintains the proximal portion of the tube in an open state. For some applications, in the event that, for example, the impeller malfunctions, the proximal portion of the pump-outlet tube is configured to collapse inwardly, in response to pressure outside of the proximal portion of the pump-outlet tube exceeding pressure inside the proximal portion of the pump-outlet tube. In this manner, the proximal portion of the pump-outlet tube acts as a safety valve, preventing retrograde blood flow into the left ventricle from the aorta.

Referring again toFIG.1C, for some applications, frame34is shaped such that the frame defines a proximal conical portion36, a central cylindrical portion38, and a distal conical portion40. Typically, the proximal conical portion is proximally-facing, i.e., facing such that the narrow end of the cone is proximal with respect to the wide end of the cone. Further typically, the distal conical portion is distally-facing, i.e., facing such that the narrow end of the cone is distal with respect to the wide end of the cone.

For some applications, within at least a portion of frame34(e.g., along all of, or a portion of, the central cylindrical portion of the frame), an inner lining39, shown inFIG.4for example, lines the frame. In accordance with respective applications, the inner lining partially overlaps or fully overlaps with pump-outlet tube24over the portion of the frame that the inner lining lines, as described in further detail hereinbelow with reference toFIGS.10A-B. For other applications, as shown inFIG.1C, the pump-head portion does not comprise inner lining39.

Typically, pump-outlet tube24includes a conical proximal portion42and a cylindrical central portion44. The proximal conical portion is typically proximally-facing, i.e., facing such that the narrow end of the cone is proximal with respect to the wide end of the cone. Typically, blood-outlet openings109are defined by pump-outlet tube24, such that the openings extend at least partially along the proximal conical portion of tube24. For some such applications, the blood-outlet openings are teardrop-shaped, as shown inFIG.1C. Typically, the teardrop-shaped nature of the blood-outlet openings in combination with the openings extending at least partially along the proximal conical portion of tube24causes blood to flow out of the blood-outlet openings along flow lines that are substantially parallel with the longitudinal axis of tube24at the location of the blood-outlet openings.

For some applications (not shown), the diameter of pump-outlet tube24changes along the length of the central portion of the pump-outlet tube, such that the central portion of the pump-outlet tube has a frustoconical shape. For example, the central portion of the pump-outlet tube may widen from its proximal end to its distal end, or may narrow from its proximal end to its distal end. For some applications, at its proximal end, the central portion of the pump-outlet tube has a diameter of between 5 and 7 mm, and at its distal end, the central portion of the pump-outlet tube has a diameter of between 8 and 12 mm.

Again referring toFIG.1C, the ventricular assist device typically includes a distal-tip element107that is disposed distally with respect to frame34and that includes an axial-shaft-receiving tube126and a distal-tip portion120. Typically, the axial-shaft-receiving tube is configured to receive a distal portion of an axial shaft92(FIG.4) of the pump-head portion during axial back-and-forth motion of the axial shaft (as described in further detail hereinbelow), and/or during delivery of the ventricular assist device. (Typically, during delivery of the ventricular assist device, the frame is maintained in a radially-constrained configuration, which typically causes the axial shaft to be disposed in a different position with respect to the frame relative to its disposition with respect to the frame during operation of the ventricular assist device). Typically, distal-tip portion120is configured to assume a curved shape upon being deployed within the subject's left ventricle, e.g., as shown inFIG.1C. For some applications, the curvature of the distal-tip portion is configured to provide an atraumatic tip to ventricular assist device20. Alternatively or additionally, the distal-tip portion is configured to space blood-inlet openings108of the ventricular assist device from walls of the left ventricle.

As shown in the enlarged portion ofFIG.1B, for some applications, pump-outlet tube24extends to the end of distal conical portion40of the frame, and the pump-outlet tube defines a plurality of lateral blood-inlet openings108, as described in further detail hereinbelow. For such applications, the pump-outlet tube typically defines a distal conical portion that is distally facing, i.e., such that the narrow end of the cone is distal with respect to the wide end of the conc. For some such applications (not shown), the pump-outlet tube defines two to four lateral blood-inlet openings (e.g., four lateral blood-inlet openings). Typically, for such applications, each of the blood-inlet openings defines an area of more than 20 square mm (e.g., more than 30 square mm), and/or less than 60 square mm (e.g., less than 50 square mm), e.g., 20-60 square mm, or 30-50 square mm. Alternatively or additionally, the outlet tube defines a greater number of smaller lateral blood-inlet openings, e.g., more than 10 blood-inlet openings, more than 50 blood-inlet openings, more than 200 blood-inlet openings, or more than 400 blood-inlet openings, e.g., 50-100 blood-inlet openings, 100-400 blood-inlet openings, or 400-600 blood-inlet openings. For some such applications, each of the blood-inlet openings defines an area of more than 0.05 square mm (e.g., more than 0.1 square mm), and/or less than 3 square mm (e.g., less than 1 square mm), e.g., 0.05-3 square mm, or 0.1-1 square mm. Alternatively, each of the blood-inlet openings defines an area of more than 0.1 square mm (e.g., more than 0.3 square mm), and/or less than 5 square mm (e.g., less than 1 square mm), e.g., 0.1-5 square mm, or 0.3-1 square mm. Such applications are described in further detail hereinbelow, for example, with reference toFIGS.11A-E.

It is noted that the lateral blood-inlet openings are typically defined by the distal conical portion of the pump-outlet tube. As such, even the blood-inlet openings that are described as “lateral blood-inlet openings” are typically not oriented entirely laterally with respect to the longitudinal axis of the pump-outlet tube. Rather, they are obliquely disposed with respect to the longitudinal axis of the pump-outlet tube. By contrast, in some embodiments, the blood-outlet openings are described as “laterally-facing blood-outlet openings” because in such embodiments the blood-outlet openings are disposed laterally with respect to the longitudinal axis of the pump-outlet tube, by virtue of being defined by the central cylindrical portion of the pump-outlet tube. It is noted that in other embodiments, the blood-outlet openings are disposed obliquely with respect to the longitudinal axis of the pump-outlet tube, by virtue of being defined at least partially by the proximal conical portion of the pump-outlet tube.

In general, the scope of the present disclosure includes combining a pump-outlet-tube that defines a single axially-facing blood-inlet opening108as shown inFIG.1C, or a pump-outlet-tube that defines a plurality of lateral blood-inlet openings108as shown inFIG.1Bin combination with other features of the ventricular assist device that are described herein, mutatis mutandis.

Reference is now made toFIG.2, which is schematic illustration of frame34that houses an impeller of ventricular assist device20, in accordance with some applications of the present invention. Frame34is typically made of a shape-memory alloy, such as nitinol, and the shape-memory alloy of the frame is shape set such that the central portion38of the frame (and thereby tube24) assumes a generally circular, elliptical, or polygonal cross-sectional shape in the absence of any forces being applied to pump-outlet tube24. By assuming its generally circular, elliptical, or polygonal cross-sectional shape, the frame is configured to hold the distal portion of the tube in an open state. (Given that, typically, central portion38of the frame has a circular cross-section, the central portion of the frame is also referred to herein as the “cylindrical portion” of the frame.)

Typically, the frame is a stent-like frame, in that it comprises struts that, in turn, define cells. Further typically, the frame is covered with pump-outlet tube24, and/or covered with an inner lining39, described hereinbelow with reference toFIGS.10A-B. As described hereinbelow, for some applications, impeller50undergoes axial back-and-forth motion with respect to frame34. Typically, over the course of the motion of the impeller with respect to the frame, the location of the portion of the impeller that defines the maximum span of the impeller is disposed within central cylindrical portion38of frame34. In some cases, if the cells of the central cylindrical portion38of frame34are too large, then pump-outlet tube24, and/or inner lining39, gets stretched between edges of the cells, such that the pump-outlet tube24, and/or inner lining39, does not define a circular cross-section. For some applications, if this occurs in the region in which the portion of the impeller that defines the maximum span of the impeller is disposed, this results in a substantially non-constant gap between the edges of the impeller blades and tube24(and/or inner lining) at that location, over the course of a rotation cycle of the impeller. For some applications, this may lead to increased hemolysis relative to if there were a substantially constant gap between the edges of the impeller blades and tube24(and/or inner lining) at that location, over the course of the rotation cycle of the impeller.

Referring toFIG.2, at least partially in view of the issues described in the above paragraph, within central cylindrical portion38of frame34, the frame defines a large number of relatively small cells. Typically, when the frame is disposed in its non-radially-constrained configuration, the maximum cell width CW of the each of the cells (i.e., the distance from the inner edge of the strut at the central junction on one side of the cell to the inner edge of the strut at the central junction on the other side of the cell, as measured around the circumference of cylindrical portion38) within the cylindrical portion of the frame is less than 2 mm, e.g., between 1.4 mm and 1.6 mm, or between 1.6 and 1.8 mm. Since the cells are relatively small, inner lining39defines a substantially circular cross-section within the cylindrical portion of the frame.

Still referring toFIG.2, and starting from the distal end of the frame (which is to the right of the figure), typically the frame defines the following portions: (a) coupling portion31via which the frame is coupled to a distal bearing housing118H (shown inFIG.5A) of the ventricular assist device, (b) distal conical portion40, (c) central cylindrical portion38, (d) proximal conical portion36, and (c) proximal strut junctions33. As illustrated, as the frame transitions from a proximal end of the frame toward the center of the frame (e.g., as the frame transitions from proximal strut junctions33, through proximal conical portion36, and to central cylindrical portion38), struts37of the frame pass through junctions35, at which the two struts branch from a single strut, in a Y-shape. As described in further detail hereinbelow, typically frame34is placed in a radially-constrained (i.e., crimped) configuration within delivery catheter143by the frame being axially elongated. Moreover, the frame typically transmits its radial narrowing to the impeller, and the impeller becomes radially constrained by becoming axially elongated within the frame. For some applications, the struts of the frame being configured in the manner described above facilitates transmission of axial elongation from the delivery catheter (or another device that is configured to crimp the frame) to the frame, which in turn facilitates transmission of axial elongation to the impeller. This is because the pairs of struts that branch from each of junctions35are configured to pivot about the junction and move closer to each other such as to close the space between them.

Still referring toFIG.2, during the assembly of the ventricular assist device, initially, distal coupling portion31is coupled to a distal bearing housing118H (shown inFIG.5A), e.g., via a snap-fit mechanism. For some applications, proximal strut junctions33are still maintained in open states at this stage, in order for the impeller to be placed within the frame via the proximal end of the frame. Typically, the structure of frame34shown inFIG.2is used in applications in which pump-outlet tube extends to the distal end of frame34(e.g., as shown inFIG.1B). In such cases, the impeller cannot be inserted via the distal end of the frame, since the distal end of the frame is covered by pump-outlet tube24. During the assembly of the ventricular assist device, subsequently to the impeller being inserted via the proximal end of the frame, the proximal strut junctions are closed. For some applications, the proximal strut junctions are closed around the outside of a proximal bearing housing116H (shown inFIG.5A), as described in further detail hereinbelow with reference toFIGS.5A-B. Typically, a securing element117(e.g., a ring shown inFIG.5A) holds the strut junctions in their closed configurations around the outside of proximal bearing housing116H.

Typically, when disposed in its non-radially constrained configuration, frame34has a total length of more than 25 mm (e.g., more than 30 mm), and/or less than 50 mm (e.g., less than 45 mm), e.g., 25-50 mm, or 30-45 mm. Typically, when disposed in its radially-constrained configuration (within delivery catheter143), the length of the frame increases by between 2 and 5 mm. Typically, when disposed in its non-radially constrained configuration, the central cylindrical portion of frame34has a length of more than 10 mm (e.g., more than 12 mm), and/or less than 25 mm (e.g., less than 20 mm), e.g., 10-25 mm, or 12-20 mm. For some applications, a ratio of the length of the central cylindrical portion of the frame to the total length of the frame is more than 1:4 and/or less than 1:2, e.g., between 1:4 and 1:2.

Reference is now made toFIGS.3A-E, which are schematic illustrations of impeller50or portions thereof, in accordance with some applications of the present invention. Typically, the impeller includes at least one outer helical elongate element52, which winds around a central axial spring54, such that the helix defined by the helical elongate element is coaxial with the central axial spring. Typically, the impeller includes two or more helical elongate elements (e.g., three helical elongate elements, as shown inFIGS.3A-C). For some applications, the helical elongate elements and the central axial spring are made of a shape-memory material, e.g., a shape-memory alloy, such as nitinol. Typically, each of the helical elongate elements and the central axial spring support a film56of a material (e.g., an elastomer, such as polyurethane, and/or silicone) therebetween. For some applications, the film of material includes pieces of nitinol embedded therein, for example in order to strengthen the film of material. For illustrative purposes, the impeller is shown in the absence of the material inFIG.3A.FIGS.3B and3Cshow respective views of the impeller with the material supported between the helical elongate elements and the spring.FIGS.3D and3Eshow similar respective views of the impeller to those shown inFIGS.3B and3C, but with certain features of the impeller differing from those shown inFIGS.3B and3C, as elaborated upon hereinbelow.

Each of the helical elongate elements, together with the film extending from the helical elongate element to the spring, defines a respective impeller blade, with the helical elongate elements defining the outer edges of the blades, and the axial spring defining the axis of the impeller. Typically, the film of material extends along and coats the spring. For some applications, sutures53(e.g., polyester sutures, shown inFIGS.3A-C) are wound around the helical elongate elements. Typically, the sutures are configured to facilitate bonding between the film of material (which is typically an elastomer, such as polyurethane, or silicone) and the helical elongate element (which is typically a shape-memory alloy, such as nitinol). For some applications, sutures (e.g., polyester sutures, not shown) are wound around spring54. Typically, the sutures are configured to facilitate bonding between the film of material (which is typically an elastomer, such as polyurethane, or silicone) and the spring (which is typically a shape-memory alloy, such as nitinol).

Typically, proximal ends of spring54and helical elongate elements52extend from a proximal bushing (i.e., sleeve bearing)64of the impeller, such that the proximal ends of spring54and helical elongate elements52are disposed at a similar radial distance from the longitudinal axis of the impeller, as each other. Similarly, typically, distal ends of spring54and helical elongate elements52extend from a distal bushing58of the impeller, such that the distal ends of spring54and helical elongate elements52are disposed at a similar radial distance from the longitudinal axis of the impeller, as each other. The helical elongate elements typically rise gradually from the proximal bushing before reaching a maximum span and then falling gradually toward the distal bushing. Typically, the helical elongate elements are symmetrical along their lengths, such that the rising portions of their lengths are symmetrical with respect to the falling portions of their lengths. Typically, the impeller defines a lumen62therethrough (shown inFIG.3C), with the lumen typically extending through, and being defined by, spring54, as well as proximal bushing64and distal bushing58, of the impeller.

Reference is now made toFIG.4, which is a schematic illustration of impeller50disposed inside frame34of ventricular assist device20, in accordance with some applications of the present invention. For some applications, within at least a portion of frame34(e.g., along all of, or a portion of, central cylindrical portion38of the frame), inner lining39lines the frame. In accordance with respective applications, the inner lining partially overlaps or fully overlaps with pump-outlet tube24over the portion of the frame that the inner lining lines, as described in further detail hereinbelow with reference toFIGS.9A-B.

As shown inFIG.4, typically there is a gap G between the outer edge of impeller50and inner lining39, even at a location at which the span of the impeller is at its maximum. For some applications, it is desirable that the gap between the outer edge of the blade of the impeller and inner lining39be relatively small, in order for the impeller to efficiently pump blood from the subject's left ventricle into the subject's aorta. (It is noted that, by virtue of the relatively small gap between the outer edge of impeller50and inner lining39even at a location at which the span of the impeller is at its maximum, as well as the shape of the impeller, the impeller functions as an axial-flow impeller, with the impeller pumping blood in the axial direction from a distal end of pump-outlet tube24to the proximal end of the pump-outlet tube.) It is also desirable that a gap between the outer edge of the blade of the impeller and the inner surface of frame34be maintained throughout the rotation of the impeller within frame34, for example, in order to reduce the risk of hemolysis.

For some applications, when impeller50and frame34are both disposed in non-radially-constrained configurations and prior to operation of the impeller, gap G between the outer edge of the impeller and the inner lining39, at the location at which the span of the impeller is at its maximum, is greater than 0.05 mm (e.g., greater than 0.1 mm), and/or less than 1 mm (e.g., less than 0.4 mm), e.g., 0.05-1 mm, or 0.1-0.4 mm. For some applications, when the impeller is disposed in its non-radially-constrained configurations and prior to operation of the impeller, the outer diameter of the impeller at the location at which the outer diameter of the impeller is at its maximum is more than 7 mm (e.g., more than 8 mm), and/or less than 10 mm (e.g., less than 9 mm), e.g., 7-10 mm, or 8-9 mm. For some applications, when frame34is disposed in its non-radially-constrained configuration, the inner diameter of frame34(as measured from the inside of inner lining39on one side of the frame to the inside of inner lining on the opposite side of the frame) is greater than 7.5 mm (e.g., greater than 8.5 mm), and/or less than 10.5 mm (e.g., less than 9.5 mm), e.g., 7.5-10.5 mm, or 8.5-9.5 mm. For some applications, when the frame is disposed in its non-radially-constrained configuration, the outer diameter of frame34is greater than 8 mm (e.g., greater than 9 mm), and/or less than 13 mm (e.g., less than 12 mm), e.g., 8-13 mm, or 9-12 mm.

Typically, an axial shaft92passes through the axis of impeller50, via lumen62of the impeller. For some applications, the axial shaft is rigid, e.g., a rigid tube. For some applications, the axial shaft is made of a shape-memory material (e.g., a shape memory alloy, such as nitinol). Typically, such materials have some elasticity, such that in the event that the axial shaft becomes bent (e.g., during delivery of the pump head to the left ventricle), the axial shaft still assumes a straight shape, once deployed inside the subject's body.

Proximal bushing64is disposed over axial shaft92, and distal bushing58is disposed over the axial shaft distally from the proximal bushing. For some applications, proximal bushing64of the impeller is coupled to the shaft such that the axial position of the proximal bushing with respect to the shaft is fixed, and distal bushing58of the impeller is slidable with respect to (i.e., is slidable along) the shaft. For example, the proximal bushing may be coupled to a coupling element65disposed on the axial shaft (shown inFIG.4), for example via a snap-fit mechanism. Alternatively, distal bushing58of the impeller is coupled to the shaft such that the axial position of the distal bushing with respect to the shaft is fixed, and proximal bushing64of the impeller is slidable with respect to the shaft.

The axial shaft itself is radially stabilized via a proximal radial bearing116and a distal radial bearing118(FIG.5A). In turn, the axial shaft, by passing through lumen62defined by the impeller, radially stabilizes the impeller with respect to the inner surface of frame34, such that even a relatively small gap between the outer edge of the blade of the impeller and the inner surface of frame34(e.g., a gap that is as described above) is maintained, during rotation of the impeller.

Referring again toFIGS.3A-C, for some applications, the impeller includes a plurality of elongate elements67extending radially from central axial spring54to outer helical elongate elements52. The elongate elements are typically flexible but are substantially non-stretchable along the axis defined by the elongate elements. Further typically, each of the elongate elements is configured not to exert force upon the helical elongate element, unless force is acting upon the impeller that is causing the helical elongate element to move radially outward, such that (in the absence of the elongate element) a separation between the helical elongate element and the central axial spring would be greater than a length of the elongate element. For example, the elongate elements may include strings (such as polyester, and/or another polymer or a natural material that contains fibers) and/or wires (such as nitinol wires, and/or wires made of a different alloy, or a metal).

For some applications, the elongate elements67maintain helical elongate element52(which defines the outer edge of the impeller blade) within a given distance with respect to the central axial spring. In this manner, the elongate elements are configured to prevent the outer edge of the impeller from being forced radially outward due to forces exerted upon the impeller during the rotation of the impeller. The elongate elements are thereby configured to maintain the gap between the outer edge of the blade of the impeller and the inner surface of frame34, during rotation of the impeller. Typically, more than one (e.g., more than two) and/or fewer than eight (e.g., fewer than four) elongate elements67are used in the impeller, with each of the elongate elements typically being doubled (i.e., extending radially from central axial spring54to an outer helical elongate element52, and then returning from the helical elongate element back to the central axial spring). For some applications, a plurality of elongate elements, each of which extends from the spring to a respective helical elongate element and back to the spring, are formed from a single piece of string or a single wire.

Reference is now made toFIGS.3D and3E, which are schematic illustrations of impeller50, the impeller including a single integrated impeller-overexpansion-prevention element72that defines a plurality of elongate elements67, in accordance with some applications of the present invention. For some applications, impeller-overexpansion-prevention element72(which defines a plurality of elongate elements67) is used as an alternative to elongate elements67as shown inFIGS.3A-C. For some applications, element72defines a ring73and the plurality of elongate elements67extending radially from the ring. For some applications, rather than threading strings and/or wire around spring54, ring73of element72is placed around (and coupled to) the spring, e.g., by being placed around a tube70, which is typically disposed at the longitudinally-central location of the spring. The ends of respective elongate elements67are then coupled to respective helical elongate elements52. As described hereinabove, elongate elements67are typically flexible but are substantially non-stretchable along the axis defined by the elongate elements. Further typically, each of elongate elements67is configured to substantially not resist compression. Rather, each elongate element67is configured to exert a tensile force upon helical elongate element52that prevents helical elongate element52from moving radially outward, such that (in the absence of elongate element67) a separation between helical elongate element52and central axial spring54would be greater than a length of elongate element67. When a force is acting upon the impeller that would cause the helical elongate element52to move radially outward (in the absence of elongate element67), the impeller-overexpansion-prevention element is configured to prevent radial expansion of the impeller. Typically, a respective elongate element67is disposed within each one of the impeller blades and is configured to prevent the impeller blade from radially expanding. For some applications, element72is made of polyester, and/or another polymer or a natural material that contains fibers, and/or nitinol (or a similar shape-memory alloy).

It is noted that the scope of the present application includes using single integrated impeller-overexpansion-prevention element72with an impeller having a different construction from that shown inFIGS.3D-E. For example, the single integrated impeller-overexpansion-prevention element72could be used with an impeller having an axial structure constructed differently from spring54(but, typically nonetheless, defining a lumen therethrough, such that the impeller defines lumen62therethrough). Alternatively or additionally, the single integrated impeller-overexpansion-prevention element72could be used with the impeller described below with reference toFIGS.3F-G.

For some applications, the following assembly technique is used to manufacture the impeller while enhancing bonding of an elastomeric material that is used to form film56to the at least one helical elongate element. Typically, bonding of the elastomeric material to the at least one helical elongate element is performed in a manner that does not cause a protrusion from the effective edge of the impeller blade. Further typically, bonding of the elastomeric material to the at least one helical elongate element is performed in a manner that provides the impeller blade with a rounded outer edge, by the elastomeric material rounding edges of the helical elongate element. Proximal bushing64, distal bushing58, and helical elongate elements52are cut from a tube of shape-memory material, such as nitinol. The cutting of the tube, as well as the shape setting of the shape-memory material, is typically performed such that the helical elongate elements and the bushings are defined by a tube of shape-memory material that is cut and shape set.

For some applications, prior to being coupled to spring54, a plasma treatment is applied to the helical elongate elements. Alternatively or additionally, prior to being coupled to spring54, the helical elongate elements are coated with a coupling agent. Typically, a coupling agent is selected that has at least two functional groups that are configured to bond respectively with the helical elongate elements and with the elastomeric material. For example, a silane compound, such as n-(2-aminoethyl)-3-aminopropyltrimethoxysilane, may be used, with the silane compound containing a first functional group (e.g., (OH)) which is configured to bond with the helical elongate elements (which are typically made of an alloy, such a nitinol), and the silane compound containing a second functional group (e.g., (NH2)) which is configured to bond with the elastomeric material. Typically, the functional groups in the coupling agent are active only for a given time period (e.g., approximately an hour or less). Therefore, during this time period, a coat of elastomeric material is applied around the helical elongate elements. Typically, the coat of elastomeric material is the same elastomeric material or a similar elastomeric material to that used in film56. For example, a polycarbonate-based thermoplastic polyurethane, such as Aromatic Carbothane™ (e.g., Aromatic Carbothane™ 75A) may be used in film56, and the coating may be the same polycarbonate-based thermoplastic polyurethane, or a similar polycarbonate-based thermoplastic polyurethane, such as Pellethane® (e.g., Pellethane® 90A).

As described hereinabove, proximal bushing64, distal bushing58, and helical elongate elements52are typically cut from a tube of shape-memory material, such as nitinol. For some applications, subsequently to the coating having been applied to the helical elongate elements52, spring54is coupled to the helical elongate elements. Typically, spring54is inserted into the cut and shape-set tube, such that the spring extends along the length of the tube from at least the proximal bushing to the distal bushing. For some applications, the spring is inserted into the cut and shape-set tube while the spring is in an axially compressed state, and the spring is configured to be held in position with respect to the tube, by exerting a radial force upon the proximal and distal bushings. Alternatively or additionally, portions of the spring are welded to the proximal and distal bushings. For some applications, the spring is cut from a tube of a shape-memory material, such as nitinol. For some such applications, the spring is configured such that, when the spring is disposed in a non-radially-constrained configuration (in which the spring is typically disposed during operation of the impeller), there are substantially no gaps between windings of the spring and adjacent windings thereto.

Typically, at this stage, overexpansion-prevention element72is placed between the spring and the helical elongate elements, as described hereinabove, such that an assembly is formed that includes coated helical elongate elements52, spring54, and overexpansion-prevention element72.

For some applications, at this stage, the assembly of coated helical elongate elements52, spring54, and overexpansion-prevention element72is sprayed with a further layer of an elastomeric material. Typically, the elastomeric material that is sprayed is the same elastomeric material or a similar elastomeric material to that used as film56. For example, a polycarbonate-based thermoplastic polyurethane, such as Aromatic Carbothane™ (e.g., Aromatic Carbothane™ 75A) may be used as film56, and the sprayed material may be the same polycarbonate-based thermoplastic polyurethane, or a similar polycarbonate-based thermoplastic polyurethane, such as Pellethane® (e.g., Pellethane® 90A). For some applications, applying the spray to the helical elongate elements rounds the helical elongate elements. Typically, when the helical elongate element has a rounded cross section, the elastomeric material forms a layer having a substantially uniform thickness at the interface with the helical elongate element. For some applications, the step of applying the coat of elastomeric material to the helical elongate elements, as described above, at least partially rounds the helical elongate elements.

For some applications, subsequently to the spray having been applied, the assembly of coated helical elongate elements52, spring54, and overexpansion-prevention element72is dipped in the elastomer from which film56is made. For some applications, the material from which the film is made is an elastomer having an ultimate elongation of more than 300 percent, e.g., more than 400 percent. Typically, the material has a relatively low molecular weight. For some applications, the material has a melt flow index (which is an indirect measure of molecular weight) of at least 4, e.g., at least 4.3. For some applications, the material has an ultimate tensile strength of more than 6000 psi, e.g., more than 7000 psi, or more than 7500 psi. For some applications, the material is a polycarbonate-based thermoplastic polyurethane, e.g., a Carbothane™. For some applications, Aromatic Carbothane™ (e.g., Aromatic Carbothane™ 75A) is used. Typically, such materials combine one or more of the following properties: no outer diameter loss caused during the dip process, resistance to fatigue, resistance to becoming misshaped by being crimped, and low outer diameter loss during crimping. Subsequently, the material is cured such that it solidifies, e.g., by being left to dry. Typically, during this stage, the impeller is disposed on a mandrel, such that the mandrel passes through lumen62defined by the bushings and the spring, thereby maintaining the lumen during the drying. For some applications, while the material from which the film is made is drying, the impeller is rotated, which typically facilitates the formation of a film of material having a substantially uniform thickness within each of the impeller blades. Once the material has dried, the mandrel is typically removed from lumen62.

In accordance with the above description of the application of film56to the helical elongate elements, the scope of the present disclosure includes any technique whereby, prior to the helical elongate elements being dipped into the elastomeric material from which film56is made, additional layers of the same elastomeric material, a different elastomeric material, and/or a mediating material are applied to the helical elongate elements, whether by spraying, dipping, or a different coating method. For some applications, additional layers of elastomeric material are configured to round the helical elongate elements, and/or to act as mediators to enhance bonding between the helical elongate elements and film56of material. For some applications, a mediating material (such as silane) is configured to act as a mediator to enhance bonding between the helical elongate elements and film56of material.

Typically, impeller50is inserted into the left ventricle transcatheterally, while impeller50is in a radially-constrained configuration. In the radially-constrained configuration, both helical elongate elements52and central axial spring54are axially elongated and radially constrained. Typically, film56of the material (e.g., silicone and/or polyurethane) changes shape to conform to the shape changes of the helical elongate elements and the axial support spring, both of which support the film of material. Typically, using a spring to support the inner edge of the film allows the film to change shape without the film becoming broken or collapsing, due to the spring providing a large surface area to which the inner edge of the film bonds. For some applications, using a spring to support the inner edge of the film reduces a diameter to which the impeller can be radially constrained, relative to if, for example, a rigid shaft were to be used to support the inner edge of the film, since the diameter of the spring itself can be reduced by axially elongating the spring.

As described hereinabove, for some applications, proximal bushing64of impeller50is coupled to axial shaft92such that the axial position of the proximal bushing with respect to the shaft is fixed, and distal bushing58of the impeller is slidable with respect to the shaft. For example, the proximal bushing may be coupled to coupling element65disposed on the axial shaft (shown inFIG.4), for example via a snap-fit mechanism. For some applications, when the impeller is radially constrained for the purpose of inserting the impeller into the ventricle or for the purpose of withdrawing the impeller from the subject's body, the impeller axially elongates by the distal bushing sliding along the axial shaft distally. Alternatively (not shown), distal bushing58of the impeller is coupled to the shaft such that the axial position of the distal bushing with respect to the shaft is fixed, and proximal bushing64of the impeller is slidable with respect to the shaft. For some such applications, when the impeller is radially constrained for the purpose of inserting the impeller into the ventricle or for the purpose of withdrawing the impeller from the subject's body, the impeller axially elongates by the proximal bushing sliding along the axial shaft proximally. Subsequent to being released inside the subject's body, the impeller assumes its non-radially-constrained configuration (in which the impeller is typically disposed during operation of the impeller), which is as shown inFIGS.3A-E.

Reference is now made toFIGS.3F and3G, which are schematic illustrations of impeller50, in accordance with some applications of the present invention. The impeller as shown inFIGS.3F and3Gis generally similar to the impeller described hereinabove with reference toFIGS.3A-E, except for the differences described below.

For some applications, each of the impeller blades comprises an inner helical elongate element52i, an outer helical elongate element520, and a film56of material extending between the inner helical elongate element and the outer helical elongate element. Each blade is proximally coupled to proximal bushing64and distally coupled to distal bushing58such that, as axial shaft92(FIG.4) rotates, the blades rotate, thereby pumping blood of the subject. For example, as shown inFIG.3F, each of the inner helical elongate elements and each of the outer helical elongate elements may be proximally coupled to the proximal bushing and distally coupled to the distal bushing. Alternatively, at the proximal and/or distal end of the blade, only one of the helical elongate elements may be coupled to the bushing, and the other helical elongate element may be coupled to the former helical elongate element. Typically, the inner helical elongate element extends between proximal bushing64and distal bushing58so as to define a radial gap55between the axial shaft (which, as noted above with reference toFIG.4, may be surrounded by central axial spring54) and the inner helical elongate element. Radial gap55is open, i.e., there is no film of material extending between the axial shaft (or central axial spring) and the inner helical elongate element. Thus, there is typically a radial gap55between the central axial spring and the impeller blade. For some applications, by defining the radial gap, the efficiency with which the impeller pumps blood is increased, and/or hemolysis that is generated by the impeller is reduced relative to a generally similar impeller that does not define such a gap.

Typically, the materials used for the blades in the impeller as shown inFIGS.3F-Gare generally similar to those described hereinabove, with reference toFIGS.3A-E. For example, the films of material may be elastomeric, and/or pieces of nitinol may be embedded into the films of material. Further typically, the films of material are bonded to the helical elongate elements using generally similar techniques to those described hereinabove. For example, sutures53(FIGS.3A-C) may couple the films of material to the inner helical elongate elements and to the outer helical elongate elements.

In some applications, proximal bushing64, distal bushing58, inner helical elongate elements52i, and outer helical elongate elements520are cut from a tube of shape-memory material, such as nitinol. For some applications, the inner helical elongate elements, the outer helical elongate elements and the proximal and distal bushings are all formed from a single integral structure, for example, a single tube of a shape-memory alloy, such as nitinol. The structure is typically cut and shaped such as to define the aforementioned structures. Alternatively, the outer helical elongate elements and the proximal and distal bushings are formed from a first structure (which is a single integral structure) and the inner helical elongate elements are formed from one or more additional structures that are coupled to the first structure. For example, the outer helical elongate elements and the proximal and distal bushings may be formed from a first tube of a shape-memory alloy, such as nitinol, and the inner helical elongate elements may be cut and formed from a second tube of a shape-memory alloy, such as nitinol. Further alternatively, the inner helical elongate elements and the proximal and distal bushings are formed from a first structure (which is a single integral structure) and the outer helical elongate elements are formed from one or more additional structures that are coupled to the first structure. For example, the inner helical elongate elements and the proximal and distal bushings may be formed from a first tube of a shape-memory alloy, such as nitinol, and the outer helical elongate elements may be cut and formed from a second tube of a shape-memory alloy, such as nitinol.

As shown inFIG.3G, for some applications, the impeller additionally comprises respective elongate elements67that couple the blades to spring54(which extends between the proximal bushing and the distal bushing) and are configured to inhibit radial expansion of the impeller by resisting tensile force. For example, elongate elements67may couple the blades directly to spring54, as inFIGS.3A-C. Alternatively, as shown inFIG.3D, the impeller may comprise impeller-overexpansion-prevention element72, which comprises elongate elements67, such that the elongate elements couple the blades to the spring by virtue of being coupled to ring73.

Typically, in such applications, as shown inFIG.3G, elongate elements67are coupled to outer helical elongate elements520. Optionally, the elongate elements may loop around outer helical elongate elements520and/or pass through inner helical elongate elements52i.

Typically, for each of the blades, the distance DO between the inner helical elongate element and the outer helical elongate element increases moving, from the proximal or distal end of the blade, toward a middle portion of the blade. For example, DO may attain its maximum value midway between proximal bushing64and distal bushing58and/or at a maximal radial span of the outer helical elongate element. For some applications, the maximum value of DO is between 1 and 3.5 mm.

Reference is now made toFIGS.5A and5B, which are schematic illustrations of impeller50and frame34of ventricular assist device20, respectively in non-radially-constrained and radially-constrained states thereof, in accordance with some applications of the present invention. The impeller and the frame are typically disposed in the radially-constrained states during the transcatheteral insertion of the impeller and the frame into the subject's body, and are disposed in the non-radially-constrained states during operation of the impeller inside the subject's left ventricle. Reference is also made toFIG.5C, which is an enlarged schematic illustration of the proximal end of the frame of the ventricular assist device, in accordance with some applications of the present invention.

As indicated inFIG.5B, the frame and the impeller are typically maintained in radially-constrained configurations by delivery catheter143. Typically, in the radially-constrained configuration of the impeller, the impeller has a total length of more than 15 mm (e.g., more than 20 mm), and/or less than 30 mm (e.g., less than 25 mm), e.g., 15-30 mm, or 20-25 mm. Further typically, in the non-radially constrained configuration of the impeller, the impeller has a length of more than 8 mm (e.g., more than 10 mm), and/or less than 18 mm (e.g., less than 15 mm), e.g., 8-18 mm, or 10-15 mm. Still further typically, when the impeller and frame34are disposed in radially-constrained configurations (as shown inFIG.5B), the impeller has an outer diameter of less than 2 mm (e.g., less than 1.6 mm) and the frame has an outer diameter of less than 2.5 mm (e.g., less than 2.1 mm).

As described hereinabove, typically, axial shaft92passes through the axis of impeller50, via lumen62of the impeller. Typically, proximal bushing64of the impeller is coupled to the shaft via a coupling element65such that the axial position of the proximal bushing with respect to the shaft is fixed, and distal bushing58of the impeller is slidable with respect to the shaft. Alternatively, distal bushing58of the impeller is coupled to the shaft such that the axial position of the distal bushing with respect to the shaft is fixed, and proximal bushing64of the impeller is slidable with respect to the shaft.

The axial shaft itself is radially stabilized via a proximal radial bearing116and a distal radial bearing118. Typically, proximal bearing housing116H is disposed around, and houses, the proximal bearing, and distal bearing housing118H is disposed around, and houses, the distal bearing. For some such applications, the radial bearings and the bearing housings are made of respective, different materials from each other. For example, the radial bearings may be made of a first material that has a relatively high hardness, such as ceramic (e.g., zirconia), and the bearing housings may be made of a second material that is moldable into a desired shape, such as a metal or an alloy (e.g., stainless steel, cobalt chromium, and/or nitinol).

For some applications, axial shaft92is made of a metal or an alloy, such as stainless steel. For some such applications, the axial shaft is covered with ceramic sleeves240(e.g., zirconia sleeves) along regions of the axial shaft that come into contact with either of the proximal and distal bearings116,118during operation of the ventricular assist device. In this manner, the radial interfaces between the axial shaft and the proximal and distal bearings are ceramic-ceramic interfaces. As described in further detail herein, typically, the impeller and the axial shaft are configured to undergo axial back-and-forth motion during operation of the ventricular assist device. Therefore, for some applications, at locations along the axial shaft corresponding to each of the proximal and distal bearings, the axial shaft is covered with the ceramic sleeve along a length of more than 5 mm, e.g., more than 7 mm. In this manner, over the course of the axial back-and-forth motion of the axial shaft, the ceramic sleeves remain in contact with the radial bearings.

For some applications, along each portion of the axial shaft that is covered with a ceramic sleeve, the shaft is shaped (e.g., via milling, molding, or a different shaping process) to define one or more grooves or indents95, as shown in the transverse cross-sectional view ofFIG.5C. Alternatively or additionally (not shown), the inner surface of the ceramic sleeve is shaped to define or more grooves or indents. For some such applications, in order to bond the sleeve to the axial shaft, an adhesive is injected into the groove or indent and the adhesive then spreads from the groove or indent across the interface between the axial shaft and the sleeve.

For some applications, the proximal bearing housing116H and distal bearing housing118H perform additional functions. Referring first to the proximal bearing housing, as described hereinabove, for some applications, proximal strut junctions33of frame34are closed around the outside of the proximal bearing housing. For some applications, the outer surface of the proximal bearing housing defines grooves that are shaped such as to receive the proximal strut junctions. For example, as shown, the proximal strut junctions have widened heads, and the outer surface of the proximal bearing housing defines grooves that are shaped to conform with the widened heads of the proximal strut junctions. Typically, securing element117(which typically includes a ring) holds the strut junctions in their closed configurations around the outside of proximal bearing housing116H.

For some applications, additional portions of the ventricular assist device are coupled to the proximal bearing housing. For example, for some applications, a drive cable130extends from outside the subject's body to axial shaft92, and is coupled to the axial shaft such that the axial shaft rotates with the drive cable. Typically, the drive cable rotates within a first outer tube140, which functions as a drive-cable-bearing tube, and which extends from outside the subject's body to the proximal bearing housing. For some applications, the first outer tube is disposed within a second outer tube142(also referred to herein as a “delivery tube”), which also extends from outside the subject's body to the proximal bearing housing. For some applications, first outer tube140and/or second outer tube142is coupled to the proximal bearing housing (e.g., using an adhesive). For example, first outer tube140may be coupled to an inner surface of the proximal bearing housing, and second outer tube142may be coupled to an outer surface of the proximal bearing housing.

Referring now to distal bearing housing118H, for some applications, distal coupling portion31of frame34is coupled to an outer surface of distal bearing housing118H, e.g., via a snap-fit mechanism. For example, the outer surface of a proximal-most portion119of the distal bearing housing may include a snap-fit mechanism to which distal coupling portion31of frame34is coupled. For some applications, distal bearing118is disposed within the proximal-most portion119of the distal bearing housing, as shown inFIG.5A. As described hereinabove, for some applications, pump-outlet tube24extends to the distal end of frame34and defines lateral blood-inlet openings108. For some such applications, a coupling portion41(e.g., a tubular coupling portion) extends distally from the pump-outlet tube, and the coupling portion is coupled to the distal bearing housing in order to anchor the distal end of the pump-outlet tube. For some applications, an intermediate portion123of the distal bearing housing defines a ridged or a threaded outer surface, to which coupling portion41of the pump-outlet tube is coupled (e.g., via an adhesive). For some applications, the outer surface is ridged in order to enhance bonding between the distal bearing housing and coupling portion41of the pump-outlet tube. For some applications, the outer surface is threaded in order to enhance bonding between the distal bearing housing and coupling portion41of the pump-outlet tube and to facilitate the application of adhesive between the outer surface and coupling portion41of the pump-outlet tube, as described in further detail hereinbelow with reference toFIG.12B. For some applications, a distal portion121of the distal bearing housing is configured to stiffen a region of distal-tip element107into which the distal end of shaft92moves (e.g., axial-shaft-receiving tube126, or a portion thereof). Typically, distal-tip element107is coupled to an outer surface of distal portion121of the distal bearing housing (e.g., via adhesive). For some applications, at least a portion of the outer surface of distal portion121of the distal bearing housing is ridged and/or threaded in order to enhance bonding between distal-tip element107and the distal bearing housing.

As described above, axial shaft92is radially stabilized via proximal radial bearing116and distal radial bearing118. In turn, the axial shaft, by passing through lumen62defined by the impeller, radially stabilizes the impeller with respect to the inner surface of frame34and inner lining39, such that even a relatively small gap between the outer edge of the blade of the impeller and inner lining39(e.g., a gap that is as described above) is maintained, during rotation of the impeller, as described hereinabove. Typically, the impeller itself is not directly disposed within any radial bearings or thrust bearings. Rather, bearings116and118act as radial bearings with respect to the axial shaft.

In some embodiments, pump-head portion27(and more generally ventricular assist device20) does not include any thrust bearing that is configured to be disposed within the subject's body and that is configured to oppose thrust generated by the rotation of the impeller. For some applications, one or more thrust bearings are disposed outside the subject's body (e.g., within motor unit23, shown inFIGS.1A and7A-B), and opposition to thrust generated by the rotation of the impeller is provided solely by the one or more thrust bearings disposed outside the subject's body. For some applications, a mechanical element and/or a magnetic element is configured to maintain the impeller within a given range of axial positions. For example, a magnet (e.g., magnet82, described hereinbelow with reference toFIG.7B) that is disposed at the proximal end of the drive cable (e.g., outside the subject's body) may be configured to impart axial motion to the impeller, and/or to maintain the impeller within a given range of axial positions.

In alternate embodiments, axial shaft92is omitted, and the impeller is instead coupled to a distal portion of drive cable130; for example, the drive cable may pass through lumen62(FIG.3E) of the impeller. In other words, the distal portion of the drive cable may function as an axial shaft. It should thus be understood that throughout the present description, the distal portion of the drive cable, which may also be referred to as an “axial shaft,” may substitute for axial shaft92.

Reference is now made toFIGS.5D,5E, and5F, which are schematic illustrations of coupling element65, in accordance with some applications of the present invention. As described hereinabove, for some applications, proximal bushing64of impeller50is coupled to axial shaft92such that the axial position of the proximal bushing with respect to the shaft is fixed, and distal bushing58of the impeller is slidable with respect to the shaft. For some applications, the proximal bushing is coupled to the axial shaft via coupling element65, for example via a snap-fit mechanism. Typically, the coupling element includes a first region (or “portion”)66disposed around axial shaft92, and a second region (or “portion”)71, which may also be disposed around the axial shaft.

The coupling element is coupled to proximal bushing64at second region71. This coupling may be effected via a snap-fit mechanism, as noted above. For example, second region71may be shaped to define one or more protrusions19, proximal bushing64may be shaped to define one or more indentations18, and the proximal bushing may couple to second region71by virtue of protrusions19snapping into indentations18. Alternatively, the proximal bushing may be shaped to define protrusions19, second region71may be shaped to define indentations18, and the proximal bushing may couple to second region71by virtue of the protrusions snapping into the indentations.

The coupling element is coupled to axial shaft92at first region66. For example, for some applications, the first region of the coupling element is welded to the shaft. For other applications, the coupling element (or at least first region66) is made of a shape-memory material (e.g., a shape-memory alloy, such as nitinol or cobalt chromium). For example, the coupling element may comprise a tube of the shape-memory material that is cut to define the first and second regions. For some such applications, at least the first region of the coupling element (or the entire coupling element) is shape set to have an inner diameter that is smaller (e.g., between 0.01 and 0.1 mm smaller) than the outer diameter of the axial shaft. For example, the axial shaft may have an outer diameter of 0.9 mm and the inner diameter of the first region of the coupling element may be between 0.85 and 0.89 mm (e.g., 0.87 mm). Thus, following the placement of the first region around the axial shaft, the first region becomes radially contracted around, and thus locked in place with respect to, the axial shaft. For some applications, coupling the coupling element to the axial shaft via this method, rather than via welding, is desirable, since the coupling element and/or the axial shaft can be weakened by being heated during the welding.

For some applications, the first region of the coupling element is shaped to define one or more slits75, e.g., by virtue of comprising a tube that defines slits75. Slits75facilitate a radial expansion of the first region such that the first region is placeable around the axial shaft. Following the placement around the axial shaft, the first region may radially contract around the axial shaft, as described above.

Slits75may incorporate various features for facilitating the expansion of first region66. For example, in some embodiments, one or more of slits75are open-ended slits750, each of which has an open end. Open-ended slits750may include one or more proximally-open slits75op, which are open at the proximal end of first region66, and/or one or more distally-open slits75od, which are open at the distal end of the first region. Optionally, the length L0of each of the open-ended slits may be 5-40% of the length L1of the coupling element. Alternatively or additionally to open-ended slits750, one or more of slits75may be closed-ended slits75c, each of which does not have any open end. In some embodiments, as shown inFIGS.5D-F, closed-ended slits75calternate with open-ended slits750around the circumference of first region66.

During manufacture of the blood pump, first region66is placed around axial shaft92such that, as described above, the first region becomes radially contracted around the axial shaft. Typically, in addition to first region66, second region71is placed around the axial shaft.

Typically, the coupling element is coupled to the axial shaft in the above-described manner without altering the temperature of the coupling element or of portions thereof. Alternatively, for some applications, the temperature of the coupling element is altered in order to facilitate the coupling of the coupling element to the axial shaft. For example, the first region may be radially expanded (with slits75facilitating the radial expansion) while the temperature of the first region is below the transformation temperature of the shape-memory material, and the first region may then be placed around the axial shaft while the first region is radially expanded. In other words, prior to the radial expansion, the coupling element may be cooled to below the transformation temperature of the shape-memory material, which is typically below ambient temperature, such as to increase the flexibility of the shape-memory material. The coupling element may then be placed around the axial shaft, while it is in its expanded configuration. Subsequently, the coupling element heats up to above its transformation temperature, causing it to radially contract towards the shape to which it was shape set. As described above, at least the first region of the coupling element (or the entire coupling element) is shape set to have an inner diameter that is smaller (e.g., between 0.01 and 0.1 mm smaller) than the outer diameter of the axial shaft. Therefore, as the coupling element radially contracts it exerts inward radial pressure on the axial shaft, causing it to become locked in place with respect to the axial shaft.

Subsequent to coupling the coupling element to the axial shaft, the impeller is coupled to the axial shaft, by coupling proximal bushing64to the second region of the coupling element. As described above, this coupling may be performed via a snap-fit mechanism; for example, protrusions19may be snapped into indentations18. Thus, as the axial shaft rotates, the blades of the impeller rotate, thereby pumping blood of the subject.

In alternate embodiments, second region71is coupled to distal bushing58(e.g., via a snap-fit mechanism, as described), such that the distal bushing is fixed in place with respect to the axial shaft, and proximal bushing64is slidable along the axial shaft.

Reference is now made toFIGS.6A and6B, which are schematic illustrations of ventricular assist device20at respective stages of a motion cycle of impeller50of the ventricular assist device with respect to frame34of the ventricular assist device, in accordance with some applications of the present invention. For some applications, while the impeller is pumping blood through tube24by rotating, axial shaft92(to which the impeller is fixated) is driven to move the impeller axially back-and-forth within frame34, by the axial shaft moving in an axial back-and-forth motion, as described in further detail hereinbelow with reference toFIG.7A. Alternatively or additionally, the impeller and the axial shaft are configured to move axially back-and-forth within frame34in response to forces that are acting upon the impeller, and without requiring the axial shaft to be actively driven to move in the axial back-and-forth motion. Typically, over the course of the subject's cardiac cycle, the pressure difference between the left ventricle and the aorta varies from being approximately zero during ventricular systole (hereinafter “systole”) to a relatively large pressure difference (e.g., 50-70 mmHg) during ventricular diastole (hereinafter “diastole”). For some applications, due to the increased pressure difference that the impeller is pumping against during diastole (and due to the fact that drive cable130is stretchable), the impeller is pushed distally with respect to frame34during diastole, relative to the location of the impeller with respect to frame34during systole. In turn, since the impeller is connected to the axial shaft, the axial shaft is moved forward. During systole, the impeller (and, in turn, the axial shaft) move back to their systolic positions. In this manner, the axial back-and-forth motion of the impeller and the axial shaft is generated in a passive manner, i.e., without requiring active driving of the axial shaft and the impeller to cause them to undergo this motion.FIGS.6A and6Bshow the impeller and axial shaft disposed at respective positions within frame34during the above-described axial back-and-forth motion cycle.

For some applications, by moving in the axial back-and-forth motion, the portions of the axial shaft that are in contact with proximal bearing116and distal bearing118are constantly changing. For some such applications, in this manner, the frictional force that is exerted upon the axial shaft by the bearings is spread over a larger area of the axial shaft than if the axial shaft were not to move relative to the bearings, thereby reducing wear upon the axial shaft, ceteris paribus. Alternatively or additionally, by moving in the back-and-forth motion with respect to the bearing, the axial shaft cleans the interface between the axial shaft and the bearings from any residues, such as blood residues.

For some applications, at the proximal-most position of the impeller during its motion cycle, the proximal end of the impeller is disposed within the proximal conical section of frame34, as shown inFIG.6A. For some applications, at the distal-most position of the impeller during its motion cycle, the distal end of the impeller is disposed at the distal end of the cylindrical section of frame34. Alternatively, even at the distal-most position of the impeller during its motion cycle, the distal end of the impeller is disposed proximally to the distal end of the cylindrical section of frame34, as shown inFIG.6B. Typically, over the course of the entire cardiac cycle, the section of the impeller at which the span of the impeller is at its maximum is disposed within the cylindrical portion of the frame34. However, a proximal portion of the impeller is typically disposed within the proximal conical section of the frame during at least a portion of the cardiac cycle.

Reference is again made toFIGS.6A and6B. Typically, distal-tip element107is a single integrated element that includes both axial-shaft-receiving tube126and distal-tip portion120. Typically, the axial-shaft-receiving tube is configured to receive a distal portion of axial shaft92of the pump-head portion during axial back-and-forth motion of the axial shaft (as described in further detail hereinbelow), and/or during delivery of the ventricular assist device. (Typically, during delivery of the ventricular assist device, the frame is maintained in a radially-constrained configuration, which typically causes the axial shaft to be disposed in a different position with respect to the frame relative to its disposition with respect to the frame during operation of the ventricular assist device). For some applications, distal-tip portion120is configured to be soft, such that the distal-tip portion is configured not to injure tissue of the subject, even if the distal-tip portion comes into contact with the tissue (e.g., tissue of the left ventricle). For example, distal-tip portion120or the entire distal-tip element may be made of silicone, polyethylene terephthalate (PET) and/or polyether block amide (e.g., PEBAX®). For some applications, the distal-tip portion defines a lumen122therethrough. For some such applications, during insertion of the ventricular assist device into the left ventricle, guidewire10(FIG.1B) is first inserted into the left ventricle, for example, in accordance with known techniques. The distal-tip portion of the ventricular assist device is then guided to the left ventricle by advancing the distal-tip portion over the guidewire, with the guidewire disposed inside lumen122. For some applications, a duckbill valve390(or a different type of valve) is disposed at the distal end of lumen122of distal-tip portion120.

Typically, during the insertion of the ventricular assist device into the subject's ventricle, delivery catheter143(FIG.5B) is placed over impeller50and frame34and maintains the impeller and the frame in their radially-constrained configurations. For some applications, distal-tip element107extends distally from the delivery catheter during the insertion of the delivery catheter into the subject's ventricle, as shown inFIG.1B. For some applications, toward the proximal end of the distal-tip element, the distal-tip element has a protrusion110. Referring toFIG.5B(which shows the pump-head portion disposed inside delivery catheter143), for some applications, during the insertion of the ventricular assist device into the subject's ventricle, the delivery catheter extends until the proximal side of the protrusion, such that the delivery catheter and the protrusion form a smooth continuous surface. The distal side of protrusion110is tapered, such that the vasculature is exposed to a tapered diameter change, and is not exposed to any edges arising from a sharp change in diameter at the interface between the delivery catheter and the distal-tip element.

For some applications, distal-tip element107defines an overall curvature that is similar to that of a question mark or a tennis-racket, with the distal-tip element defining a straight proximal portion and a bulge on one side of the longitudinal axis of the straight proximal portion. Typically, the ventricular assist device is introduced into the subject's ventricle over a guidewire, as described hereinabove. Distal-tip portion120defines lumen122, such that the distal-tip portion is held in a straightened configuration during the introduction of the ventricular assist device into the subject's ventricle (e.g., as shown in the left frame ofFIG.1B). For some applications, upon the guidewire being removed, distal-tip portion is configured to assume its curved shape.

Referring again toFIGS.6A-B, for some applications, axial-shaft-receiving tube126extends proximally from distal-tip portion120of distal-tip element107. As described hereinabove, typically, the axial shaft undergoes axial back-and-forth motion during the operation of impeller50. Axial-shaft-receiving tube126defines lumen127, which is configured to receive the axial shaft when the axial shaft extends beyond distal bearing118. For some applications, the axial shaft-receiving tube defines a stopper128at its distal end, the stopper being configured to prevent advancement of the axial shaft beyond the stopper. For some applications, the stopper comprises a rigid component that is inserted (e.g., embedded) into the distal end of the shaft-receiving tube. Alternatively (not shown), the stopper comprises a shoulder between lumen127of the axial-shaft-receiving tube and lumen122of distal-tip portion120.

Typically, during normal operation of the impeller, the axial shaft does not come into contact with stopper128, even when drive cable130(shown inFIG.5A) is maximally elongated (e.g., during diastole). However, stopper128is configured to prevent the axial shaft from protruding into the tip portion when the delivery catheter is advanced over impeller50and frame34, during retraction of ventricular assist device20from the subject's ventricle. In some cases, during the advancement of the delivery catheter over the frame and the impeller, the drive cable is at risk of snapping. In the absence of stopper128, in such cases, the axial shaft may protrude into the tip portion. Stopper128prevents this from happening, even in the event that the drive cable snaps.

It is noted that, at the proximal end of frame34, proximal radial bearing116also functions as a stopper, by preventing coupling element65and/or proximal bushing64of impeller50from being able to move beyond the proximal radial bearing. Typically, during normal operation of the impeller, coupling element65and proximal bushing64do not come into contact with proximal radial bearing116. However, proximal radial bearing116is configured to prevent coupling element65and/or proximal bushing64of impeller50from migrating proximally from inside the frame, for example, when the impeller and the frame are held in radially-constrained (i.e., crimped) configurations inside delivery catheter143. Typically, the coupling element and/or the proximal bushing is proximally-extended such as to prevent a central region of the impeller (at which the span of the impeller is at its maximum) from sliding proximally into the proximal conical portion of frame34. For example, in the systolic phase of the impeller's motion cycle (shown inFIG.6A), if the impeller were to slide further proximally by more than a given amount, the coupling element would contact proximal radial bearing116, thereby preventing further proximal motion of the impeller. For some applications, the coupling element and/or the proximal bushing is proximally extended such that it has a total length of more than 1.5 mm, e.g., more than 4 mm. For some applications (not shown), a separate stopper element is disposed upon the axial shaft proximally with respect to coupling element and/or proximal bushing64. Typically, the stopper is configured as described with reference to the coupling element. Namely, if the impeller were to slide further proximally by more than a given amount, the stopper element would contact proximal radial bearing116, thereby preventing further proximal motion of the impeller.

Typically, during operation of the ventricular assist device, and throughout the axial back-and-forth motion cycle of the impeller, the impeller is disposed relatively close to the distal-tip portion. For example, the distance of the impeller to the distal-tip portion may be within the distal-most 50 percent, e.g., the distal-most 30 percent (or the distal-most 20 percent) of tube24, throughout the axial back-and-forth motion cycle of the impeller.

Reference is now made toFIGS.6C and6D, which are schematic illustrations of a ventricular assist device that includes a motion-cushioning spring68, in accordance with some applications of the present invention. As described in further detail hereinbelow, typically, during operation of the ventricular assist device (i.e., as the impeller is rotating), the impeller undergoes axial back-and-forth motion. For some applications, as the impeller undergoes the axial back-and-forth motion, the motion-cushioning spring is configured to act as a shock absorber, to provide cushioning to the motion.FIG.6Cshows the impeller at the systolic phase of the motion cycle of the impeller andFIG.6Dshows the impeller at the diastolic phase of its motion cycle. As shown, as the impeller moves distally from its systolic position to its diastolic position, the motion-cushioning spring becomes more compressed. For some applications, the impeller is configured to be radially constrained (i.e., crimped) by becoming axially elongated, and the motion-cushioning spring is configured to become compressed such as to accommodate the axial elongation of the impeller. Typically, when the impeller is in a radially-constrained configuration, during insertion of the pump head into the left ventricle, the impeller is axially elongated such that the distal end of the impeller is disposed further distally within frame34and the spring is further compressed relative to the configurations of the impeller and the spring shown inFIG.6D.

Typically, the motion-cushioning spring is disposed around axial shaft92between the distal end of the impeller (e.g., distal bushing58of the impeller) and distal bearing118. For some applications, the motion-cushioning spring is coupled to distal bearing118or the distal bearing housing118H and extends proximally over axial shaft92from the distal bearing or the distal bearing housing118H. Typically, in such cases, the motion-cushioning spring remains rotationally stationary as the impeller rotates, and the impeller is configured to rotate with respect to the motion-cushioning spring. Alternatively or additionally, the motion-cushioning spring is coupled to the distal end of the impeller (e.g., distal bushing58of the impeller) and/or extends distally over axial shaft92from the distal end of the impeller (e.g., distal bushing58of the impeller). For some such applications, the motion-cushioning spring is configured to rotate together with the impeller. Alternatively, the motion-cushioning spring extends from a radial bearing that is disposed around the distal end of the impeller (e.g., distal bushing58of the impeller), such that motion-cushioning spring remains rotationally stationary as the impeller rotates, and the impeller is configured to rotate with respect to the motion-cushioning spring.

For some applications, the motion-cushioning spring is coupled to an elastomeric material69(such as polyurethane, and/or silicone), such that at least a portion of axial shaft92that is between the distal end of the impeller and the distal radial bearing is covered by the elastomeric material. For some applications, coupling the elastomeric material to the spring reduces a risk of the generation of thrombi and/or hemolysis by the spring, relative to if the spring were not coupled to the elastomeric material. It is noted that the scope of the present disclosure includes providing the motion-cushioning spring in the absence of elastomeric material, as may be desirable in some cases.

FIGS.6C and6Dshow the spring as being coated with the elastomeric material with the elastomeric material extending between adjacent windings of the spring. Alternatively, the spring is embedded within the elastomeric material. Typically, the elastomeric material is generally similar to the elastomeric material that is used for film56of material within impeller50. Further typically, the elastomeric material is coupled to the motion-cushioning spring in a generally similar manner to that described hereinabove with reference to the coupling of the film of elastomeric material to the spring of the impeller. Typically, the elastomeric material is coupled to the motion-cushioning spring in such a manner that the elastomeric material changes shape (e.g., by stretching and compressing) to conform to shape changes that the motion-cushioning spring undergoes (e.g., when the motion-cushioning spring undergoes elongation and compression). Further typically, the elastomeric material is configured to undergo the above-described shape changes without the elastomeric material becoming broken or collapsing, and without the elastomeric material becoming creased when the spring is compressed.

Typically, purging fluid is pumped between first outer tube140and second outer tube142. Typically, within the pump head, a portion of the purging fluid flows through the lumen defined by axial shaft92and then exits the axial shaft in the vicinity of distal bearing118, in order to purge the interface between the axial shaft and the distal bearing. For some applications, the purging system is configured such that purging fluid flows proximally from the distal bearing along an interface between the axial shaft and the elastomeric material. In this manner, the interface between the axial shaft and the elastomeric material is purged and/or lubricated.

For some applications (not shown), a proximal motion-cushioning spring is disposed on the proximal side of the impeller. For some such applications, the proximal motion-cushioning spring is disposed around axial shaft92between the proximal end of the impeller (e.g., proximal bushing64of the impeller) and proximal bearing116. For some applications, the proximal motion-cushioning spring is coupled to proximal bearing116or proximal bearing housing116H and extends distally over axial shaft92from the proximal bearing or the proximal bearing housing. Typically, in such cases, the proximal motion-cushioning spring remains rotationally stationary as the impeller rotates, and the impeller is configured to rotate with respect to the motion-cushioning spring. Alternatively or additionally, the proximal motion-cushioning spring is coupled to the proximal end of the impeller (e.g., proximal bushing64of the impeller) and/or extends distally over axial shaft92from the proximal end of the impeller (e.g., proximal bushing64of the impeller). For some such applications, the motion-cushioning spring is configured to rotate together with the impeller. Alternatively, the proximal motion-cushioning spring extends from a radial bearing that is disposed around the proximal end of the impeller (e.g., proximal bushing64of the impeller), such that motion-cushioning spring remains rotationally stationary as the impeller rotates, and the impeller is configured to rotate with respect to the motion-cushioning spring.

For some applications, the pump head includes both a proximal motion-cushioning spring disposed on a proximal side of the impeller, and a distal motion-cushioning spring disposed on a distal side of the impeller, such that axial movement of the impeller in either the distal or the proximal direction is cushioned by the motion-cushioning springs.

Reference is now made toFIGS.7A-D, which are schematic illustrations of motor unit23and/or a driven-magnet unit310of ventricular assist device20, in accordance with some applications of the present invention.FIG.7Ashows an oblique view of the motor unit and the driven-magnet unit,FIG.7Bshows an exploded view of the motor unit and the driven-magnet unit,FIG.7Cshows a cross-sectional view of the driven-magnet unit, andFIG.7Dshows an oblique view of the driven-magnet unit.

Referring toFIG.7B, typically, motor unit23includes a motor74that is configured to impart rotational motion to impeller50, via drive cable130. For some applications, the motor unit includes ribs90that are configured to dissipate heat that is generated by the motor. For some applications, the motor unit includes vibration dampeners94and96that are configured to dampen vibration of the motor unit that is caused by rotational motion and/or axial back-and-forth motion of components of the ventricular assist device.

Typically, the motor is magnetically coupled to the drive cable. For some applications, the magnetic coupling of the motor to the drive cable is as shown inFIG.7B. As shown in the cross-sectional frame ofFIG.7B, a set of driving magnets (or “drive magnets”)77are coupled to the motor via a driving magnet housing78. For some applications, the driving magnet housing includes a ring (as shown), and the driving magnets are adhered to an inner surface of the ring. For some applications, a spacer85is adhered to the inner surface of the driving magnet housing, between the two driving magnets, as shown. A driven magnet82is disposed between the driving magnets such that there is axial overlap between the driving magnets and the driven magnet. The driven magnet is coupled to a pin131, which extends to beyond the distal end of driven magnet82, where the pin is coupled to the proximal end of drive cable130, as shown inFIG.7C. For example, the driven magnet may be cylindrical and define a hole therethrough, and pin131may be adhered to an inner surface of the driven magnet that defines the hole. For some applications, the driven magnet is cylindrical, and the magnet includes a North pole and a South pole, which are divided from each other along the length of the cylinder, as shown. For some applications, the driven magnet is housed inside a cylindrical housing87. Typically, pin131defines a lumen133, which comprises a portion of a continuous lumen that extends through the ventricular assist device, as described in further detail hereinbelow.

It is noted that in the application shown inFIG.7B, the driving magnets are disposed outside the driven magnet. However, the scope of the present application includes reversing the configurations of the driving magnets and the driven magnet, mutatis mutandis. For example, the proximal end of the drive cable may be coupled to two or more driven magnets, which are disposed around a driving magnet, such that there is axial overlap between the driven magnets and the driving magnet.

Typically, driven magnet82and pin131are held in axially fixed positions within motor unit23. For some applications, driven-magnet unit310includes snap-fit prongs91(shown inFIG.7C), via which the magnetic-coupling unit is reversibly coupled to motor unit23. In addition, magnetic forces between the driving magnets and the driven magnet typically hold the driven magnet in a relatively fixed axial position.

Referring now toFIG.7C, for some applications, pin131is disposed within radial bearings96, which support the pin as the pin rotates. Typically, at its distal end, the pin is coupled to the proximal end of drive cable130(e.g., using coupling techniques described hereinbelow with reference toFIGS.14A-H). Thus, the proximal end of the drive cable is typically held in an axially fixed position by the pin. Typically, drive cable130extends from pin131to axial shaft92(which is disposed within pump-head portion27) and thereby at least partially fixes the axial position of the axial shaft, and in turn impeller50. (Alternatively, as described above, the distal end of drive cable130may function as an axial shaft.) For some applications, the drive cable is somewhat stretchable. For example, the drive cable may be made of coiled wires that are stretchable. For some applications, the device does not include a thrust bearing in the pump-head portion or in the vicinity thereof. For such applications, the drive cable typically allows the axial shaft (and in turn the impeller) to assume a range of axial positions (by the drive cable becoming more or less stretched), but limits the axial motion of the axial shaft and the impeller to being within a certain range of motion (by virtue of the proximal end of the drive cable being held in an axially fixed position, and the stretchability of the drive cable being limited).

Still referring toFIG.7C, typically, the drive cable rotates within a first outer tube140, which functions as a drive-cable-bearing tube, and which extends from driven-magnet unit310(which is disposed outside the subject's body) to pump-head portion27(e.g., to proximal bearing housing116H of the pump-head portion). For some applications, the first outer tube is disposed within a second outer tube142, which also extends from driven-magnet unit310(which is disposed outside the subject's body) to pump-head portion27(e.g., to proximal bearing housing116H of the pump-head portion).

Referring again toFIG.7B, it is noted that, typically, motor74is mechanically coupled to, and configured to rotate, driving magnets77. Driven magnet82axially overlaps driving magnets77such that, as the motor rotates the driving magnets, the driven magnet rotates. Drive cable130is coupled to driven magnet82(e.g., via pin131) such that, as the driven magnet rotates, the impeller pumps blood of the subject within the body of the subject while the axial position of the impeller, and hence of the driven magnet, varies with the cardiac cycle of the subject.

For some applications, ventricular assist device20includes a magnetic sensor84configured to detect changes in the magnetic field caused by the variation in the axial position of the driven magnet. For example, sensor84may include a magnetometer (e.g., a Hall sensor) that is disposed within motor unit23, as shown inFIG.7B. (In some cases, sensor84is referred to herein as magnetometer84.)

Reference is now made toFIG.7E, which is a schematic illustration of motor unit23, in accordance with some applications of the present invention. Reference is also made toFIG.7F, which shows a cross-sectional view of a portion of motor unit23, in accordance with some embodiments of the present invention.

For some applications, a printed circuit board104is disposed inside a protective container105. For some applications, magnetometer84is housed inside a magnetometer housing112that orients the magnetometer toward the interface between drive magnets77and driven magnet82, such as to facilitate the detection of changes in magnetic flux density and/or magnetic phase by the magnetometer. For some applications, the magnetometer is disposed behind a metallic element113that is shaped to act as a passive magnetic lens, by directing the changes in magnetic flux and magnetic phase that are generated at the interface between the drive magnet and the driven magnet toward the magnetometer.

Referring now toFIG.7D, typically the ventricular assist device is delivered to the subject's ventricle via delivery catheter143. For some applications, the proximal end of the delivery catheter is sealingly coupled to driven-magnet unit310via a toric joint mechanism101. Typically, a sterile sleeve100extends proximally from the toric joint mechanism, thereby allowing relative motion between the driven-magnet unit and the proximal end of the delivery catheter, while maintaining sterility. For some applications, a fixation unit97is configured to reversibly fix the position of the driven-magnet unit relative to the proximal end of the delivery catheter. For example, once the pump-head portion has been positioned within the patient's ventricle, the fixation unit may be used to fix the position of the driven-magnet unit relative to the proximal end of the delivery catheter. For some applications, the fixation unit is configured such that the position of the driven-magnet unit is fixed relative to the proximal end of the delivery catheter by screwing a portion of the fixation unit. For example, the fixation unit may include a Tuohy Borst adapter.

Referring again toFIG.7C, for some applications, an additional sheath98covers second outer tube142within at least a portion of sterile sleeve100. For some applications, driven magnet housing87defines a valve99at its proximal end. Typically, guidewire10(shown inFIG.1B) passes out of the proximal end of the ventricular assist device via valve99. Typically, the ventricular assist device is delivered to the subject's ventricle over the guidewire as described hereinabove. Further typically, once the pump-head portion has been deployed within the ventricle, the guidewire is retracted from the proximal end of the device via valve99. Subsequently to the removal of the guidewire, the driven-magnet unit is typically coupled to motor unit23.

As described hereinabove, typically purging system29(shown inFIG.1A) is used with ventricular assist device20. Typically, driven-magnet unit310includes an inlet port86and an outlet port88, for use with the purging system. For some applications, a purging fluid is continuously or periodically pumped into the ventricular assist device via inlet port86and out of the ventricular assist device via outlet port88. For some applications, an additional purging fluid inlet port89(FIG.7D) is provided for pumping fluid between delivery catheter143and outer tube142.

Reference is now made toFIG.8A, which is a graph indicating variations in the length of a drive cable of a ventricular assist device, as a function of a varying pressure gradient against which the impeller of the ventricular assist device pumps, as measured in experiments. An impeller and a drive cable as described herein were used to pump a glycerin-based solution through chambers, with the chambers set up to replicate the left ventricle and the aorta, and the solution having properties (such as, density and viscosity) similar to those of blood. The pressure gradient against which the impeller was pumping was varied in a pulsatile manner to represent the pulsatility of the pressure gradient against which the impeller typically pumps when it is pumping blood from the left ventricle to the aorta. At the same time, movement of the drive cable was imaged and changes in the length of the drive cable were determined via analysis of the images. The graph shown inFIG.8Aindicates the changes in the length of the drive cable that were measured, as a function of the pressure gradient. As shown inFIG.8A, as the pressure gradient against which the impeller pumped increased, the drive cable became increasingly elongated. As indicated by the results shown inFIG.8Aand as described hereinabove, it is typically the case that, in response to variations in the pressure against which the impeller is pumping blood (e.g., the pressure difference between the left ventricle and the aorta), the impeller moves back and forth with respect to frame34. In turn, the movement of the impeller causes drive cable130to become more or less elongated.

For some applications, during operation of the ventricular assist device, computer processor25of control console21(FIG.1A) is configured to measure an indication of the pressure exerted upon the impeller (which is indicative of the pressure difference between the left ventricle and the aorta), by measuring an indication of tension in drive cable130, and/or axial motion of the drive cable. For some applications, based upon the measured indication, the computer processor detects events in the subject's cardiac cycle, determines the subject's left-ventricular pressure, and/or determines the subject's cardiac afterload. For some applications, the computer processor controls the rotation of the impeller, and/or the axial back-and-forth motion of the axial shaft in response thereto.

For some applications, it is the case that the axial back-and-forth motion of the impeller gives rise to a measurable back-and-forth motion of the inner, driven magnet82relative to the outer, one or more drive magnets77(FIG.7B), since the driven magnet is held in place with respect to the drive magnets via magnetic coupling, rather than rigid mechanical coupling. It is noted that typically the axial motion of the magnet is substantially less than that of the impeller, since the full range of motion of the impeller is not transmitted along the length of the drive cable (the drive cable typically being somewhat stretchable). For some applications, the magnetometer measures variations in the magnetic field that is generated by one of the magnets in order to measure the axial motion of drive cable130, and, in turn, to determine the pressure against which the impeller is pumping. For example, the inner, driven magnet82may be axially longer than the outer, driving magnets77. Due to the inner magnet being longer than the outer magnets, there are magnetic field lines that emanate from the inner magnet that do not pass to the outer magnets, and the magnetic flux generated by those field lines, as measured by the magnetometer, varies as the drive cable, and, in turn, the inner magnet moves axially. During operation, motor74rotates, creating an AC signal in the magnetometer, which typically has a frequency of between 200 Hz and 800 Hz. Typically, as the tension in the drive cable changes due to the subject's cardiac cycle, this gives rise to a low frequency envelope in the signal measured by the magnetometer, the low frequency envelope typically having a frequency of 0.5-2 Hz. For some applications, the computer processor measures the low frequency envelope, and derives the subject's cardiac cycle from the measured envelope.

For some applications, the magnetometer measurements are initially calibrated, such that the change in magnetic flux per unit change in pressure against which the impeller is pumping (i.e., per unit change in the pressure difference between the left ventricle and the aorta, or per unit change in the pressure gradient) is known. It is known that, in most subjects, at systole, the left-ventricular pressure is equal to the aortic pressure. Therefore, for some applications, the subject's aortic pressure is measured, and the subject's left-ventricular pressure at a given time is then calculated by the computer processor, based upon (a) the measured aortic pressure, and (b) the difference between the magnetic flux measured by the magnetometer at that time, and the magnetic flux measured by the magnetometer during systole (when the pressure in the left ventricle is assumed to be equal to that of the aorta). For example, the subject's aortic pressure may be measured by measuring pressure in a channel224defined by delivery catheter143, as described in further detail hereinbelow with reference toFIG.12A. For some applications, alternative or additional physiological parameters are determined using the above-described technique. For example, events in the subject's cardiac cycle and/or the subject's cardiac afterload may be determined.

For some applications, generally similar techniques to those described in the above paragraph are used, but as an alternative to or in addition to utilizing magnetometer measurements, a different parameter is measured in order to determine left ventricular blood pressure (and/or a different physiological parameter, e.g., events in the subject's cardiac cycle and/or the subject's cardiac afterload) at a given time. For example, it is typically the case that there is a relationship between the amount of power (and/or current) that is required to power the rotation of the impeller at a given rotation rate and the pressure difference that is generated by the impeller. (It is noted that some of the pressure difference that is generated by the impeller is used to overcome the pressure gradient against which the impeller is pumping, and some of the pressure difference that is generated by the impeller is used to actively pump the blood from the left ventricle to the aorta, by generating a positive pressure difference between the left ventricle and the aorta. Moreover, the relationship between the aforementioned components typically varies over the course of the cardiac cycle.) For some applications, calibration measurements are performed, such that the relationship between (a) power (and/or current) consumption by the motor that is required to rotate the impeller at a given rotation rate and (b) the pressure difference that is generated by the impeller, is known. For some applications, the subject's aortic pressure is measured, and the subject's left-ventricular pressure at a given time is then calculated by the computer processor, based upon (a) the measured aortic pressure, (b) the power (and/or current) consumption by the motor that is required to rotate the impeller at a given rotation rate at that time, and (c) the predetermined relationship between power (and/or current) consumption by the motor that is required to rotate the impeller at a given rotation rate and the pressure difference that is generated by the impeller. For some applications, the above-described technique is performed while maintaining the rotation rate of the impeller at a constant rate. Alternatively or additionally, the rotation rate of the impeller is varied, and the variation of the rotation rate of the impeller is accounted for in the above-described calculations. For some applications, alternative or additional physiological parameters are determined using the above-described technique. For example, events in the subject's cardiac cycle and/or the subject's cardiac afterload may be determined.

Typically, tube24has a known cross-sectional area (when the tube is in an open state due to blood flow through the tube). For some applications, the flow through tube24that is generated by the impeller is determined based on the determined pressure difference that is generated by the impeller, and the known cross-sectional area of the tube. For some applications, such flow calculations incorporate calibration parameters in order to account for factors such as flow resistance that are specific to the ventricular assist device (or type of ventricular assist device) upon which the calculations are performed. For some applications, the ventricular pressure-volume loop is derived, based upon the determined ventricular pressure.

For some applications, the computer processor determines one or more physiological parameters of the subject based upon data that are received via the ventricular assist device. Typically, the following parameters are detected via the ventricular assist device: arterial pressure (AP), electrical current consumption by the motor (I) and revolutions per minute (RPM) of the impeller. For some applications, the left ventricular pressure (LVP), the left ventricular end-diastolic pressure (LVEDP) and/or the pump flow rate are derived.

As described hereinabove, for some applications, the pressure difference between the left ventricle and the aorta (dP) is derived based upon electrical current consumption by the motor (I) and revolutions per minute (RPM) of the impeller. Left ventricular pressure (LVP) is derived by subtracting the pressure difference (dP) from the aortic pressure (AP). Left ventricular end-diastolic pressure (LVEDP) is then derived by analyzing the LVP curve, and pump flow rate is derived by integrating the pressure difference over time (along with empirically determined pumping performance parameters of the device).

For some applications, the dP curve is derived by determining (a) the maximum dP within a given first current time period (e.g., within the last second, within the last 0.5 seconds, or within a different time period that is typically between 0.1 seconds and 2 seconds, i.e., at a first frequency of between 0.5 Hz and 2 Hz), (b) the minimum dP within the given first current time period (e.g., within the last second, within the last 0.5 seconds, or within a different time period that is typically between 0.1 seconds and 2 seconds, i.e., at a first frequency of between 0.5 Hz and 2 Hz), and (c) the real-time value of the dP within a given second current time period (e.g., the current hundredth of second, or within a different time period that is typically between the current fiftieth of a second and the current two-hundredth of a second, i.e., at a second frequency that is typically between 50 Hz and 200 Hz). Typically, the second time period is smaller than the first time period, i.e., the second frequency is greater than the first frequency.

Typically, the value of each of (a), (b), and (c) is derived using a linear model. Further typically, within the first current time period, the real-time values of dP (which are detected at the second frequency) are corrected based upon the maximum dP and the minimum dP (which are detected at the first frequency), such as to generate a corrected dP curve. The LVP curve is then derived from the corrected dP curve. Typically, this results in more accurate values of LVP than if uncorrected real-time values of dP were to be used.

Referring again toFIG.7B, for some applications, in addition to magnetometer84, which is configured to measure the magnetic flux density generated by the driven magnet, a second magnetometer84A (e.g., a second Hall sensor) measures an indication of the magnetic flux density generated by the driving magnet. For some applications, the second magnetometer measures magnetic flux density of the motor, which is indicative of the flux density cycle of the driving magnet, since the motor directly drives the driving magnet to rotate. Typically, as the impeller rotates such as to pump blood, torque is generated upon the impeller. Further typically, the strength of the torque is dependent upon various parameters, such as the flow that is generated by the impeller, the rotational speed of the impeller, and/or the pressure gradient against which the impeller is pumping. For some applications, it is the case that the torque generated upon the impeller gives rise to a measurable torque on the inner, driven magnet82relative to the outer, drive magnets77, since the driven magnet is held in place with respect to the drive magnets via magnetic coupling, rather than rigid mechanical coupling. It is noted that typically the torque generated upon the driven magnet is substantially less than that generated upon the impeller, since the torque that is generated upon the impeller is not transmitted along the length of the drive cable. However, it is typically the case that the torque that is generated upon the impeller is at least partially transmitted to the driven magnet via the drive cable.

The torque that is transmitted to the driven magnet typically gives rise to a phase difference between the signal that is measured by magnetometer84(which measures magnetic flux density of the driven magnet) and the signal that is measured by second magnetometer84A (which measures magnetic flux density of the motor and/or the driving magnet). For some applications, as the torque upon the impeller varies, this gives rise to a variation in the phase difference between the signal that is measured by magnetometer84and the signal that is measured by second magnetometer84A. For some applications, the computer processor detects the variation in the aforementioned phase difference, and determines a physiological parameter of the subject, at least partially in response thereto. For example, at least partially based upon variations in the phase difference, the computer processor may determine the difference between the subject's left-ventricular pressure and the subject's aortic pressure, the subject's left ventricular pressure, an event in the subject's cardiac cycle, the subject's cardiac afterload, and/or a different physiological parameter. For some applications, the technique described in the present paragraph is used as an alternative to the above-described technique for using magnetic flux density measurements and/or power consumption measurements to determine physiological parameters. Alternatively, two or more of these techniques are used in combination with each other. For example, the subject's physiological parameters may be determined based upon a mathematical model that incorporates two or more measurements, and/or one of the techniques may be used to validate estimations of the subject's physiological parameters that are made using another one of the techniques.

Reference is now made toFIGS.8B and8C, which are graphs that demonstrate the correlation between the phase difference signal and the pressure gradient against which impeller50pumps, in accordance with some applications of the present invention.

The graph shown inFIG.8Bshows the results of an experiment in which a ventricular assist device as described herein was used to pump blood against respective pressure gradients within a static in-vitro system (i.e., an in-vitro system in which when each measurement was taken, the pressure gradient was constant). A linear regression model was used to estimate the pressure gradient against which the impeller was pumping based upon a combination of the phase difference signal, the magnetic flux amplitude signal, and the current consumed by the motor. The graph shown inFIG.8Bshows the estimated pressure gradient versus the measured pressure gradient. As shown, the linear regression model which incorporates phase difference measurements provides a reliable method for estimating the pressure gradient against which the impeller is pumping.

The graph shown inFIG.8Cshows the results of an experiment in which a ventricular assist device as described herein was used to pump blood against respective pressure gradients within a pulsatile in-vitro system (i.e., an in-vitro system in which the pressure gradient was varied in a pulsatile manner). A space state model was used to estimate the pressure gradient against which the impeller was pumping based upon a combination of the phase difference signal, the magnetic flux amplitude signal, and the current consumed by the motor. The graph shown inFIG.8Cshows the estimated pressure gradient overlaid upon the measured pressure gradient. As shown, a space state model which incorporates phase difference measurements provides a reliable method for estimating the pressure gradient against which the impeller is pumping.

In accordance with the above, and in accordance with some applications of the invention, a magnetic phase difference between the one or more driven magnets and the one or more drive magnets is measured, and a physiological parameter of the subject is determined, at least partially in response thereto. For example, at least partially based upon variations in the phase difference, the computer processor may determine the difference between the subject's left-ventricular pressure and the subject's aortic pressure, the subject's left-ventricular pressure, an event in the subject's cardiac cycle, the subject's cardiac afterload, and/or a different physiological parameter. For some applications, the physiological parameter is determined based upon the phase difference measurements in combination with one or more additional measurements, such as magnetic flux amplitude measurements, power consumed by the motor, and/or current consumed by the motor. Typically, such measurements are combined into a mathematical model, such as a linear regression model, and/or a space state model.

Reference is now made toFIGS.9A and9B, which are schematic illustrations of a ventricular assist device that includes one or more blood-pressure-measurement tubes222and/or fibers228, in accordance with some applications of the present invention.

FIG.9Ais a schematic illustration of a ventricular assist device that includes one or more blood-pressure-measurement tubes222, in accordance with some applications of the present invention. As described hereinabove, typically, the ventricular assist device includes pump-outlet tube24, which traverses the subject's aortic valve, such that a proximal end of the tube is disposed within the subject's aorta and a distal end of the tube is disposed within the subject's left ventricle. Typically, a blood pump (which typically includes impeller50) is disposed within the subject's left ventricle within tube24, and is configured to pump blood through tube24from the left ventricle into the subject's aorta. For some applications, ventricular blood-pressure-measurement tube222is configured to extend at least to an outer surface213of tube24, such that an opening214at the distal end of the blood-pressure-measurement tube is in direct fluid communication with the patient's bloodstream outside tube24. Typically, opening214is configured to be within the subject's left ventricle proximally to the blood pump (e.g., proximally to impeller50). A pressure sensor216(illustrated schematically inFIG.1A) measures pressure of blood within the ventricular blood-pressure-measurement tube. Typically, by measuring pressure of blood within the left-ventricular blood-pressure-measurement tube, the pressure sensor measures the subject's blood pressure outside tube24(i.e., left-ventricular blood pressure). Typically, blood-pressure-measurement tube222extends from outside the subject's body to opening214at the distal end of tube222, and pressure sensor216is disposed toward a proximal end of tube222, e.g., outside the subject's body. For some applications, computer processor25(FIG.1A) receives an indication of the measured blood pressure and controls the pumping of blood by the impeller, in response to the measured blood pressure.

For some applications, the ventricular assist device includes two or more such ventricular blood-pressure-measurement tubes222, e.g., as shown inFIG.9A. For some applications, based upon the blood pressure as measured within each of the left-ventricular blood-pressure-measurement tubes, computer processor25determines whether the opening of one of the two or more ventricular blood-pressure-measurement tubes is occluded. This may occur, for example, due to the opening coming into contact with the wall of the interventricular septum, and/or a different intraventricular portion. Typically, in response to determining that the opening of one of the two or more ventricular blood-pressure-measurement tubes is occluded, the computer processor determines the subject's left-ventricular pressure based upon the blood pressure measured within a different one of the two or more ventricular blood-pressure-measurement tubes.

For some applications, second outer tube142defines a groove215in a portion of the outer surface of the second outer tube that is configured to be disposed within tube24. Typically, during insertion of the ventricular assist device into the subject's body, the portion of ventricular blood-pressure-measurement tube222that extends from within tube24at least to an outer surface of tube24, is configured to be disposed within the groove, such that the portion of the ventricular blood-pressure-measurement tube does not protrude from the outer surface of the outer tube.

For some applications (not shown), distal portions of blood-pressure-measurement tubes222are disposed on the outside of pump-outlet tube24. For example, blood-pressure-measurement tubes222may extend from outer tube142to the proximal end of pump-outlet tube24, and thereafter the blood pressure measurement tubes may be built into the outer surface of tube pump-outlet tube24, as shown in FIG. 16D of U.S. Pat. No. 10,881,770 to Tuval, which is incorporated herein by reference, for example.

As described hereinabove, for some applications, drive cable130extends from a motor outside the subject's body to axial shaft92upon which impeller50is disposed. Typically, the drive cable is disposed within first outer tube140and second outer tube142, as described hereinabove. For some applications, a proximal portion of blood-pressure-measurement tube222comprises a channel between first outer tube140and second outer tube142, as shown in the cross-section ofFIG.9A. In this regard, it is noted that blood-pressure-measurement tube should be understood to refer to a continuous lumen extending from pressure sensor216to the outside of pump-outlet tube24within the subject's left ventricle, regardless of whether there are changes in the structure of the lumen along the length of the lumen. As described hereinabove, typically purging fluid is also pumped between outer tube140and outer tube142, and for some applications, the purging fluid is pumped via a purging-fluid channel226. Typically, blood-pressure-measurement tube222occupies more of the cross-sectional area defined between first outer tube140and second outer tube142than purging-fluid channel226, as shown inFIG.9A. For example, the ratio of (a) the cross-sectional area defined between first outer tube140and second outer tube142that is occupied by blood-pressure-measurement tube222to (b) the cross-sectional area defined between first outer tube140and second outer tube142that is occupied by purging-fluid channel226is typically more than 3:2, more than 3:1, or more than 5:1. For some applications, pressure-measurement tube222occupies a relatively large proportion of the cross-sectional area defined between first outer tube140and second outer tube142, in order for the blood pressure outside of pump-outlet tube24within the subject's left ventricle to be conveyed proximally to pressure sensor216.

Referring toFIG.9B, for some applications, an optical fiber228is configured to extend to at least an outer surface213(FIG.9A) of tube24, such that the distal end230of the optical fiber is directly exposed to the patient's bloodstream outside tube24. Typically, the optical fiber extends from a proximal end of the fiber that is outside the subject's body (e.g., within motor unit23) to distal end230. Further typically, a light source and light detector (not shown) are disposed at the proximal end of the optical fiber and are configured to detect blood pressure at the distal end of the optical fiber by directing light via the optical fiber and detecting reflected light.

Typically, distal end230of optical fiber228is configured to be within the subject's left ventricle proximal to the blood pump (e.g., proximal to impeller50). Typically, by measuring pressure of blood at distal end230of optical fiber228, the pressure sensor thereby measures the subject's blood pressure outside tube24(i.e., left-ventricular blood pressure). For some applications, computer processor25(FIG.1A) receives an indication of the measured blood pressure and controls the pumping of blood by the impeller, in response to the measured blood pressure.

For some applications, the ventricular assist device includes two or more such optical fibers228, e.g., as shown inFIG.9B. For some applications, based upon the blood pressure as measured using each of the optical fibers, computer processor25determines whether the distal end of one of the optical fibers is not exposed to the left ventricular bloodstream. This may occur, for example, due to the distal end of one of the optical fibers coming into contact with the wall of the interventricular septum, and/or a different intraventricular portion. Typically, in response to determining that the distal end of the one of the optical fibers is not exposed to the left ventricular bloodstream, the computer processor determines the subject's left-ventricular pressure based upon the blood pressure measured using a different one of the two or more optical fibers228.

For some applications, along the length of second outer tube142, the optical fibers are disposed within the second outer tube. Typically, at the distal end of second outer tube, the optical fibers are coupled to proximal conical portion36of frame34, such that the optical fibers extend radially to the outer surface of pump-outlet tube24. For example, as shown inFIG.9B, the optical fibers may be sutured or tied to the proximal conical portion of frame34, using coupling elements232(e.g., strings). For some applications (not shown), distal portions of optical fibers228are disposed on the outside of pump-outlet tube24. For example (not shown), optical fibers228may extend from second outer tube142to the proximal end of pump-outlet tube24, and thereafter the optical fibers may be coupled to the outer surface of pump-outlet tube24.

Referring to both blood-pressure measurement tube222and optical fiber228, it is noted that the distal end of the tube or the fiber is typically in direct fluid communication with the left-ventricular bloodstream of the subject at a location that is proximal to the proximal-most portion of blood-inlet opening(s)108(e.g., at least 1 cm, or at least 1.5 cm proximal to the proximal-most portion of blood-inlet opening(s)108). Thus, the distal end of the tube or the fiber is typically exposed to blood that has a pressure that reflects the blood pressure of the left ventricle itself and that is not affected by any pressure variations that are generated in the vicinity of the blood-inlet openings as a result of fluid-flow dynamics generated at the blood-inlet openings.

Referring to both blood-pressure measurement tube222and optical fiber228, for some applications, the tube or the fiber is coupled to the outer surface of second outer tube142. For some applications, the tube or the fiber terminates within the aorta (e.g., on second outer tube142or at the proximal end of pump-outlet tube24) in order to measure aortic pressure. For some applications, the tube or the fiber terminates at the outside of the pump-outlet tube, in the vicinity of frame34, and is configured to measure left-ventricular pressure. For some such applications, the computer processor determines whether the frame is properly placed with reference to the aortic valve and the left ventricle based upon the blood pressure that is measured using the tube or the fiber. For some applications, the tube or the fiber terminates at the distal bearing housing118H or at the distal-tip element107, and is configured to measure pressure in the vicinity of the blood-inlet openings. For some applications, the tube or the fiber runs along the outside of at least a portion of the pump-outlet tube (e.g. from the proximal end of the pump-outlet tube until the vicinity of frame34).

For some applications, the tube or the fiber is coupled to the outside of the pump-outlet tube by sandwiching the tube or the fiber between a polymeric membrane and the outer surface of the pump-outlet tube, using a bonding method, such as thermal boding, dipping, or gluing. For some applications, along the entire length of the overlap between the tube or the fiber and the pump-outlet tube, the tube or the fiber is bonded to the outside of the pump-outlet tube (e.g., using the aforementioned method(s)). Alternatively, the tube or the fiber is bonded to the outside of the pump-outlet tube (e.g., using the aforementioned method(s)) only at segmented locations along the length of the overlap between the tube or the fiber and the pump-outlet tube. For some applications, by bonding the tube or the fiber only at the segmented locations, the flexibility of the pump-outlet tube is greater than if the tube or the fiber were bonded to the pump-outlet tube along the entire length of the overlap between the tube or the fiber and the pump-outlet tube. As described above, for some applications, the tube or the fiber is coupled to the outside of the pump-outlet tube by sandwiching the tube or the fiber between a polymeric membrane and the outer surface of the pump-outlet tube. For some applications, the polymeric layer is applied at the segmented locations. For some such applications, a ring-shaped polymeric layer is applied around the full circumference of the pump-outlet tube at each of the segmented locations. Alternatively, as described below with reference toFIG.31, at each of the segmented locations (or along the entire the entire length of the overlap between the tube or the fiber and the pump-outlet tube) one or more bands are applied only around part of the circumference of the pump-outlet tube. For some applications, the bands are bonded to the pump-outlet tube in such a manner that they cause the pump-outlet tube to assume a desired curvature upon being deployed inside the subject's body. For some applications, the bands are bonded to the pump-outlet tube in such a manner that they cause different regions of the pump-outlet tube to assume respective degrees of curvature and/or or such that they cause the pump-outlet tube to twist, upon blood being pumped through the pump-outlet tube.

Reference is now made toFIGS.10A and10B, which are schematic illustrations of ventricular assist device20, the device including inner lining39that lines the inside of frame34that houses impeller50, in accordance with some applications of the present invention. For some applications, inner lining39is disposed inside frame34, in order to provide a smooth inner surface (e.g., a smooth inner surface having a substantially circular cross-sectional shape) through which blood is pumped by impeller. Typically, by providing a smooth surface, the covering material reduces hemolysis that is caused by the pumping of blood by the impeller, relative to if the blood were pumped between the impeller and struts of frame34. For some applications, inner lining39includes polyurethane, polyester, and/or silicone. Alternatively or additionally, the inner lining includes polyethylene terephthalate (PET) and/or polyether block amide (PEBAX®).

Typically, the inner lining is disposed over the inner surface of at least a portion of central cylindrical portion38of frame34. For some applications, pump-outlet tube24also covers central cylindrical portion38of frame34around the outside of the frame, for example, such that pump-outlet tube24and inner lining39overlap over at least 50 percent of the length of the inner lining, for example, over the entire length of the cylindrical portion of frame34, e.g., as shown inFIG.10A. For some applications, there is only partial overlap between pump-outlet tube24and inner lining39, e.g., as shown inFIG.10B. For example, pump-outlet tube24may overlap with inner lining39along less than 50 percent (e.g., along less than 25 percent) of the length of the inner lining. For some such applications, during insertion of ventricular assist device20into the subject's body, the impeller is advanced distally within frame34, such that the impeller is not disposed within the area of overlap between the pump-outlet tube and the inner lining, such that there is no longitudinal location at which the impeller, pump-outlet tube24, frame34, and inner lining39all overlap with each other. As shown inFIGS.10A and10B, for some applications, a single axially-facing blood inlet opening108is defined at the distal end of the pump-outlet tube and/or the inner lining. Alternatively, the inner lining is disposed over the inner surface of at least a portion of central cylindrical portion38of frame34, and the pump-outlet tube extends to the distal end of the frame and defines a plurality of lateral blood-inlet openings108. Such applications are described in further detail hereinbelow with reference toFIGS.11A-E, for example.

Typically, over the area of overlap between inner lining39and pump-outlet tube24, the inner lining is shaped to form a smooth surface (e.g., in order to reduce hemolysis, as described hereinabove), and pump-outlet tube24is shaped to conform with the struts of frame34(e.g., as shown in the cross-section inFIG.10A). Further typically, the inner lining has a substantially circular cross-section (for example, due to the relatively small cell width within the central cylindrical portion of the frame, as described hereinabove with reference toFIG.2). For some applications, over the area of overlap between inner lining39and pump-outlet tube24, the pump-outlet tube and the inner lining are coupled to each other, e.g., via vacuum, via an adhesive, and/or using a thermoforming procedure.

For some applications, inner lining39and pump-outlet tube24are made of different materials from each other. For example, the inner lining may be made of polyurethane, and the pump-outlet tube may be made of polyether block amide (PEBAX®). Typically, for such applications, the material from which the inner lining is made has a higher thermoforming temperature than that of the material from which the pump-outlet tube is made. Alternatively, inner lining39and pump-outlet tube24are made of the same material as each other. For example, both the inner lining and the pump-outlet tube may be made of polyurethane or polyether block amide (PEBAX®).

For some applications, the pump-outlet tube and the inner lining are bonded to each other and/or the frame in the following manner. For some applications, the inner lining is directly bonded to the inner surface of the frame before the pump-outlet tube is bonded to the outside of the frame. It is noted that, by bonding the inner lining directly to the inner surface of the frame (rather than simply bonding the inner lining to the pump-outlet tube and thereby sandwiching the frame between the inner lining to the pump-outlet tube), any air bubbles, folds, and other discontinuities in the smoothness of the surface provided by the inner lining are typically avoided. For some applications, similar techniques to those described hereinabove for enhancing bonding between the elastomeric film and the helical elongate elements of the impeller, are used to enhance bonding between the inner lining and the inner surface of the frame. For some applications, initially, the frame is treated so as to enhance bonding between the inner lining and the inner surface of the frame. For some applications, the treatment of the frame includes applying a plasma treatment to the frame (e.g., to the inner surface of the frame), dipping the frame in a coupling agent that has at least two functional groups that are configured to bond respectively with the frame and with the material form which the inner lining is made (e.g., silane solution), and/or dipping the frame in a solution that contains the material from which the inner lining is made (e.g., polyurethane solution). For some applications, the inner lining is made of an elastomeric material (e.g., polyurethane) and the coupling agent is a silane solution, such as a solution of n-(2-aminoethyl)-3-aminopropyltrimethoxysilane, with the silane containing a first functional group (e.g., (OH)) which is configured to bond with the frame (which is typically made of an alloy, such a nitinol), and the silane containing a second functional group (e.g., (NH2)) which is configured to bond with the elastomeric material.

For some applications, subsequently, a solution that contains the material from which the inner lining is made (e.g., polyurethane solution) is sprayed over the central cylindrical portion of the frame. Once the inner surface of the frame has been treated, the inner lining is bonded to the inner surface of the central cylindrical portion of the frame (e.g., to the inner surface of a central cylindrical portion of the frame). Typically, the inner lining (which is shaped as a tube), is placed over a mandrel, the frame is placed over the inner lining, and pressure is applied by a heat shrinking process. Further typically, the assembly of the inner lining and the frame is heated in an oven.

Subsequently to the inner lining having been bonded to the frame, a portion of pump-outlet tube24is placed around the outside of the frame. As described above, for some applications, inner lining39and pump-outlet tube24are made of different materials from each other. For example, the inner lining may be made of polyurethane, and the pump-outlet tube may be made of polyether block amide (PEBAX®). Typically, for such applications, the material from which the inner lining is made has a higher thermoforming temperature than that of the material from which the pump-outlet tube is made. For some applications, in order to mold pump-outlet tube24to conform with the struts of frame34, without causing the inner lining to deform, the frame is heated to a temperature that is above the thermoforming temperature of pump-outlet tube24but below the thermoforming temperature of inner lining39.

Typically, the frame is heated from inside the frame, using the mandrel. Typically, while the frame is heated to the aforementioned temperature, an outer tube (which is typically made from silicone) applies pressure to pump-outlet tube24that causes pump-outlet tube24to be pushed radially inwardly, in order to cause the pump-outlet tube to conform with the shapes of the struts of the frame, as shown in the cross-section ofFIG.10A. For some applications, during this stage, the mandrel that is placed inside the inner lining and that heats the inner lining is shorter than the length of the inner lining. The mandrel is typically placed within the inner lining such that margins are left outside of the mandrel at each of the ends of the inner lining. Typically, the inner lining acts as a shield to protect the pump-outlet tube from being overheated and becoming damaged by the heating of the mandrel. Placing the inner lining on the mandrel in the aforementioned manner prevents the mandrel from coming into direct contact with the frame and/or the pump-outlet tube. For some applications, the combination of the frame, the inner lining, and the portion of pump-outlet tube24disposed around the frame is subsequently shape set to a desired shape and dimensions using shape setting techniques that are known in the art.

Reference is now made toFIG.10C, which is a schematic illustration of an inner lining that includes an extension39eextending proximally beyond cylindrical portion38of frame34, in accordance with some applications of the present invention. For some applications, the inner lining extension39eis not coupled to the inner surface of frame34, but rather, the material that comprises the extension is free to flap within the blood flow that is generated by the impeller. In some cases, the inner lining extension increases the efficiency of pumping of the blood by the impeller, for example by rectifying non-linear flow paths of the blood that are generated by the pumping of the impeller.

Reference is now made toFIGS.11A-E, which are schematic illustrations of pump-outlet tube24or a portion thereof, the pump-outlet tube being configured to define lateral blood-inlet openings108at a distal end thereof, in accordance with some applications of the present invention. For some applications, the pump-outlet tube extends substantially until the distal end of distal conical portion40of frame34. For such applications, the pump-outlet tube typically defines a distal conical portion46which is distally facing, i.e., facing such that the narrow end of the cone is distal with respect to the wide end of the cone. Typically, the pump-outlet tube includes coupling portion41(e.g., a tubular coupling portion, as shown), which extends distally from the pump-outlet tube. As described hereinabove, the coupling portion is coupled to the distal bearing housing in order to anchor the distal end of the pump-outlet tube.

For some applications (not shown), the pump-outlet tube defines two to four lateral blood-inlet openings. Typically, for such applications, each of the blood-inlet openings defines an area of more than 20 square mm (e.g., more than 30 square mm), and/or less than 60 square mm (e.g., less than 50 square mm), e.g., 20-60 square mm, or 30-50 square mm. Alternatively or additionally, the outlet tube defines a greater number of smaller blood-inlet openings108, e.g., more than 10 blood-inlet openings, more than 50 blood-inlet openings, more than 100 blood-inlet openings, or more than 150 blood-inlet openings, e.g., 50-100 blood-inlet openings, 100-150 blood-inlet openings, or 150-200 blood-inlet openings. For some applications, the blood-inlet openings are sized such as (a) to allow blood to flow from the subject's left ventricle into the tube and (b) to block structures from the subject's left ventricle from entering into the frame. Typically, for such applications, the distal conical portion46of pump-outlet tube24is configured to reduce a risk of structures from the left ventricle (such as chordae tendineae, trabeculae carneae, and/or papillary muscles) entering into frame34and potentially being damaged by the impeller and/or the axial shaft, and/or causing damage to the left ventricular assist device. Therefore, for some applications, the blood-inlet openings are shaped such that, in at least one direction, the widths (or spans) of the openings are less than 1 mm, e.g., 0.1-1 mm, or 0.3-0.8 mm. By defining such a small width (or span), it is typically the case that structures from the left ventricle (such as chordae tendineae, trabeculae carneae, and/or papillary muscles) are blocked from entering into frame34. For some such applications, each of the blood-inlet openings defines an area of more than 0.05 square mm (e.g., more than 0.1 square mm), and/or less than 3 square mm (e.g., less than 1 square mm), e.g., 0.05-3 square mm, or 0.1-1 square mm. Alternatively, each of the blood-inlet openings defines an area of more than 0.1 square mm (e.g., more than 0.3 square mm), and/or less than 5 square mm (e.g., less than 1 square mm), e.g., 0.1-5 square mm, or 0.3-1 square mm.

Typically, the portion of the pump-outlet tube that defines the blood-inlet openings has a porosity of more than 40 percent, e.g., more than 50 percent, or more than 60 percent (where porosity is defined as the percentage of the area of this portion that is porous to blood flow). Thus, on the one hand, the blood-inlet openings are relatively small (in order to prevent structures of the left ventricular from entering the frame), but on the other hand, the porosity of the portion of the pump-outlet tube that defines the blood-inlet openings is relatively high, such as to allow sufficient blood flow into the pump-outlet tube.

For some applications, each of the blood-inlet openings has a circular or a polygonal shape. For some applications, each of the blood-inlet openings has a hexagonal shape, as shown inFIGS.11A-E. Typically, using openings having a hexagonal shape allows the portion of the pump-outlet tube that defines the blood-inlet openings to have a relatively high porosity (e.g., as described hereinabove), while providing the portion of the pump-outlet tube that defines the blood-inlet openings with sufficient material between the blood-inlet openings to prevent tearing and/or stretching of the material. As shown inFIG.11B, for some applications, a width W of gaps between adjacent hexagonal (or other polygonal) holes is more than 0.01 mm (e.g., more than 0.04 mm), and/or less than 0.1 mm (e.g., less than 0.08 mm), for example, 0.01-0.1 mm, or 0.04-0.08 mm. For some applications, the distance D between opposing sides of each of the hexagons (or other types of polygons) is more than 0.2 mm (e.g., more than 0.4 mm) and/or less than 0.8 mm (e.g., less than 0.6 mm), e.g., 0.2-0.8 mm, or 0.4-0.6 mm. As indicated inFIG.11B, typically each of the polygons encloses a circle (such that any structure that cannot pass through such a circle would be unable to pass through the polygon). Typically, the diameter of the circle enclosed by the polygon is the equivalent of distance D, e.g., more than 0.2 mm (e.g., more than 0.4 mm) and/or less than 0.8 mm (e.g., less than 0.6 mm), e.g., 0.2-0.8 mm, or 0.4-0.6 mm.

FIG.11Dshows a segment of distal conical portion46of pump-outlet tube24, in accordance with some applications of the present invention. In the view shown inFIG.11D, the segment is laid out flat for illustrative purposes. As shown inFIG.11D, for some applications, within a proximal region46P of distal conical portion46of pump-outlet tube24, the widths W1 of the gaps between the hexagonal (or other type of polygonal) holes are larger than widths W of the gaps between the hexagonal (or other type of polygonal) holes within a distal region46D of distal conical portion46of the pump-outlet tube. For some applications, the ratio of the widths of gaps between adjacent blood-inlet openings with the proximal region of the distal portion of the pump-outlet tube to the widths of gaps between adjacent blood-inlet openings within the distal region of the distal portion of the pump-outlet tube is greater than 3:2, e.g., between 3:2 and 5:2. Typically, for such applications, within proximal region46P of distal conical portion46of pump-outlet tube24, a distance D1 between opposing sides of each of the hexagons (or other type of polygons) is smaller than distance D between opposing sides of each of the hexagons (or other type of polygons) within distal region46D of distal conical portion46of the pump-outlet tube. (As described hereinabove, typically, distances D and D1 also represent the diameter of a circle that is enclosed by the respectively sized polygons.) For some applications, the ratio of the diameter of a circle enclosed by each of the blood-inlet openings with the distal region of the distal portion of the pump-outlet tube to a diameter of a circle enclosed by each of the blood-inlet openings with the proximal region of the distal portion of the pump-outlet tube is greater than 7:6, e.g., between 7:6 and 4:3. Further typically, the distal conical portion of pump-outlet tube24has a higher porosity within distal region46D than within proximal region46P of the distal conical portion46of the pump-outlet tube. For example, the ratio of the porosity within distal region46D to the porosity within proximal region46P is more than 4:3, or more than 3:2. For some applications, the proximal region extends along a length of more than 0.5 mm, and/or less than 2 mm (e.g., less than 1.5 mm), for example, between 0.5 and 2 mm or 0.5-1.5 mm. For some applications, the total length of the distal conical portion is more than 6 mm and/or or less than 12 mm (e.g., less than 10 mm), for example, 6-12 mm, or 6-10 mm.

As described hereinabove with reference toFIGS.10A-B, typically, the pump-outlet tube is coupled to frame34via heating. For some applications, within the proximal region46P of distal conical portion46of pump-outlet tube24, the gaps between the blood-inlet holes are wider and/or the blood-inlet holes are smaller than within distal region46D, and/or the porosity is lower than within distal region46D, in order to prevent and/or reduce damage (e.g., tearing, thinning, and/or stretching) that may be caused to the material that defines the blood-inlet holes during the above-described heating process. For some applications, as a result of the above-described heating process being applied, any difference between the sizes of the gaps between the blood-inlet holes and/or the sizes of the blood-inlet holes themselves and/or porosity between distal region46D and proximal region46P is reduced or even removed.

Typically, width W of the gaps between the hexagonal (or other type of polygonal) holes and distance D between opposing sides of each of the hexagons (or other type of polygons) within distal region46D of distal conical portion46of the pump-outlet tube are as described hereinabove. For some applications, width W1 of gaps between adjacent hexagonal (or other polygonal) holes within proximal region46P of distal conical portion46of pump-outlet tube24is more than 0.05 mm (e.g., more than 0.07 mm), and/or less than 0.2 mm (e.g., less than 0.15 mm), for example, 0.05-0.2 mm, or 0.07-0.15 mm. For some applications, distance D1 between opposing sides of each of the hexagons (or other types of polygons) within proximal region46P of distal conical portion46of pump-outlet tube24is more than 0.1 mm (e.g., more than 0.3 mm) and/or less than 0.6 mm (e.g., less than 0.5 mm), e.g., 0.1-0.6 mm, or 0.3-0.5 mm.

The scope of the present disclosure includes having non-uniformly sized and/or shaped lateral blood-inlet openings (e.g., circular, rectangular, polygonal, and/or hexagonal lateral blood-inlet openings), disposed in any arrangement along the distal conical portion46of the pump-outlet tube. Similarly, the scope of the present disclosure includes a distal conical portion46of the pump-outlet tube that defines lateral blood-inlet openings being arranged such that the distal conical portion has a non-uniform porosity, with the porosity varying over different regions of the distal conical portion. For some applications, the shapes and/or sizes of the lateral blood-inlet openings, and/or the porosity of the distal conical portion, is varied such as to account for varying blood flow dynamics at different regions of the distal conical portion. Alternatively or additionally, the shapes and/or sizes of the lateral blood-inlet openings, and/or the porosity of the distal conical portion, is varied such as to account for changes in the shape of the distal conical portion along its length.

For some applications, along distal conical portion46of pump-outlet tube24, the thickness of the polymeric material from which the pump-outlet tube is made is greater than the thickness in other regions of the pump-outlet tube (e.g., within the central cylindrical portion and/or the proximal conical portion of the pump-outlet tube). For some such applications, the tube is manufactured in this manner in order to prevent tearing of the tube within the distal conical portion46, which defines blood-inlet openings108, and may (in some cases) be at greater risk of tearing than other portions of the pump-outlet tube.

Reference is now made toFIG.11E, which is an enlarged schematic illustration of the interface between the distal end of pump-outlet tube24and distal-tip element107. Typically, the pump-outlet tube includes a coupling portion41(e.g., a tubular coupling portion, as shown), which extends distally from the pump-outlet tube. As described hereinabove, the coupling portion is coupled to distal bearing housing118H in order to anchor the distal end of the pump-outlet tube. Also as described hereinabove, typically, the pump-outlet tube is coupled to the outside of the central cylindrical portion of the frame. For some applications, distal conical portion46of the pump-outlet tube is not itself bonded to distal conical portion40of the frame. Rather, distal conical portion46of the pump-outlet tube is held in place with respect to distal conical portion40of the frame, by virtue of coupling portion41being coupled to distal bearing housing118H and the pump-outlet tube being coupled to the outside of the central cylindrical portion of the frame. Alternatively, the distal conical portion46of the pump-outlet tube is directly coupled to distal conical portion40of the frame (e.g., via heat shrinking).

As described hereinabove, for some applications, coupling portion41is coupled to the outer surface of portion123of distal bearing housing118H. For some applications, coupling portion41defines a hole111(e.g., toward the distal end of the coupling portion), as shown inFIG.11E. For some applications, an adhesive is applied between coupling portion41and the outer surface of portion123of distal bearing housing118H, via the hole. For some applications, the outer surface of portion123of distal bearing housing118H is threaded. Typically, the threaded outer surface allows the adhesive to gradually and uniformly spread between coupling portion41and the outer surface of portion123of distal bearing housing118H. Further typically, the coupling portion is transparent, such that the spread of the adhesive is visible through the coupling portion. Therefore, for some applications, once the adhesive has sufficiently spread between coupling portion41and the outer surface of portion123of distal bearing housing118H (e.g., once the outer surface of portion123has been covered with the adhesive), application of the adhesive is terminated.

Reference is now made toFIGS.12A,12B, and12C, which are schematic illustrations of drive cable130of ventricular assist device20, in accordance with some applications of the present invention. Typically, the rotational motion of the of the motor is transmitted to the axial shaft via the drive cable. Typically, the drive cable extends from motor unit23(which is typically disposed outside the subject's body) to the proximal end of axial shaft92(with the connection between the distal end of the drive cable and the proximal end of the axial shaft being shown in one of the enlarged portions ofFIG.5Afor example). For some applications, the drive cable includes a plurality of wires134that are disposed in a coiled configuration in order to impart sufficient strength and flexibility to the drive cable, such that a portion of the cable is able to be maintained within the aortic arch, while the cable is rotating and moving in the axial back-and-forth motion.

For some applications, the drive cable includes a plurality of coaxial layers of coiled wires. For example, as shown inFIGS.12A-C, the drive cable may include an outer layer136, an intermediate layer138, and an inner layer156, which are coaxial with each other and each of which comprises coiled wires. For some applications, the inner and outer layers are wound in a first direction (e.g., left hand wound) and the intermediate layer is wound in a different direction (e.g., right hand wound, or vice versa). For some applications, the drive cable includes a greater number or smaller number of layers of coiled wires (e.g., four layers or two layers). For some applications, each of the layers includes between 2 and 8, e.g., between 4 and 6, wires within the coil. For some such applications, a diameter of the drive cable is greater than 1.05 mm, e.g., between 1.1 and 1.4 mm, e.g., 1.1 mm, 1.2 mm, or 1.3 mm.

The drive cable is typically disposed within a first outer tube140, which is configured to remain stationary while the drive cable undergoes rotational and/or axial back-and-forth motion. The first outer tube is configured to effectively act as a bearing tube for the drive cable, along the length of the drive cable. As such, first outer tube is also referred to herein as the drive-cable-bearing tube. The drive-cable-bearing tube is described in further detail hereinbelow with reference toFIGS.12D-I. For some applications, the drive-cable-bearing tube is disposed within a second outer tube142, which is typically made of a material having greater flexibility than that of the drive-cable-bearing tube (e.g., nylon, and/or polyether block amide), and typically has a thickness that is greater than that of the drive cable bearing tube. Second outer tube142may also be referred to herein as mechanical-property-providing tube142or as delivery tube142. Second outer tube142is described in further detail hereinbelow, with reference toFIG.13.

Typically, during insertion of the impeller and the frame into the left ventricle, impeller50and frame34are maintained in radially-constrained configurations by delivery catheter143. As described hereinabove, in order for the impeller and the frame to assume non-radially-constrained configurations, the delivery catheter is retracted. For some applications, as shown inFIG.12A, the delivery catheter remains in the subject's aorta during operation of the left ventricular device, and second outer tube142is disposed inside the delivery catheter. (AlthoughFIG.12Ashows the distal end of the delivery catheter disposed within the aortic arch, for some application, the distal end of the delivery catheter is disposed within the descending aorta during operation of the left ventricular device.) For some applications, during operation of the left ventricular device, a channel224is defined between delivery catheter143and outer tube142. (It is noted that the channel as shown inFIG.12Ais not to scale, for illustrative purposes.) For some such applications, the subject's aortic blood pressure is measured by measuring the pressure of blood within channel224. For example, pressure sensor216(illustrated schematically inFIG.1A) may be in fluid communication with channel224, and may be configured to measure the subject's aortic pressure by measuring the pressure of blood within channel224. Typically, in order to retract the left ventricular device from the subject's body, the delivery catheter is advanced over the impeller and the frame, such that the impeller and the frame assume their radially-constrained configurations. The catheter is then withdrawn from the subject's body.

Reference is now made toFIG.12D, which is a schematic illustration of first outer tube140, which acts as a drive-cable-bearing tube, in accordance with some applications of the present invention. For some applications, the drive-cable-bearing tube includes an outer layer141and inner layer144, each of which is typically made of a biocompatible polymeric material, and a coil153embedded between the outer and inner layers. For some applications, outer layer141is made of Pebax, inner layer144is made of PTFE and/or polyimide (e.g., a mixture of PTFE and/or polyimide), and the coil is made of an alloy, such as stainless steel. Typically, the inner layer includes materials that are configured to provide low levels of friction and high wear resistance. Further typically, the outer layer is configured to provide additional strength to the drive-cable-bearing tube, while still providing the drive-cable-bearing tube with sufficient flexibility that it is able to conform with the curvature of the aortic arch, for example. Typically, the coil is configured such as to maintain a substantially circular cross-section for the drive-cable-bearing tube, even within regions in which the drive-cable-bearing tube undergoes a substantial curve (e.g., within the aortic arch). Typically, in the absence of the coil, the drive-cable-bearing tube would have a tendency to flatten and form an elliptical cross-section within such regions.

For some applications, within at least some portions of the drive cable, the outer coiled wires of the drive cable are swaged such as to flatten the wires in the coil. For some applications, along a portion of the drive cable that is disposed within the aortic arch, the outer coiled wires of the drive cable are swaged in the above-described manner. Typically, if the outer coiled wires are rounded, then the outer coiled wires of the drive cable contact the inner surface of the drive-cable-bearing tube only at the outside of the circular cross-section of each of the wires, such that the frictional force that each wire exerts on the inner surface of the drive-cable-bearing tube is concentrated at that location. By contrast, if the wires are flattened, then the frictional force that each wire exerts on the inner surface of the drive-cable-bearing tube is evenly spread around the full circumference of the drive cable.

Reference is now made toFIGS.12E,12F,12G, and12H, which are schematic illustrations of a laser-cut tube145that comprises at least a portion of first outer tube (i.e., drive-cable-bearing tube)140, in accordance with some applications of the present invention. For some applications, the first outer tube is made of a laser-cut shape memory material (e.g., a shape-memory alloy, such as nitinol or cobalt chromium) that is covered with and/or embedded within a polymeric material such as polyether block amide (e.g., PEBAX®) or thermoplastic polyurethane (e.g., Pellethane®).FIGS.12E-Hshow the laser-cut tube as it would appear if it were cut longitudinally and laid flat. As shown, slits, openings, curves, and/or other shapes are typically cut into the tube in order to provide the tube with flexibility. For some applications, different regions of the tube are cut differently from each other, in order to provide different regions of the tube with different respective flexibilities. For example, a region of the tube that is configured to be positioned within the aortic arch, or within the ascending aorta, may be configured to have a greater flexibility than a portion that is configured to be positioned within the descending aorta.

For some applications (not shown), the bearing tube is made of one or more layers of coiled wires (e.g., as described hereinabove with reference to the drive cable). Typically the coiled wires are covered with and/or embedded within a polymeric material such as polyether block amide (e.g., PEBAX®) or thermoplastic polyurethane (e.g., Pellethane®). For some applications, the coiled wires are flattened, for example, using swaging. For some applications, different regions of the tube are formed using different numbers of layers of wires, and or different numbers of wires within each layer, in order to provide different regions of the tube with respective flexibilities. For example, a region of the tube that is configured to be positioned within the aortic arch, or within the ascending aorta, may be configured to have a greater flexibility than a portion that is configured to be positioned within the descending aorta.

For some applications, in order to reduce frictional forces between the drive cable and the drive-cable-bearing tube, beads are disposed between the drive cable and the drive-cable-bearing tube. Typically, the beads have diameters of between 0.05 mm and 0.15 mm. Further typically, the beads are not held in fixed positions within the space between the drive cable and the drive-cable-bearing tube, but rather are able to move within this space thereby facilitating movement of the drive cable relative to the drive-cable-bearing tube.

Reference is now made toFIG.12I, which is a schematic illustration of a portion of drive cable130and drive-cable-bearing tube140, in accordance with some applications of the present invention.

For some applications, drive-cable-bearing tube140comprises one or more inwardly-facing ceramic portions and drive cable130comprises one or more outwardly-facing ceramic portions. Drive cable130is configured to pass through the drive-cable-bearing tube such that the outwardly-facing ceramic portions are aligned with the inwardly-facing ceramic portions, i.e., such that, at one or more regions, the interface between the drive cable and the inner surface of the drive-cable-bearing tube is ceramic on ceramic. Drive cable130is further configured to couple to an intracorporeal device such as a blood pump (as described above), at the distal end of the drive-cable-bearing tube, and to rotate within the drive-cable-bearing tube, thereby rotating the intracorporeal device, while the outwardly-facing ceramic portions are aligned with the inwardly-facing ceramic portions. In some embodiments, the inwardly-facing ceramic portions and/or the outwardly-facing ceramic portions comprise zirconia.

In some embodiments, the inwardly-facing ceramic portions of the drive-cable-bearing tube comprise respective ceramic sleeves147, which may line the inside of the main body of the drive-cable-bearing tube, or alternatively, may be integrated with the main body of the drive-cable-bearing tube. For example, the main body of the drive-cable-bearing tube may comprise a structure such as laser-cut tube145(or wires or a braid or a coil) that is covered with or embedded within a polymeric material148. Ceramic sleeves147may line the structure, or alternatively, replace the structure, at certain locations, e.g., as shown inFIG.12I.

Similarly, the outwardly-facing ceramic portions of the drive cable may comprise respective ceramic sleeves146, which may cover the main body of the drive cable or alternatively, may be integrated with the main body of the drive cable.

Typically, regions of drive cable130and drive-cable-bearing tube140along which the drive cable exerts strong frictional forces on the inner surface of the drive-cable-bearing tube are configured in this manner, in order to prevent wear and heating along such regions. For example, for embodiments in which the drive-cable-bearing tube is configured to pass through the aorta of the subject, the drive-cable-bearing tube may comprise the inwardly-facing ceramic portions at a section of the drive-cable-bearing tube that is configured to sit (i.e., to be disposed) within the aortic arch of the subject, and/or a section of the drive-cable-bearing tube that is configured to sit within the ascending aorta of the subject. Thus, the drive-cable-bearing tube may be inserted through the aorta such that the inwardly-facing ceramic portions are within the aortic arch and/or the ascending aorta of the subject.

In some embodiments, each of the inwardly-facing ceramic portions is 10-200% longer than each of the outwardly-facing ceramic portions. Thus, even if the drive cable moves axially within the drive-cable-bearing tube, the outwardly-facing ceramic portions may remain aligned with the inwardly-facing ceramic portions.

Reference is now made toFIG.13, which is a schematic illustration of delivery tube142, in accordance with some applications of the present invention.

As described hereinabove, drive cable130passes through delivery tube142, typically by virtue of passing through first outer tube140, which functions as a drive-cable-bearing tube, which in turn passes through the delivery tube. Typically, both drive-cable-bearing tube140and the delivery tube extend from motor unit23until the proximal bearing housing116H. Further typically, during delivery of pump-head portion27to the left ventricle, the pump-head portion, drive cable130, drive-cable-bearing tube140, and delivery tube142are disposed inside delivery catheter143, with the delivery catheter maintaining the pump head in a radially constrained configuration. Once pump-head portion27has been delivered to the left ventricle, the delivery catheter is retracted (e.g., to the descending aorta, as shown inFIG.13), which allows the pump-head portion to radially expand to the non-radially constrained configurations of the frame and impeller. Subsequently, drive cable130rotates the impeller so as to pump blood of the subject from the left ventricle into the aorta.

Typically, delivery tube142is configured to impart respective mechanical properties to respective portions of the length of the drive cable. Further typically, a purging fluid is pumped between drive-cable-bearing tube140and the delivery tube, such that the purging fluid flows to the pump head and purges interfaces between axial shaft92and the proximal and distal radial bearings116and118, for example, as described with reference toFIGS.11A-Cof US 2022/0226632 to Tuval, which is incorporated herein by reference.

As shown in the enlarged cross-section through delivery tube142, typically, the delivery tube comprises an outer layer167, with an optional braid157(e.g., a metal or alloy (e.g., a stainless steel) braid) or coil disposed within outer layer167.

For some applications, delivery tube142is configured to provide different mechanical properties to respective regions along the delivery tube. For example, outer layer167may vary along the length of the delivery tube, e.g., by virtue of having a variable composition and/or a variable thickness. Alternatively or additionally to outer layer167varying along the length of the delivery tube, braid157may have a pick density that varies along the length of the delivery tube, or the aforementioned coil (substituting for braid157) may have a pitch that varies along the length of the delivery tube. Due to this variation in outer layer167and/or braid157(or the aforementioned coil), the flexural rigidity (or “stiffness”) of the delivery tube at a first portion P1of the delivery tube, which is configured to traverse the aortic valve of the subject, is less than the flexural rigidity at a second portion P2of the delivery tube, which is configured to traverse at least a portion of the aortic arch of the subject, and the flexural rigidity at second portion P2is less than the flexural rigidity at a third portion P3of the delivery tube, which is configured to traverse the descending aorta of the subject. Advantageously, this variation in flexural rigidity facilitates the passage of the delivery tube through the aorta and into the left ventricle.

For example, outer layer167may comprise a (thermoplastic) polyurethane (e.g., Pellethane®) jacket (or “coating”) at first portion P1, a polyether block amide (e.g. PEBAX®) jacket at second portion P2, and a polyamide (e.g. Grilamid®) jacket at third portion P3.

For some applications, the ratio between the pick density of the braid within first portion P1of the delivery tube to pick density of the braid within third portion P3of the delivery tube is between 3:2 and 5:2. For some applications, the pick density of the braid within the first portion of the delivery tube is between 30 and 50 picks per inch and the pick density of the braid within the third portion of the delivery tube is between 15 and 25 picks per inch.

In some embodiments, the flexural rigidity of delivery tube142monotonically increases between first portion P1and third portion P3. For example, between a polyurethane jacket at first portion P1and a polyamide jacket at third portion P3, outer layer167may comprise a single polyether block amide jacket, or multiple polyether block amide jackets having an increasing durometer moving proximally along the tube.

For some applications, the delivery tube defines a first region R1, a second region R2, and a third region R3. First region R1 comprises first portion P1, second region R2 comprises second portion P2, and third region R3 comprises third portion P3.

Typically, first region R1 is configured to extend proximally from the proximal end of the frame (e.g., from proximal bearing housing116H) through the aortic valve and into the ascending aorta. For some applications, first region R1 has a length of between 70 mm and 100 mm, e.g., between 75 mm and 95 mm.

Typically, second region R2 extends proximally from the proximal end of first region R1 and around at least a portion of the aortic arch. For some applications, second region R2 has a length of between 40 mm and 80 mm (e.g., between 50 mm and 70 mm). Alternatively, second region R2 has a length of between 150 mm and 210 mm (e.g., between 160 mm and 200 mm). For some applications, along second region R2, there is a gradual decrease in the flexibility (i.e., a gradual increase in the flexural rigidity) of the delivery tube. For example, along second region R2, outer layer167may comprise multiple polyether block amide jackets having an increasing durometer moving proximally along the tube. Alternatively, the flexural rigidity of delivery tube142may be uniform along second region R2.

Typically, third region R3 extends from the proximal end of second region R2 along the descending aorta, out of the patient's vasculature (e.g., via a femoral puncture FP), and to motor unit23. For some applications, third region R3 has a length of between 1200 mm and 1500 mm (e.g., between 1250 mm and 1450 mm). In some embodiments, there is a gradual decrease in flexibility from the distal end of region R3 to third portion P3. For example, along region R3 distally to third portion P3, outer layer167may comprise multiple polyether block amide jackets having an increasing durometer moving proximally along the tube.

In some embodiments, the flexural rigidity at the distal end of the delivery tube, which is identified inFIG.13as a distal region R1d of first region R1 and which couples to the pump head (e.g., to the proximal end of the proximal bearing housing), is greater than the flexural rigidity at first portion P1. Advantageously, this increased flexural rigidity may facilitate the coupling of the delivery tube to the pump head. Typically, in such embodiments, the flexural rigidity at the distal end of the delivery tube is between the flexural rigidity at first portion P1and the flexural rigidity at third portion P3. For some applications, R1d has a length of between 5 mm and 15 mm (e.g., between 8 mm and 12 mm).

For example, outer layer167may comprise a polyurethane jacket at first portion P1and a polyether block amide (e.g. PEBAX®) jacket at the distal end. The outer layer may further comprise a polyamide jacket at third portion P3, such that the flexural rigidity at the distal end is between that of first portion P1and that of third portion P3.

In some such embodiments, the flexural rigidity monotonically increases between first portion P1and the distal end of the delivery tube. (The flexural rigidity may thus increase monotonically moving both proximally and distally from first portion P1.) For example, distally to first portion P1, outer layer167may comprise multiple polyether block amide jackets having an increasing durometer moving distally along the tube.

In some embodiments, the flexural rigidity at a fourth portion P4of the delivery tube, which is configured to span the point of insertion into the body of the subject (e.g., femoral puncture FP), is less than the flexural rigidity at third portion P3. Advantageously, this lesser flexural rigidity may facilitate the insertion into the body. Nonetheless, even in such embodiments, the flexural rigidity at the proximal end of the delivery tube—i.e., the flexural rigidity proximally to fourth portion P4, along the most proximal portion of third region R3, which may have a length of between 360 mm and 460 mm (e.g., between 380 mm and 440 mm)—is greater than the flexural rigidity at fourth portion P4(and, optionally, greater than that of third portion P3), in order to facilitate advancing and retracting this portion of the delivery tube with respect to a handle149, and/or in order to facilitate clamping of the delivery tube without causing the delivery tube to kink. For example, proximally to fourth portion P4, outer layer167may comprise a polyether block amide (e.g. PEBAX®) jacket with a high-performance polyamide (e.g., Grilamid®) outer coating. In some embodiments, along this most proximal portion of the delivery tube, the delivery tube comprises a coil or metal frame positioned between braid157and outer layer167, e.g., between fiber159, which is described immediately below, and outer layer167. Advantageously, the coil or metal frame may further inhibit kinking of the delivery tube.

In some embodiments, the delivery tube further comprises at least one fiber159(e.g., an aromatic polyamide (i.e., aramid) fiber) that extends along the length of delivery tube142and increases the tensile rigidity (or “tensile strength”) of the delivery tube, relative to if the delivery tube would not comprise fiber159. In other words, fiber159resists elongation, thereby inhibiting elongation of the delivery tube. Thus, braid157resists bending (the degree of this resistance optionally varying along the length of the delivery tube) and elongation, while fiber159provides additional resistance to elongation (i.e., the fiber provides additional tensile strength). For some applications, a relatively high resistance to elongation (i.e., a relatively high tensile strength) is desirable along the full length of the delivery tube, in order to prevent elongation of the delivery tube when the delivery tube is disposed inside delivery catheter143.

Typically, the tensile rigidity of the delivery tube is more uniform than the flexural rigidity of the delivery tube.

For example, (a) the ratio between the tensile strength of third portion P3and that of first portion P1may be lower than (b) the ratio between the flexural rigidity of third portion P3and that of first portion P1. For example, the ratio between the tensile strength of third portion P3and that of first portion P1is typically between 3:2 and 5:2 (e.g. 2:1), and the ratio between the flexural rigidity of third portion P3and that of first portion P1is typically between 5:2 and 7:2 (e.g., 3:1). Thus, for some applications, a ratio between the former ratio and the latter ratio is between 4:3 and 5:3, e.g., 3:2.

Alternatively or additionally, the ratio between the tensile rigidity at second portion P2and the tensile rigidity at first portion P1may be less than the ratio between the flexural rigidity at second portion P2and the flexural rigidity at first portion P1. Alternatively or additionally, the ratio between the tensile rigidity at third portion P3and the tensile rigidity at second portion P2may be less than the ratio between the flexural rigidity at third portion P3and the flexural rigidity at second portion P2. This latter ratio—or, for embodiments in which second region R2 has a variable flexural rigidity, the ratio between the flexural rigidity of third portion P3and the mean flexural rigidity of second region R2—is typically between 3:2 and 5:2 (e.g. 2:1).

Typically, notwithstanding the nonuniformity of outer layer167and/or of braid157, the delivery tube comprises a uniform inner surface. For example, braid157may be coated with a uniform polymeric inner lining155(e.g., a thermoplastic polyurethane (e.g., Pellethane®) inner lining) along the length of delivery tube142. Thus, advantageously, outer tube140(or, in some embodiments, the drive cable itself) is exposed to the uniform inner surface.

In some embodiments, outer layer167becomes at least partly fused with the more inner portions of the delivery tube142during the manufacture thereof, e.g., via a reflow soldering process.

In the context of the above discussion regarding the flexural rigidity of respective portions of delivery tube142, it is noted that a metric for quantifying flexural rigidity is Young's modulus multiplied by the second moment of area (which is also known as area moment of inertia).

Reference is now made toFIGS.14A and14B, which are schematic illustrations of an interface between drive cable130and axial shaft92, in accordance with some applications of the present invention.

As described hereinabove, typically, the drive cable, which comprises a plurality of (e.g., three layers of) coiled wires134, extends from the motor unit23(which is disposed outside the subject's body) until the axial shaft92, which is hollow. Rotational motion that is generated by the motor within the motor unit is typically transmitted to the axial shaft (and thereby transmitted to the impeller) via the drive cable.

For some applications, the drive cable is coupled to the axial shaft via welding. In some cases, however, the interface between the drive cable and the axial shaft can be weakened as a result of the heating that the drive cable and the axial shaft undergo during welding. Therefore, for some applications, the drive cable is coupled to the axial shaft via a coupling technique other than welding. To facilitate this coupling technique, the proximal end of the axial shaft may be shaped to define multiple shaft pores152.

More specifically, for some applications, a porous coupling tube150(which may be polymeric) is placed around the distal end of drive cable and the proximal end of the axial shaft. In other words, the distal end of the drive cable and the proximal end of the axial shaft are inserted into opposing ends of coupling tube150, which is shaped to define multiple coupling-tube pores173. Subsequently, while the ends of the drive cable and axial shaft are inside the coupling tube, a molten bonding material is flowed between coiled wires134at the distal end of the drive cable via coupling-tube pores173, and into the proximal end of the axial shaft via the coupling-tube pores and shaft pores152, such that, upon solidifying, the material bonds the drive cable to the axial shaft. Typically, while the molten bonding material solidifies, the distal end of the drive cable, the proximal end of the axial shaft, and the coupling tube are compressed and heated.

Typically, the drive cable is also hollow, and it is desired that the drive cable and axial shaft be shaped to define a continuous lumen even after the bonding material solidifies. Hence, prior to the flowing of the molten bonding material, the drive cable and axial shaft are placed over a mandrel, such that the continuous lumen is maintained. (This continuous lumen is shown, for example, inFIG.15Bwith reference number132.)

Typically, the molten material includes a molten polymer including, for example, polyether ether ketone (PEEK). The polymer is heated such that it passes through the coupling-tube pores173and flows between the wires of the drive cable and into the shaft. Upon drying and solidifying, the polymeric material bonds the drive cable to the axial shaft.FIG.14Ashows the porous coupling tube disposed around the distal end of the drive cable and the proximal end of the axial shaft, whereasFIG.14Bshows an exploded view (for illustrative purposes) in which the porous coupling tube is shown separately from the distal end of the drive cable and the proximal end of the axial shaft.

In some embodiments, the molten material is flowed using a heat-shrinking process. In particular, a sleeve of the bonding material (e.g., a PEEK sleeve) is placed around the coupling tube, and an outer sleeve (made of polytetrafluoroethylene, for example) is placed around the sleeve of the material. Next, heat is applied to the sleeve of the material and to the outer sleeve. The applied heat melts the sleeve of the material, thereby forming the molten material, and shrinks the outer sleeve such that the outer sleeve forces the molten material between coiled wires134and into the proximal end of the axial shaft. Following the solidification of the molten material, the outer sleeve and any protruding pieces of the solidified material may be removed.

For some applications, the axial shaft (or at least the proximal end thereof) has a smaller outer diameter than the drive cable (or at least the distal end thereof). For example, the axial shaft may have a diameter of 0.9 mm and the drive cable may have a diameter of 1.1 mm, 1.2 mm, or 1.3 mm. For some such applications, prior to the insertion of the proximal end of the axial shaft into the coupling tube, a porous adaptor tube151, which is shaped to define multiple adaptor-tube pores129, is placed around the proximal end of the axial shaft so as to add to the outer diameter of the axial shaft, e.g., such that the outer diameter of the axial shaft with porous adaptor tube151disposed around it is approximately equal to the outer diameter of the distal end of the drive cable. Porous coupling tube150is then placed over the distal end of the drive cable and over the porous adaptor tube (which, in turn, is disposed over the proximal end of the axial shaft). The molten material is then flowed into the axial shaft via adaptor-tube pores129.

In some embodiments, adaptor tube151is polymeric. In other embodiments, the adaptor tube comprises a metallic alloy.

For some applications, the distal end of the drive cable is coupled to the proximal end of the axial shaft using the following procedure. Porous adaptor tube151is placed over the proximal end of the axial shaft, and the porous coupling tube is then placed over the distal end of the drive cable and over the porous adaptor tube. The above-described assembly is then placed over a mandrel, which passes through the lumen defined by the axial shaft and the distal end of the drive cable. A polymeric material, such as polyether ether ketone, is heated such that it passes through the pores of the porous coupling tube, the porous adaptor tube, and into pores152of the axial shaft (e.g., using a heat-shrinking process). Typically, the mandrel prevents the polymeric material from entering the lumen defined by the axial shaft and the distal end of the drive cable. Upon drying and solidifying, the polymeric material bonds the drive cable to the axial shaft.

As described above, typically, delivery tube142is coupled to proximal bearing housing116H (FIG.5A). In some embodiments, after coupling the drive cable to the axial shaft, the drive cable (together with the axial shaft) is inserted through the delivery tube, such that the distal end of the drive cable is disposed within the proximal bearing housing proximally to proximal radial bearing116. Thus, advantageously, the proximal bearing housing may help prevent any weakening in the coupling between the drive cable and the axial shaft.

With reference toFIG.14B, it is noted that for some applications, prior to the insertion of the distal end of the drive cable into the coupling tube, wires134are merged together at the distal end of the drive cable so as to strengthen the distal end. For example, a molten material (e.g., PEEK) may be flowed between wires134at the distal end such that, upon solidification of the material, the distal end of the drive cable is shaped to define a tube.

Reference is now made toFIGS.14C-D, which are schematic illustrations of an interface between drive cable130and axial shaft92, in accordance with some applications of the present invention.FIG.14Cshows the porous coupling tube disposed around the distal end of the drive cable and the proximal end of the axial shaft, whereasFIG.14Dshows an exploded view (for illustrative purposes) in which the porous coupling tube is shown separately from the distal end of the drive cable and the proximal end of the axial shaft.

In some embodiments, coupling tube150is compliant so as to conform both to the larger outer diameter of the drive cable and to the smaller outer diameter of the axial shaft, thus typically obviating the need for adaptor tube151.

Reference is now made toFIGS.14E-F, which are schematic illustrations of an interface between drive cable130and axial shaft92, in accordance with some applications of the present invention.FIG.14Eshows the porous coupling tube disposed around the distal end of the drive cable and the proximal end of the axial shaft, whereasFIG.14Fshows an exploded view (for illustrative purposes) in which the porous coupling tube is shown separately from the distal end of the drive cable and the proximal end of the axial shaft.

In some embodiments, the wall of coupling tube150is shaped to define multiple tabs154. Prior to the flowing of the molten material, at least some of tabs154are pushed into shaft pores152(FIG.14D) such that these tabs protrude into the shaft pores, thereby strengthening the coupling of the drive cable to the axial shaft. For embodiments in which adaptor tube151is used, these tabs are typically pushed into the shaft pores via adaptor-tube pores129, such that the tabs protrude into the shaft pores via the adaptor-tube pores.

In some embodiments, to strengthen the coupling even further, at least two of the tabs pushed into shaft pores152have different respective orientations with respect to the longitudinal axis of the coupling tube. In other words, at least two of these tabs are rotatable about rotation axes having different respective orientations with respect to the longitudinal axis. For example, in the embodiment shown inFIGS.14E-F, one tab154ais oriented parallelly to the longitudinal axis of the coupling tube, in that tab154ais rotatable about a rotation axis139athat is perpendicular to the longitudinal axis of the coupling tube, while another tab154bis oriented perpendicularly to the longitudinal axis, in that tab154bis rotatable about another rotation axis139bthat is parallel to the longitudinal axis. Together, tabs154aand154bhelp withstand both axial and rotational forces.

In some embodiments, the tabs pushed into shaft pores152are shaped to define respective holes137, which allow the molten material to flow therethrough.

Reference is now made toFIGS.14G-H, which are schematic illustrations of an interface between drive cable130and axial shaft92, in accordance with some applications of the present invention.FIG.14Gshows the porous coupling tube disposed around the distal end of the drive cable and the proximal end of the axial shaft, whereasFIG.14Hshows an exploded view (for illustrative purposes) in which the porous coupling tube is shown separately from the distal end of the drive cable and the proximal end of the axial shaft.

In some embodiments, prior to flowing the molten material between coiled wires134at the distal end of the drive cable and into the proximal end of the axial shaft, some of tabs154are pushed between the coiled wires such that these tabs protrude between the coiled wires, thereby strengthening the coupling of the drive cable to the axial shaft. These tabs may have any shape that allows the tabs to fit between the coiled wires, which may be different from the shape of those of the tabs pushed into the shaft pores. For example, in some embodiments, U-shaped tabs154care pushed between the coiled wires.

In some embodiments, at least one of the coiled wires (in any one or more of the layers of the drive cable) is cut, at the distal end of the drive cable, so as to define one or more enlarged gaps between successive windings of the coiled wires, and some of the tabs (such as U-shaped tabs154c) are pushed into the enlarged gaps. Such cuts135are shown inFIG.14H.

It is noted that the scope of the present invention includes the use of tabs154to strengthen the coupling of the drive cable to the axial shaft even without the subsequent flowing of a molten material through coupling-tube pores173.

In some embodiments, as shown inFIGS.14E-H, coupling-tube pores173are preformed, and tabs154extend into the coupling-tube pores prior to being pushed into the shaft pores or between the coiled wires. (Thus, as the tabs are pushed, the coupling-tube pores are expanded.) In other embodiments, coupling-tube pores173are formed as the tabs are pushed, i.e., the pushing of the tabs forms coupling-tube pores173in the wall of the coupling tube.

Alternatively or additionally to coupling the distal end of the drive cable to the proximal end of the axial shaft, the techniques described above may be used to couple the proximal end of the drive cable to the distal end of another hollow shaft configured to couple the drive cable to a rotating element configured to rotate the drive cable. For example, as noted hereinabove with reference toFIG.7A, for some applications, at the proximal end of the drive cable, the drive cable is coupled to driven magnet82(which is a rotating element configured to rotate the drive cable) via pin131(which is a hollow shaft). For some applications, the proximal end of the drive cable is coupled to the distal end of pin131using generally similar techniques to those described with reference toFIGS.14A-Hin relation to bonding the distal end of the drive cable to the proximal end of axial shaft92. The scope of the present disclosure includes using the apparatus and techniques described with reference toFIGS.14A-Hto couple any two tubular and/or rod shaped components to each other, mutatis mutandis.

For some applications, the drive cable continues into the pump-head portion, rather than the drive cable ending at the coupling to the axial shaft. For some such applications, a reinforcing element is added inside the drive cable within the pump-head portion, in order to add to the rigidity of the drive cable within the pump-head portion. Typically, the reinforcing element defines a lumen therethrough, which functions as a portion of the continuous lumen described hereinbelow with reference toFIGS.15A-C.

For some applications, the ventricular assist device does not define a lumen within the pump head, i.e., the axial shaft, which is coupled to the impeller, is solid. For some such applications, the axial shaft may have a relatively small diameter, such as a diameter less than 1 mm, e.g., less than 0.8 mm, 0.6 mm, or 0.4 mm. Thus, the axial shaft may have a smaller diameter than the drive cable (which, as noted above, is typically hollow, to facilitate the flow of purging fluid therethrough). For example, the drive cable may have an outer diameter greater than 1 mm (e.g., between 1 mm and 1.5 mm), and the axial shaft may have a diameter of less than 1 mm, e.g., less than 0.8 mm, 0.6 mm, or 0.4 mm. Advantageously, as a result of the axial shaft defining a smaller diameter, the pump-head portion can be radially constrained (i.e., crimped) to a smaller diameter.

Optionally, the drive cable may be coupled to the smaller-diameter axial shaft as described above with reference toFIGS.14A-H. Alternatively or additionally, the drive cable may be coupled to the smaller-diameter axial shaft within the proximal bearing housing, such that the proximal bearing housing helps keep these two elements from separating from one another.

Reference is now made toFIGS.15A-C, which are schematic illustrations showing a lumen that is used as a guidewire lumen and as a purging fluid lumen, in accordance with some applications of the present invention. For some applications, a continuous lumen extends from the proximal end of the driven-magnet unit310(shown inFIG.15C) to the distal end of distal-tip element107(shown inFIGS.15A-B). Typically, at respective portions along the length of the ventricular assist device, the lumen is defined by respective components. Further typically, starting from its proximal end, the lumen is defined by lumen133of pin131(FIG.15C), a lumen132of drive cable130and axial shaft92(FIG.15B), and then lumen122of distal-tip element107(FIG.15A).

Referring toFIG.15A, for some such applications, the ventricular device is guided to the aorta and to the left ventricle over guidewire10. Typically, a valve160(such as a duckbill valve) is disposed at the distal end of lumen122, within distal-tip element107, and the guidewire is inserted through the valve. The guidewire passes through lumen122(of the distal-tip portion), and then passes into lumen132which is defined by the axial shaft at that point. The guidewire then continues to pass through lumen132all the way until the proximal end of the drive cable. From the proximal end of the drive cable, the guidewire passes through lumen133defined by pin131, which is disposed outside of the subject's body even after insertion of the distal end of ventricular assist device20into the subject's left ventricle. Typically, as shown inFIG.15C, a valve99(such as a duckbill valve) is disposed at the proximal end of lumen133within driven-magnet unit310, and the guidewire passes out of the proximal end of the driven-magnet unit via the valve. Typically, when the distal end of the ventricular assist device is disposed inside the subject's left ventricle, the guidewire is retracted from the subject's body by pulling the guidewire out of the proximal end of lumen133. Subsequently, the axial position of driven magnet82(within which pin131is disposed) is fixed such as to be disposed between driving magnets77, as shown inFIG.7A. For example, a portion of motor unit23in which the driven magnet is disposed may be coupled to a portion of the motor unit in which driving magnets77are disposed using snap-fit prongs91(FIG.15C).

For some applications, by using lumen132of the axial shaft and the cable in the above-described manner, it is not necessary to provide an additional guidewire guide to be used during insertion of left-ventricular assist device20. For some applications, the axial shaft and the drive cable each have outer diameters of more than 0.6 mm (e.g., more than 0.8 mm), and/or less than 1.4 mm (e.g., less than 1 mm), e.g., 0.6-1.4 mm, such as 0.6-1.2 mm or 0.8-1 mm. For some applications, the diameter of lumen132, defined by the shaft and the cable, is more than 0.3 mm (e.g., more than 0.4 mm), and/or less than 0.7 mm (e.g., less than 0.6 mm), e.g., 0.3-0.7 mm, or 0.4-0.6 mm. For some applications, drive cable130has a total length of more than 1 m (e.g., more than 1.1 m), and/or less than 1.4 m (e.g., less than 1.3 m), e.g., 1-1.4 m, or 1.1-1.3 m. Typically, the diameters of lumen122and lumen133are generally similar to that of lumen132.

For some applications, the continuous lumen is additionally used by purging system29(shown inFIG.1A) of the ventricular assist device. Typically, both the first and second outer tubes140,142remain stationary, during rotation of the drive cable. For some applications, purging system29controls the flow of a purging fluid (e.g., a fluid containing glucose or dextrose) via inlet port86and outlet port88(FIG.15C). The fluid is configured to remove air from the space between the drive cable and first outer tube140, and/or to reduce frictional forces between drive cable130(which rotates), and first outer tube140(which remains stationary during rotation of the drive cable), and/or to reduce frictional forces between axial shaft92and proximal bearing116(FIG.5A) and/or distal bearing118.

Referring toFIG.15B, for some applications, the purging fluid is pumped between the first and second outer tubes140,142, and there is an opening161within the first outer tube in the vicinity of the proximal bearing. For some applications, the purging fluid is pumped via a purging-fluid channel226defined between the first and second outer tubes. For some applications, the purging fluid flows between first outer tube140and drive cable130via opening161, as indicated by a purging-fluid-flow arrow162ainFIG.15B. In this manner, the interface between drive cable130(which rotates), and first outer tube140(which acts as the drive-cable-bearing tube and remains stationary, during rotation of the drive cable) is purged. For some applications, some of the purging fluid additionally flows to the interface between the axial shaft and proximal bearing116, thereby purging the interface (and/or reducing frictional forces at the interface), as indicated by another purging-fluid-flow arrow162binFIG.15B. Typically, the flow of the purging fluid in the direction of arrow162balso prevents blood from flowing into the interface between the axial shaft and the proximal bearing.

As described hereinabove, typically, the drive cable includes a plurality of coiled wires. For some applications, purging fluid passes into lumen132defined by the drive cable via gaps in the coiled wires. Once the purging fluid is disposed within lumen132it flows in both proximal and distal directions, as indicated by a bidirectional purging-fluid-flow arrow162cofFIG.15B. The purging fluid that flows in the distal direction typically flows out of the distal end of lumen132and toward lumen122defined by distal-tip portion120(FIG.15A). At the end of distal-tip portion, the purging fluid is typically prevented from flowing out of the distal-tip portion by valve160. Therefore, some of the purging fluid typically flows to the interface between the axial shaft and distal bearing118, thereby purging the interface (and/or reducing frictional forces at the interface), as indicated by purging-fluid-flow arrows162dinFIG.15B. Typically, the flow of the purging fluid in the direction of arrows162dalso prevents blood from flowing into the interface between the axial shaft and the distal bearing.

As described above, once the purging fluid is disposed within lumen132it flows in both proximal and distal directions, as indicated by arrow162cofFIG.15B. Referring now toFIG.15C, typically, the purging fluid is pumped into the channel between the first and second outer tubes140,142via inlet port86. Typically, purging fluid flows back toward the proximal end of the device via the drive cable. Further typically, at the proximal end of the device, the purging fluid flows out of the proximal end of lumen132and then out of the proximal end of lumen133defined by pin131. Typically, the purging fluid flows such as to purge interfaces between pin131and radial bearing96, thereby reducing frictional forces at these interfaces. For some applications, the purging fluid flows around driven magnet82, such as to reduce frictional forces that the driven magnet is exposed to. For some applications, the purging fluid then flows out of outlet port88. Typically, the purging fluid is then disposed of. Alternatively, the purging fluid is pumped back into the device, via inlet port86.

Referring again toFIG.7D, for some applications, ventricular assist device20includes an additional purging fluid inlet port89, which is typically used to pump purging fluid into channel224(FIG.12A) between delivery catheter143and second outer tube142. For some applications, rather than continuously pumping purging fluid into channel224, fluid is pumped into this channel periodically, in order to flush the channel. For some applications, port89and channel224are used for aortic pressure sensing. For example, a pressure sensor216(which is illustrated schematically inFIG.1A) may be configured to measure pressure within channel224, within port89, and/or at a different location that is in fluid communication with channel224.

Reference is now made toFIG.16A, which is a schematic illustration of a pump-head portion of a ventricular assist device that includes a distal thrust bearing260, in accordance with some applications of the present invention. Typically, bearing260functions as both a distal radial bearing and a distal thrust bearing. The upper view of the ventricular device shown inFIG.16Ashows the device in its radially-constrained (i.e., crimped) configuration, while the lower view shows the device in its non-radially constrained configuration, with the arrows representing the movement of respective portions of the device between these two configurations.

Some applications of the present invention are described hereinabove as being directed toward a ventricular assist device that does not include any thrust bearing disposed within the subject's body and that is configured to allow axial back-and-forth motion of impeller50and axial shaft92. For some alternative applications, the ventricular assist device does include a thrust bearing that is disposed distally from axial shaft92so as to inhibit distal movement of the axial shaft beyond the thrust bearing. Optionally, the thrust bearing may further inhibit proximal movement of the axial shaft, such that the thrust bearing may prevent any axial movement of the axial shaft. For example, in some embodiments, thrust bearing260prevents axial shaft92from undergoing axial motion in response to variations in the pressure gradient against which the impeller is pumping (thereby, typically, preventing the impeller from undergoing axial motion in response to variations in the pressure gradient against which the impeller is pumping).

For some applications, thrust bearing260is disposed within frame34as shown. For example, the thrust bearing may be disposed within the cylindrical portion of the frame or within the distal conical portion of the frame. For some applications, at a distal end of axial shaft, the axial shaft defines a widened portion262that is configured to engage the thrust bearing and to prevent the axial shaft (and thereby prevent the impeller) from undergoing axial motion. (In addition, the widened portion of the axial shaft is constrained radially by bearing260, such that the bearing also functions as a distal radial bearing.) For some applications, the thrust bearing is coupled to the frame via connecting struts264, which extend radially inwardly from the frame to the thrust bearing. Typically, in order to manufacture frame34, the frame is cut from a tube of a shape-memory alloy, such as nitinol. For some applications, connecting struts264are cut from the tube from which the frame is cut, such that the frame and the connecting struts form a single integral body, without requiring coupling to each other (e.g., via adhesive, welding, etc.). In general, for some applications, the frame and the connecting struts are cut from a single piece of a material such as to form a single integral body. For some applications, connecting struts264as well as thrust bearing260itself are cut from the tube from which the frame is cut, such that the frame, the connecting struts, and the thrust bearing form a single integral body, without requiring coupling to each other (e.g., via adhesive, welding, etc.). In general, for some applications, the frame, the connecting struts, and thrust bearing260itself are cut from a single piece of a material such as to form a single integral body.

Reference is now made toFIG.16B, which is a schematic illustration of pump-head portion27of ventricular assist device20, in accordance with some embodiments of the present invention. As indicated inFIG.16B, a distal portion of pump-head portion27is enlarged inFIG.16C, while a proximal portion of pump-head portion27is enlarged inFIG.16D.

Reference is now made toFIG.16C. As stated with reference toFIG.16A, although some applications of the present invention are described hereinabove as being directed toward a ventricular assist device that does not include any thrust bearing disposed within the subject's body and that is configured to allow axial back-and-forth motion of impeller50and axial shaft92, for some alternative applications, the ventricular assist device does include a thrust bearing that is configured to reduce or prevent axial motion of axial shaft92in response to variations in the pressure gradient against which the impeller is pumping (and, typically, to thereby prevent the impeller from undergoing axial motion in response to variations in the pressure gradient against which the impeller is pumping).

For some applications, a thrust bearing270is disposed within distal bearing housing118H, adjacent to the distal end of axial shaft92. Typically, the distal end of the axial shaft is configured to be in contact with the distal thrust bearing. When the impeller pumps blood in the proximal direction, this causes the axial shaft to be biased in the distal direction. The thrust bearing prevents distal movement of the axial shaft, such that the axial shaft and the impeller typically remain in axially fixed positions within frame34.

In some embodiments, thrust bearing270comprises a proximally-facing ceramic surface271comprising, for example, zirconia. A distally-facing ceramic cover273, comprising zirconia for example, covers the distal end of axial shaft92, such that ceramic cover273contacts ceramic surface271as the axial shaft rotates. Thus, advantageously, there is a low-friction interface between the axial shaft and the thrust bearing.

For some applications, the thrust bearing is made of a ceramic material (e.g., zirconia), i.e., the thrust bearing comprises a piece of ceramic comprising proximally-facing ceramic surface271. In other embodiments, the thrust bearing is made of a different material but is coated with a ceramic material, i.e., the thrust bearing comprises the proximally-facing ceramic surface by virtue of being coated with a ceramic coating.

It is noted that thrust bearing270differs from thrust bearing260ofFIG.16A, in that thrust bearing270is entirely distal to the distal end of the axial shaft. Nonetheless, the features described herein with reference to thrust bearing270, such as a low-friction ceramic-ceramic interface, may likewise be implemented for thrust bearing260.

As described hereinabove, for some such applications, the axial shaft is covered with ceramic sleeves240(e.g., zirconia sleeves) along regions of the axial shaft that come into contact with either of the proximal and distal bearings116,118during operation of the ventricular assist device. In this manner, the radial interfaces between the axial shaft and the proximal and distal bearings are ceramic-ceramic interfaces. For some applications, the ceramic sleeve at the distal end of the axial shaft extends around the distal end of the axial shaft, such that the axial interface between the axial shaft and thrust bearing270is a ceramic-ceramic interface. In other words, a ceramic cap, which comprises ceramic cover273, fits around the distal end of the axial shaft, such that the ceramic cap covers both the distal end and side wall of the axial shaft. (Thus, in effect, ceramic cover273and ceramic sleeve240are combined into a single element.) In other embodiments, as shown inFIG.16C, a separate ceramic covering is applied to the distal end of the axial shaft, such that the axial interface between the axial shaft and thrust bearing270is a ceramic-ceramic interface. In other words, ceramic cover273covers the distal end of the axial shaft, while ceramic sleeve240covers the side wall of the axial shaft. (Typically, in such embodiments, ceramic cover273is coupled to ceramic sleeve240.)

In some embodiments, axial shaft92is shaped to define a shaft lumen132d, which is the distal portion of lumen132(FIG.15B), which extends through the drive cable and axial shaft. For some applications, the thrust bearing defines a bearing lumen272therethrough. Typically, bearing lumen272functions as a portion of the continuous lumen through the device that is used for guidewire insertion and/or as a purging fluid channel, as described hereinabove with reference toFIG.15A. In other words, bearing lumen272is continuous with shaft lumen132d.

For some applications, at least a portion of bearing lumen272is frustoconically-shaped, with a wider end of the bearing lumen facing distally. In other words, at its distal end272d, the lumen defines a portion of a cone, with a wide end of the cone-portion facing the distal direction. Typically, the frustoconical portion of the lumen facilitates the advancement of the guidewire through the thrust bearing in the distal-to-proximal direction.

For some applications, thrust bearing270is configured such as to allow proximal (or “reverse”) flow of fluid (in particular, purging fluid) into the interface between the distal end of the axial shaft and the thrust bearing (e.g., between ceramic cover273and proximally-facing ceramic surface271). For some such applications, the thrust bearing defines at least one additional bearing lumen (not shown) for proximal flow of the purging fluid, through the additional bearing lumen, into the interface. In other words, to facilitate the flow of purging fluid in the distal-to-proximal direction and through the interface between the thrust bearing and the axial shaft, the thrust bearing may define one or more additional flow channels.

Reference is now made toFIG.16D.

Typically, when the ventricular assist device includes a thrust bearing within the pump-head portion, the impeller and the axial shaft do not undergo axial back-and-forth motion. This is because, when the impeller pumps blood in the proximal direction, this causes the axial shaft to be biased in the distal direction. The thrust bearing prevents distal movement of the axial shaft, such that the axial shaft and the impeller remain in axially fixed positions within frame34.

For some such applications, and even for some applications in which there is no thrust bearing (such that proximal bearing housing116H houses proximal radial bearing116but no thrust bearing, and the impeller and axial shaft undergo axial motion), drive cable130is coupled to axial shaft92within the proximal bearing housing. In other words, the interface between the distal end of drive cable130and the proximal end of axial shaft92(i.e., the location at which the drive cable and the axial shaft are coupled to each other) is within the proximal bearing housing116H. Typically, the proximal bearing housing is rigid, such that the proximal bearing housing protects the location at which the drive cable and the axial shaft are coupled to each other from being exposed to forces, such as bending forces, that could weaken the coupling and thereby damage the ventricular assist device. For example, the proximal bearing housing may inhibit bending of the drive cable within the bearing housing.

Typically, in such embodiments, the radial separation between the proximal bearing housing and the distal end of the drive cable, which is within the proximal bearing housing, is less than 2 mm, e.g., less than 1 mm. Advantageously, this small radial separation may help inhibit bending of the drive cable.

For some applications, the drive cable and the axial shaft are coupled to each other using the apparatus and methods described with reference toFIGS.14A-H. Thus, for example,FIG.16Dshows coupling tube150disposed partly within the proximal bearing housing. For some applications, coupling tube150(which is used to couple the drive cable to the axial shaft) is itself made out of ceramic (e.g., zirconia) and is configured to provide the function of sleeve240, thereby typically obviating the need for sleeve240in addition to coupling tube150.

As described above, proximal radial bearing116surrounds the axial shaft and is configured to radially stabilize the axial shaft while the axial shaft rotates. A proximal sleeve240is disposed around the axial shaft such that proximal sleeve240contacts proximal radial bearing116as the axial shaft rotates. Similarly, distal radial bearing118(FIG.16C) surrounds the axial shaft distally from the proximal radial bearing and is configured to radially stabilize the axial shaft while the axial shaft rotates. A distal sleeve240is disposed around the axial shaft distally from the proximal sleeve such that the distal sleeve contacts the distal radial bearing as the axial shaft rotates.

For some applications, when the ventricular assist device includes a thrust bearing within the pump-head portion (and therefore the impeller and axial shaft do not undergo axial back-and-forth motion), sleeves240(which are placed around the axial shaft at the interfaces between the axial shaft and the proximal and distal radial bearings) are shorter than in cases in which the ventricular assist device does not include a thrust bearing within the pump-head portion (and in which the impeller and axial shaft undergo axial back-and-forth motion). This is because the portion of the axial shaft that is located at the interfaces with the proximal and distal radial bearings is relatively fixed. For some applications, the length of each of the proximal and distal radial bearings are between 2 and 4 mm (e.g., approximately 3 mm) and the length of each of the sleeves is between 2.5 and 4.5 mm (e.g., approximately 3.5 mm). For some applications, the length of each of the sleeves is between 4 and 6 mm (e.g., approximately 5 mm), in order to provide a larger margin at the ends of the radial bearings.

For some applications, the axial shaft is more flexible than the sleeves. For example, the sleeves may be made of a ceramic material (e.g., zirconia), and the axial shaft may be made of a material more flexible than ceramic, such as nitinol or another shape-memory alloy. Typically, it is desirable for the axial shaft to be sufficiently flexible as to pass through curved portions of the vasculature during delivery of the device to the left ventricle, but also sufficiently rigid to be able to provide support to the impeller during the operation of the impeller without the axial shaft undergoing vibration.

For some such applications, the lengths of the sleeves relative to the length of the shaft is configured to provide a desired overall flexibility to the axial shaft and/or to provide a desired flexibility to respective portions of the axial shaft. For example, the ratio between the length of the axial shaft and the combined length of the proximal and distal sleeves may be more than 2:1 (e.g., more than 3:1) and/or less than 6:1 (e.g., less than 5:1), for example between 2:1 and 6:1 or between 3:1 and 5:1. For some such applications, the length of the axial shaft is between 30 and 50 mm, e.g., between 35 mm and 45 mm. As described hereinabove, the length of each of the proximal and distal sleeves may be between 4 and 6 mm (e.g., approximately 5 mm), such that the combined length of the sleeves is 8-12 mm, in order to provide a larger margin at the ends of the radial bearings.

Reference is now made toFIG.16E, which is a schematic illustration of pump-head portion27of ventricular assist device20that includes a thrust bearing274, in accordance with some applications of the present invention, in accordance with some applications of the present invention. For some applications, thrust bearing274is disposed within proximal bearing housing116H, adjacent to the proximal portion of axial shaft92. For some such applications, a flange276extends radially from the proximal portion of axial shaft92. For some applications, the thrust bearing is made of a ceramic material (e.g., zirconia). For some such applications, flange276is coated with (or made of) a ceramic material (e.g., zirconia). Thus, the interface between the flange and the thrust bearing is a ceramic-ceramic interface. Typically, the flange is configured to be in contact with the proximal thrust bearing. When the impeller pumps blood in the proximal direction, this causes the axial shaft to be biased in the distal direction. The thrust bearing prevents distal movement of the axial shaft, by holding the flange in place, such that the axial shaft and the impeller remain in axially fixed positions within frame34. For some applications, thrust bearing274comprises a first portion of a shaped element (e.g., a shaped ceramic element), with the shaped element defining a second portion that functions as proximal radial bearing116.

Reference is now made toFIGS.17A,17B,17C, and17D, which are schematic illustrations of a distal-tip portion120of ventricular assist device20, in accordance with some applications of the present invention. As described hereinabove, distal-tip portion120typically forms a portion of distal-tip element107which also includes axial-shaft-receiving tube126. Typically, distal-tip element107is configured such that in its non-constrained configuration (i.e., in the absence of any forces acting upon the distal-tip portion), the distal-tip element is at least partially curved. Referring toFIG.17A, for some applications, within a given plane, distal-tip element107has a proximal, straight portion346(at least a portion of which typically comprises axial-shaft-receiving tube126). The proximal straight portion of distal-tip element107defines a longitudinal axis348. The curved portion of distal-tip element107curves away from longitudinal axis348in a first direction, and then passes through an inflection point and curves in the opposite direction with respect to longitudinal axis348. For example, as shown inFIG.17A, within the plane of the page, the distal-tip element first curves to the top of the page, then curves to the bottom of the page. Typically, when shaped as shown inFIG.17A, the distal-tip element defines an overall curvature that is similar to that of a question mark or a tennis-racket, the distal-tip element defining a bulge351on one side of the longitudinal axis of the straight proximal straight portion of the distal-tip element. For some applications, the bulge is generally shaped as a semi-ellipse. (It is noted that in this context, the term “semi-ellipse” includes a semi-circle. It is further noted that in some cases, the tip does not define a precise semi-ellipse, but rather a bulged shape that is substantially similar to a semi-ellipse.) Typically, upon being deployed within the subject's left ventricle, the curvature of the curved portion of distal-tip element107is configured to provide an atraumatic tip to ventricular assist device20. Further typically, the distal-tip element is configured to space the inlet openings108of the ventricular assist device from walls of the left ventricle.

Referring now toFIGS.17B and17C, it is first noted that these figures show a cross-sectional view of the left ventricle22in which septal wall338is disposed on the left of the page and free wall334is disposed on the right of the page. In this view, the left atrium359, and left atrial appendage358are visible above the left ventricle, and right ventricle340is visible to the left of the left ventricle. It is noted that the view of the aorta and the left ventricle as shown inFIGS.17B and17Cis different to that shown, for example, inFIG.1B.FIG.1Bis a schematic illustration, provided for illustrative purposes, and does not properly depict the scale and orientation of the ventricular assist device with respect to the anatomy.

For some applications, distal-tip element107is configured to separate the blood inlet opening from a posterior wall of the subject's left ventricle when the distal-tip element is placed against the apex of the subject's left ventricle. Typically, the distal-tip element is configured to separate the blood-inlet opening from a septal wall of the subject's left ventricle as the distal-tip element contacts the apex of the subject's left ventricle.

Typically, distal-tip element107is inserted into the left ventricle, such that bulge351bulges toward the septal wall338. When disposed in this configuration, in response to distal-tip element107being pushed against the apex (e.g., due to a physician advancing the device or in response to movement of the left ventricle), blood inlet openings108typically get pushed in the direction of free wall334and away from the septal wall338(in the direction of the arrows shown inFIG.17C). Typically, this is due to proximal straight portion346pivoting about the curved portion of the question mark shape, as shown. By contrast, other shapes of tips, if disposed in a similar orientation, may result in the blood inlet openings being pushed toward the septal wall. For example, if the distal-tip element were to have a pigtail tip (in which the tip curves in a single direction of curvature) that is oriented such that the pigtail curve is on the free wall side of the longitudinal axis of the straight portion of the distal-tip element, then pushing the tip distally would typically cause the blood inlet openings to move toward the septal wall due to the loop of the pigtail curve tightening. Notwithstanding the above, however, it is noted that in some embodiments, distal-tip element107has a pigtail shape.

Referring now toFIG.17D, for some applications, in addition to the above-described curvature of the distal-tip portion, the distal-tip portion has an additional curvature in a different plane from the above-described curvature. In particular, in addition to a distal curved portion349, the distal-tip portion comprises a proximal curved portion350, which does not lie in the same plane as the distal curved portion. (The two planes are different in the sense that they are not parallel to one another.)

In some embodiments, the tip portion further comprises a straight portion353proximal to proximal curved portion350. Straight portion353may be configured to receive the axial shaft while the axial shaft rotates, as described above, with reference toFIG.17A, for straight portion346. Alternatively or additionally, purging fluid may flow through straight portion353. For some applications, purging fluid flows until the distal end of the distal-tip portion, at which point valve160(FIG.15A, e.g., a duckbill valve) as described herein prevents the purging fluid from flowing out of the distal end of the distal-tip portion.

In some embodiments, the angle between the distal end of the proximal curved portion and the axial shaft is between 100 and 170 degrees. Alternatively or additionally, the angle between the distal end of the proximal curved portion and the plane in which distal curved portion349lies is between 25 and 65 degrees, such as between 35 and 55 degrees. For some applications, the angle between the distal end of the proximal curved portion and the plane in which distal curved portion349lies is 60 and 120 degrees.

For example, as shown inFIG.17D, proximal curved portion350may curve toward the apex342of the left ventricle, otherwise known as the apex of the heart. In other words, the proximal curved portion may be curved such that, following the insertion of the axial shaft into the left ventricle via the aorta, the distal end of the proximal curved portion (where the proximal curved portion meets distal curved portion349) points toward apex342.

In some embodiments, distally to proximal curved portion350, the distal-tip portion undergoes the above-described curvature (i.e., the curvature that is similar to that of a question mark or a tennis-racket) in a second plane. In other words, distal curved portion349comprises a straight segment355(analogous to straight portion346ofFIG.17A) that defines a longitudinal axis, and a curved segment that is distal to straight segment355and curves in a first direction with respect to the longitudinal axis before passing through an inflection point and curving in a second direction with respect to the longitudinal axis, such that the curved segment defines bulge351on one side of the longitudinal axis. In other embodiments, distal curved portion349has a pigtail shape. For some applications, the proximal curved portion and the distal curved portion are continuous with each other. Typically, for such applications, the distal-tip portion does not include straight segment355between the proximal curved portion and the distal curved portion.

Typically, lumen122(FIG.15A) passes through proximal curved portion350and distal curved portion349. Also typically, valve160(FIG.15A) is at the distal end of distal curved portion349.

In some embodiments, the impeller of the blood pump is rotated (i.e., the blood pump is operated) while distal curved portion349pushes against apex342. For some applications, the proximal curved portion is configured such that if the tip is pushed against apex342, then even in the event that distal curved portion349curls up, the distal blood-inlet openings108are still pushed in the direction of free wall334and away from the septal wall338, by the tip flexing about proximal curved portion350.

Reference is now made toFIG.18, which is a schematic illustration of a steering mechanism that is used to steer frame34with respect to a portion of the device disposed proximally thereto, in accordance with some applications of the present invention.

In some embodiments, one or more steering wires301are coupled to frame34and are configured to extend from the frame, to outside the body of the subject, while the frame is within the body. In some embodiments, steering wires301pass through delivery tube142. For example, the steering wires may be disposed between drive-cable-bearing tube140and the delivery tube.

In some embodiments, one of the steering wires is coupled to the proximal end of frame34. Alternatively or additionally, one of the steering wires may be coupled to the distal end of frame34. This coupling may be via proximal bearing housing116H or distal bearing housing118H, i.e., the steering wire may be coupled to the frame by virtue of being coupled to the proximal or distal bearing housing.

Typically, for such applications, handle149of the ventricular assist device comprises a steering control300, and the steering wires are coupled to steering control300so as to be controllable via the steering control.

Advantageously, steering wires301may be configured for orienting the frame, e.g., prior to rotating the impeller so as to pump blood of the subject. For example, the steering wires may be configured to extend from the frame while the frame is within the left ventricle of the subject's heart, and the steering wires may be configured for orienting the frame within the left ventricle, e.g., such that the frame extends toward the apex of the left ventricle. The frame may be oriented by applying a force to a steering wire coupled to the proximal end of the frame while applying a counterforce to the delivery tube or to another steering wire coupled to the distal end of the frame.

For some such applications, a controller steers the proximal end of the frame with respect to a portion of the device disposed proximally thereto, for example, in order to direct the distal end of the frame to extend in the direction of the apex of the left ventricle. Typically, upon releasing the pump head within the left ventricle, the operator steers the proximal end of the frame in the above-described manner, in order to conform with the anatomy (e.g., the shape of the left ventricle) of the subject.

As described hereinabove with reference toFIG.9B, for some applications, the ventricular assist device includes optical fibers228for blood-pressure measurement, which extend from a proximal end of the device to, for example, the proximal end of the frame. For some applications, optical fibers228are additionally used as steering wires, in the above-described manner. (In other words, the steering wires comprise respective optical fibers.) In other embodiments, one or more of the steering wires comprises an elongation-resistant fiber, such as an aramid fiber.

Reference is now made toFIGS.19A-B, which are schematic illustrations of an expandable element surrounding delivery tube142, in accordance with some applications of the present invention.

In some embodiments, an expandable element314, such as an expandable stent, an expandable braided element, or an inflatable element316(e.g., a balloon114), surrounds delivery tube142proximally to blood-outlet openings109, with the length of the delivery tube between expandable element314and the blood-outlet openings being less than 30 mm. For example, expandable element314may be disposed in the vicinity of the interface between delivery tube142and the region at which the proximal end of pump-outlet tube24is coupled to the delivery tube (as shown inFIG.19A).

In some embodiments, as shown inFIG.19A, expandable element314is entirely proximal to the pump-outlet tube. In other embodiments, as shown inFIG.19B, expandable element314is an inflatable element316(e.g., balloon114), which when inflated, is disposed at least partly within, e.g., entirely within, pump-outlet tube24.

Typically, expandable element314, whether configured as shown inFIG.19A,FIG.19B(orFIG.20AorFIG.20B) is configured to center a portion of the ventricular assist device (e.g., a portion of delivery tube142and, in particular, the portion of the delivery tube near pump-outlet tube24) within the aorta, by contacting the aorta wall or, if the expandable element is within the pump-outlet tube, by pushing the wall of the pump-outlet tube against the aorta wall.

In some embodiments, expandable element314is an inflatable element316that is shaped to direct the blood through blood-outlet openings109, as indicated inFIG.19Bby blood-flow arrows318. For example, the distal end of the inflatable element may have a width that decreases moving distally, e.g., the distal end may be frustoconical, such that the blood is directed by the distal end of the inflatable element, at an angle, through the blood-outlet openings.

For some applications, the inflatable element (e.g., balloon) is inflated using a fluid (e.g., a purging fluid), which is pumped through the ventricular-assist device. For example, as shown inFIG.19B, the wall of the delivery tube may be shaped to define one or more openings320, and inflatable element316may surround openings320such that a fluid flowing, via the openings, from the delivery tube into the inflatable element inflates the inflatable element. This fluid may include purging fluid, which, distally to openings320, purges the interface between the axial shaft and the radial bearings, which don't rotate with the axial shaft.

It is noted that expandable element314may be combined with any of the embodiments of pump-outlet tube24described below with reference toFIGS.33A-C.

As shown inFIG.19A, for some applications, regardless of whether the ventricular assist device comprises expandable element314, proximally to the proximal conical portion of the pump-outlet tube, the pump-outlet tube defines a tubular coupling portion45, via which the proximal end of the pump-outlet tube is coupled (e.g., via an adhesive) to second outer tube142of the ventricular assist device. For some applications, the pump-outlet tube is manufactured using a single continuous tube, with respective portions of the pump-outlet tube being molded to define the tubular coupling portion, the proximal conical portion, the distal conical portion, and the cylindrical central portion. Typically, in such cases, the blood-inlet openings and the blood-outlet openings are cut (e.g., laser cut) from the tube. For some applications, rather than molding the tube to define the proximal conical section and the tubular coupling portion, initially the portion of the tube that will form the proximal conical section and the tubular coupling portion are shaped as a cylinder (which is typically continuous with the cylinder shape of the central portion). At its proximal end, strips are then cut (e.g., laser cut) from the tube, leaving other strips still attached to, and extending proximally from, the central cylindrical portion of the tube. The proximal ends of the strips are then adhered to outer tube142of the ventricular assist device, in such a manner that they define a conical portion of the pump-outlet tube that defines blood-outlet openings. (For some applications, some of the blood-outlet openings are laser cut from the tube, and other blood-outlet openings are formed by adhering the above-described strips to second outer tube142of the ventricular assist device.) For some applications, by forming the conical portion of the pump-outlet tube and the blood-outlet openings using the latter method, the thickness of the layer of the pump-outlet tube that is coupled to outer tube142is less than the thickness of the tubular coupling portion as formed by the former method. For some applications, this reduces the sharpness of the diameter change at the interface between the outer tube142and the region at which the proximal end of the pump-outlet tube is coupled to the outer tube.

For some applications, the pump-outlet tube is manufactured using a single continuous tube, with respective portions of the pump-outlet tube being molded to define the tubular coupling portion, the proximal conical portion, the distal conical portion, and the cylindrical central portion, as in the former method described in the above paragraph. However, before adhering the tubular coupling portion to outer tube142of the ventricular assist device, the tubular coupling is cut (e.g., in a tapered manner), so as to reduce the thickness of the layer of the pump-outlet tube that is coupled to outer tube142and/or to prevent folds forming in the tubular coupling portion of the pump-outlet tube.

Typically, (a) blood-outlet openings109are defined by portions of the wall of the blood outlet tube that at least partially extends into the proximal conical portion of the pump-outlet tube, e.g., as shown inFIG.19A, and/or (b) blood-outlet openings109are laterally facing, by virtue of being defined by the central cylindrical portion of pump-outlet tube24, e.g., as shown inFIG.19B. (As noted hereinabove, “laterally-facing blood-outlet openings” should be interpreted to mean that the blood-outlet openings are disposed laterally with respect to the longitudinal axis of the pump-outlet tube, by virtue of being defined by the central cylindrical portion of the pump-outlet tube. This is in contrast to blood-inlet openings that are described as “lateral blood-inlet openings,” which are typically not oriented entirely laterally with respect to the longitudinal axis of the pump-outlet tube. Rather, they are obliquely disposed with respect to the longitudinal axis of the pump-outlet tube.) Examples of each type of blood-outlet openings are shown in respective figures in the present application. The scope of the present disclosure includes combining other features of the pump-outlet tube and/or other portions of the ventricular assist device with any configuration of blood-outlet openings that are described and/or shown in the present application.

Reference is now made toFIGS.20A,20B,20C, and20D, which are schematic illustrations of ventricular assist device20, in accordance with some applications of the present invention.

FIG.20Ais similar toFIG.19B, in thatFIG.20Ashows expandable element314disposed within pump-outlet tube24, typically in the vicinity of the blood-outlet openings. For some such applications, the blood-outlet openings are configured as described hereinabove. In accordance with some applications, the expandable element is an inflatable element316, e.g., balloon114, as shown inFIG.20A. For some applications, the balloon is inflated using purging fluid, which is pumped through the ventricular-assist device. Alternatively or additionally, the balloon is inflated with a different fluid, e.g., a dedicated supply of fluid (such as air or saline) for inflating the balloon.

For some applications, the expandable element is configured to act as a blood flow director, by directing blood from the proximal end of the pump-outlet tube through the blood outlet openings, as indicated by blood-flow arrows360. For some applications, the expandable element is shaped such as to direct the blood flow in this manner. For example, the expandable element may have an angled and/or a curved surface that is configured to direct the blood flow in this manner. For some applications, by directing blood flow in this manner, the overall pumping efficiency of the device is increased relative to if the device does not include an expandable element.

As described hereinabove, typically, expandable element314is configured to center a portion of the ventricular assist device (e.g., a portion of delivery tube142and, in particular, the portion of the delivery tube near pump-outlet tube24) within the aorta, by contacting the aorta wall or, if the expandable element is within the pump-outlet tube, by pushing the wall of the pump-outlet tube against the aorta wall.

Referring toFIG.20B, for some applications, expandable element314is a porous expandable element, such as an expandable cage or stent172, surrounds delivery tube142within, or at the proximal end of, the pump-outlet tube, such that the blood is pumped through the porous expandable element. For example, the porous expandable element may be disposed at the proximal end of the pump-outlet tube, and the proximal end of the pump outlet tube may be coupled to delivery tube142via the porous expandable element. For some such applications, the pump-outlet tube does not define blood-outlet openings109(FIG.20A). Rather, blood flows out of the pump-outlet tube exclusively via the porous expandable element, as indicated by blood-flow arrows174. For some applications, the porous expandable element comprises a structure made of a shape-memory alloy, such as a laser-cut shape-memory alloy and/or a braided shape-memory alloy.

As described hereinabove, typically, the porous expandable element is configured to center a portion of the ventricular assist device (e.g., a portion of delivery tube142and, in particular, the portion of the delivery tube near pump-outlet tube24) within the aorta, by contacting the aorta wall or, if the expandable element is within the pump-outlet tube, pushing the wall of the pump-outlet tube against the aorta wall.

Referring toFIG.20C, for some applications, a proximal portion178of pump-outlet tube24, which defines blood-outlet openings109, is folded inwardly toward the distal end of the pump-outlet tube. Typically, as shown, proximal portion178is folded inwardly such that the blood-outlet openings direct blood proximally (substantially parallel to the axis of outer tube142) rather than radially outwardly (away from the axis of outer tube142), as indicated by blood-flow arrows362. For some applications, proximal portion178is folded inwardly such that the blood-outlet tube forms a protective layer between blood flowing out of the blood-outlet openings and the walls of the subject's aorta.

Referring toFIG.20D, for some applications, the blood-outlet openings109are defined by a substantially proximally-facing surface190of the pump-outlet tube, rather than being defined by a lateral surface of the pump-outlet tube. Typically, for such applications, blood flow from the pump-outlet tube is axially directed, as indicated by blood-flow arrows364.

Reference is now made toFIG.20E, which is a schematic illustration of pump-outlet tube24that defines a blood-flow chamber366at its proximal end, in accordance with some applications of the present invention. For some applications, the blood-flow chamber is defined by an internal membrane368that is disposed within the proximal end of the pump-outlet tube and that defines holes370therethrough. Blood flows into the blood-flow chamber via the holes, as indicated by blood-flow arrows372. Subsequently, the blood flows out of the blood-flow-chamber and into the subject's aorta via blood-outlet openings109(which are generally as described hereinabove), as indicated by blood-flow arrows374. Typically, by virtue of the blood flowing through the blood-flow chamber, the blood-flow chamber inflates such as to center a portion of the left-ventricular assist device (e.g., the delivery tube, and in particular, the portion of the delivery tube near pump-outlet tube24) within the aorta, by contacting the aorta wall. Typically, the internal membrane is shaped so as to direct the blood flow out of the blood-outlet openings.

For some applications, internal membrane368is a continuation of pump-outlet tube24, and the internal membrane is covered with an external membrane, which defines blood-outlet openings, and which forms the external surface of blood-flow chamber366. For such applications, the proximal end of the blood-outlet tube is shaped so as to direct the blood flow out of the blood-outlet openings.

Typically, the combination of the proximal end of the blood-outlet tube and an additional membrane (whether an internal membrane or an external membrane) are configured to define blood-flow chamber366, which typically functions as described above. In general, the scope of the present disclosure includes any structure that provides a blood-flow chamber disposed at a proximal end of the pump-outlet tube, the blood-flow chamber defining (a) holes370via which blood is pumped into the blood-flow chamber and (b) blood-outlet openings109configured to be disposed with within the aorta via which the blood flows out of the blood-flow chamber and into the aorta.

Reference is now made toFIGS.21A and21B, which are schematic illustrations of distal bearing housing118H of ventricular assist device20, in accordance with some applications of the present invention. For some applications, the bearing housing is constructed using a cylindrical tube180that has substantially smooth outer and inner surfaces and a uniform thickness (i.e., without the various regions on the surfaces and varying thicknesses described with reference toFIG.11E). Typically, the cylindrical tube is made from an alloy or a metal, such as cobalt chrome and/or stainless steel. For some applications, a coupling element182is coupled to the outer surface of the cylindrical tube and the distal end of frame34is coupled to the bearing housing via the coupling element. (For example, the distal end of the frame may become coupled to the coupling element via a snap-fit mechanism.) For some applications, distal bearing118itself comprises an inner cylindrical element that is coupled to the inner surface of the distal bearing housing. As described hereinabove, typically, the bearing is made of a ceramic material such as zirconia. For some applications, distal thrust bearing270is disposed inside distal bearing housing118H, as shown. The distal thrust bearing is generally as described hereinabove. For some applications, the proximal bearing housing also comprises a cylindrical tube that has substantially smooth outer and inner surfaces and a uniform thickness, as described with respect to the distal bearing housing mutatis mutandis.

Reference is now made toFIGS.22A,22B,22C, and22D, which are schematic illustrations of portions of ventricular assist device20, the device including an inlet guard400disposed inside frame34, in accordance with some applications of the present invention. Inlet guard400is shaped to define one or more holes402, shown enlarged inFIG.22D, which are disposed around the axial shaft and within frame34distally to the impeller, such that the blood flows to the impeller via holes402. The inlet guard may be coupled to the struts of frame34, to inner lining39(FIG.4) of the frame, to the inner wall of pump-outlet tube24, and/or to distal bearing housing118H.

For some applications, inlet guard400is flat and/or is disposed such that it is perpendicular to the axial shaft (i.e., to the longitudinal axis of the frame). Thus, advantageously, the inlet guard may occupy relatively little space, and/or may provide an advantageous flow direction for the blood. Typically, the inlet guard is toric.

As described hereinabove (with reference toFIGS.16A-E), for some applications, the ventricular assist device includes a thrust bearing in pump-head portion27. Typically, for such applications, the impeller does not move distally of cylindrical portion38of frame34(either during delivery of the device to the left ventricle or during operation of the device).

For some applications, the inlet guard is placed within the frame at the distal end of the central cylindrical portion of the frame or in the vicinity thereof, e.g., within 1 mm of the distal end of the cylindrical portion. This placement may simplify the assembly of the blood pump.

Typically, the inlet guard is polymeric, i.e., is made of a polymeric material (such as polyurethane (e.g., Pellethane®), polyethylene terephthalate (“PET”), ultra-high-molecular-weight polyethylene (“UHMWPE”), and/or polyether block amide (e.g., Pebax®)) that is shaped to define holes402. For some applications, the thickness of the inlet guard is more than 40 microns (e.g., more than 50 microns), and/or less than 100 microns (e.g., less than 80 microns), for example, 40-100 microns or 50-80 microns. Thus, the inlet guard may be configured to withstand pressure yet be crimpable.

Typically, for applications in which ventricular assist device20includes inlet guard400disposed inside frame34, pump-outlet tube24does not extend until the distal end of distal conical portion40of frame34. Moreover, pump-outlet tube24may have an open distal end, rather than terminating in a distal conical portion. (Thus, the inlet guard may simplify the manufacture of the blood pump.) The distal end of the pump-outlet tube may be proximal to the distal end of the distal conical portion of the frame. For example, the distal end of the pump-outlet tube may be within 1 mm of the distal end of the central cylindrical portion of frame34, i.e., the pump-outlet tube may extend only until the end of the cylindrical portion of frame34, or the vicinity thereof. Blood may thus flow into frame34via openings defined by the distal conical portion of the frame.

For some applications, holes402of inlet guard400are sized such as (a) to allow blood to flow from the subject's left ventricle into pump-outlet tube24and (b) to block structures from the subject's left ventricle from entering into the pump-outlet tube. Typically, for such applications, the inlet guard is configured to reduce a risk of structures from the left ventricle (such as chordae tendineae, trabeculae carneae, and/or papillary muscles) entering into pump-outlet tube24and potentially being damaged by the impeller and/or the axial shaft, and/or causing damage to the ventricular assist device.

For some applications, inlet guard400defines more than 10 holes, more than 50 holes, more than 100 holes, or more than 150 holes, e.g., 50-100 holes, 100-150 holes, or 150-200 holes. For some applications, the holes are sized such as (a) to allow blood to flow from the subject's left ventricle into the tube and (b) to block structures from the subject's left ventricle from entering into the frame. Typically, for such applications, the inlet guard is configured to reduce a risk of structures from the left ventricle (such as chordae tendineae, trabeculae carneae, and/or papillary muscles) entering into the cylindrical portion of frame34and potentially being damaged by the impeller and/or the axial shaft, and/or causing damage to the left ventricular assist device. Therefore, for some applications, the holes are shaped such that, for each of the holes, the span of the hole in at least one direction is less than 1 mm, e.g., 0.1-1 mm, or 0.3-0.8 mm. By defining such a small width (or span), it is typically the case that structures from the left ventricle (such as chordae tendineae, trabeculae carneae, and/or papillary muscles) are blocked from entering into the cylindrical portion of frame34.

For some applications, each of the holes defines an area of more than 0.05 square mm (e.g., more than 0.1 or 0.3 square mm), and/or less than 5 square mm (e.g., less than 3 or 1 square mm), e.g., 0.05-5, 0.05-3, 0.1-1, 0.1-5, or 0.3-1 square mm.

Typically, the inlet guard has a porosity of at least 40 percent, e.g., more than 50 percent, or more than 60 percent (where porosity is defined as the percentage of the area of this portion that is porous to blood flow). Thus, on the one hand, the holes are relatively small (in order to prevent structures of the left ventricular from entering the frame), but on the other hand, the porosity of the portion of the pump-outlet tube that defines the holes is relatively high, such as to allow sufficient blood flow into the pump-outlet tube.

For some applications, each of the holes has a circular or a polygonal shape. For some applications, each of the holes has a hexagonal shape, as shown most clearly inFIG.22D. Typically, using openings having a hexagonal shape allows the inlet guard to have a relatively high porosity (e.g., as described hereinabove), while providing the inlet guard with sufficient material between the holes to prevent tearing and/or stretching of the material.

As shown inFIG.22D, for some applications, a width W2 of gaps between adjacent holes402(i.e., the distance between each pair of adjacent holes) is more than 0.01 mm (e.g., more than 0.04 mm), and/or less than 0.1 mm (e.g., less than 0.08 mm), for example, 0.01-0.1 mm, or 0.04-0.08 mm.

As further shown inFIG.22D, for some applications, the distance D2 between opposing sides of each of the hexagons (or other types of polygons) is more than 0.2 mm (e.g., more than 0.4 mm) and/or less than 0.8 mm (e.g., less than 0.6 mm), e.g., 0.2-0.8 mm, or 0.4-0.6 mm. Typically each of the polygons encloses a circle (such that any structure that cannot pass through such a circle would be unable to pass through the polygon). Typically, the diameter of the circle enclosed by the polygon is the equivalent of distance D2, e.g., more than 0.2 mm (e.g., more than 0.4 mm) and/or less than 0.8 mm (e.g., less than 0.6 mm), e.g., 0.2-0.8 mm, or 0.4-0.6 mm.

For some applications, the frame is assembled with the inlet guard inside in the following manner. As described hereinabove, during assembly of the pump-head portion, the proximal end of frame34is typically open. For some applications, the inlet guard is placed through the open proximal end of the frame while being supported upon a rod (e.g., a mandrel). The inlet guard typically has an overall torus shape, with the edges of the shape defining inner and outer circles, as shown inFIG.22D. The inner circle defined by the inlet guard is typically coupled to the axial shaft (or to the distal bearing housing) and the outer circle is coupled to struts of frame34, to pump-outlet tube24and/or to inner lining39. For some applications, the aforementioned coupling of the inlet guard to other portions of the device is performed via suturing, via hooks, via adhesive, and/or via heat fusion.

As noted above, in some embodiments, the inlet guard is coupled to the distal bearing housing, which may house a radial and/or thrust bearing. In such embodiments, typically, the distal bearing housing is partly or entirely disposed within frame34. For example, at least 10%, 50%, or 80% of the length of the bearing housing may be disposed within the frame. Moreover, the distal bearing housing may extend into the frame even for applications in which the blood pump does not comprise inlet guard400.

For further details in this regard, reference is now made toFIG.23, which is a schematic illustration of ventricular assist device20in which distal bearing housing118H is disposed at least partly within, i.e., extends proximally at least partially into distal conical portion40of frame34. Reference is also made again toFIGS.22A-C, which similarly show distal bearing housing118H extending proximally at least partially into the distal conical portion of frame34.

Typically, it is desirable for the relatively rigid portions of the pump-head portion of the device (such as the frame, the impeller, and the bearing housings) to have a combined length that is as short as possible, in order to safely navigate the device through curved vasculature of the subject (such as the aortic arch). For some applications, by configuring the device such that distal bearing housing118H extends at least partially into the distal conical portion of frame34, the combined length of the relatively rigid portions of the pump-head portion of the device (such as the frame, the impeller, and the bearing housings) is shortened relative to if the device were configured such distal bearing housing118H would not extend proximally at least partially into the distal conical portion of frame34.

In some embodiments, the bearing housing occupies at least 10% of the length of the distal conical portion of the frame. For some applications, the distal bearing housing extends until, or even at least partially into, cylindrical portion38of frame34.

As noted above, distal bearing housing118H houses a radial and/or thrust bearing, which is adjacent to the axial shaft and is configured to stabilize the axial shaft (radially and/or axially) while the axial shaft rotates. An advantage of a thrust bearing is that the impeller does not advance into the distal conical portion of the frame (or into a portion thereof) either during operation of the device or during delivery of the device to the left ventricle (when the impeller is in a radially constrained configuration within the frame), such that it may be easier to configure distal bearing housing118H to extend proximally into the distal conical portion of frame34. However, the scope of the present disclosure includes a distal bearing housing that extends into frame34in the manner described above, but which houses only distal radial bearing118and does not house a thrust bearing.

Likewise, the proximal bearing housing, which houses a radial and/or thrust bearing adjacent to the axial shaft and configured to stabilize the axial shaft (radially and/or axially) while the axial shaft rotates, may be disposed partly or entirely within the proximal conical portion of the frame. For example, at least 10%, 50%, or 80% of the length of the proximal bearing housing may be disposed within the frame. Alternatively or additionally, the proximal bearing housing may occupy at least 10% of the length of the proximal conical portion of the frame; for example, the proximal bearing housing may be disposed at least partly within the cylindrical portion of the frame. The scope of the present disclosure includes a proximal bearing housing that extends into frame34in the manner described above, but which houses only proximal radial bearing116and does not house a thrust bearing.

In some embodiments, both the proximal and distal bearing extend into the frame as described above. In such embodiments, the distance between the proximal end of the proximal bearing housing and the distal end of the distal bearing housing may be less than 10% greater than the length of the frame, thus reducing the total length of the more rigid elements of the blood pump.

Reference is now made toFIG.24, which is a schematic illustration of valve160disposed at the distal end of lumen122of distal-tip portion120, in accordance with some applications of the present invention. As described hereinabove, in some embodiments, the valve is a duckbill valve. For some applications, the duckbill valve is configured to facilitate insertion of a guidewire both via the distal end of the duckbill valve and via the proximal end of the duckbill valve. For some applications, at both the proximal and the distal end, the valve narrows from a wide opening to a self-sealing central portion, such as to guide the guidewire toward the self-sealing central portion from either end of the valve. Typically, a guidewire is inserted into the distal-tip portion from the distal end of the distal-tip portion in advance of the initial delivery of the ventricular assist device to the left ventricle over the guidewire. In some cases, it is desirable to insert a guidewire from the proximal end of the ventricular assist device to the distal end of the device at some stage thereafter. The shape of the duckbill valve that is shown inFIG.24typically facilitates both the distal-to-proximal insertion and the proximal-to-distal insertion of the guidewire.

Reference is now made toFIGS.25A-25B, which are schematic illustrations of valve99disposed at the proximal end of lumen133defined by pin131of driven-magnet unit310, in accordance with some applications of the present invention. As described hereinabove, in some embodiments, the valve is a duckbill valve. For some applications, the duckbill valve is configured to facilitate insertion of a guidewire via both the distal end of the duckbill valve and via the proximal end of the duckbill valve. For some applications, at both the proximal and the distal end the valve narrows from a wide opening to a self-scaling central portion, such as to guide the guidewire toward the self-sealing central portion from either end of the valve. Typically, a guidewire is inserted through valve99from the distal end of the driven-magnet unit in advance of the initial delivery of the ventricular assist device to the left ventricle over the guidewire. In some cases, it is desirable to insert a guidewire from the proximal end of the ventricular assist device to the distal end of the device at some stage thereafter. The shape of the duckbill valve that is shown inFIGS.25A-Btypically facilitates both the distal-to-proximal insertion and the proximal-to-distal insertion of the guidewire.

Reference is now made toFIGS.26A and26B, which are schematic illustrations of, respectively, an oblique view and a cross-sectional view of a locking unit279for securing delivery tube142to delivery catheter143, in accordance with some applications of the present invention.

For some applications, once the ventricular assist device has been deployed within the left ventricle, and the delivery catheter has been retracted to its intraprocedural position (e.g., such that its distal end is disposed within the descending aorta), it is desirable to secure delivery tube142and delivery catheter143in fixed positions with respect to each other. For some applications, a fixation unit such as fixation unit97(shown inFIG.7D) is used, with the fixation unit being configured such that the position of the driven-magnet unit is fixed or unfixed relative to the proximal end of the delivery catheter by screwing a portion of the fixation unit. For example, the fixation unit may include a Tuohy Borst adapter.

For other applications, locking unit279, which is configured to couple to the proximal end of the delivery catheter and comprises a clip280, is used instead of a fixation unit such as that shown inFIG.7D. For such applications, delivery tube142extends from outside the body of the subject into the left ventricle through locking unit279and delivery catheter143. Clip280is configured such that it has only two states (or “positions”): a closed state, in which the clip grips the delivery tube so as to inhibit movement of the delivery tube relative to the delivery catheter, and an open state, in which the clip does not grip the delivery tube. This differs from a fixation unit such as that shown inFIG.7D, which typically has varying degrees of closedness depending on how tightly the screw mechanism is activated. Typically, clip280is an external component of the locking unit, such that the state of clip280is readily visible to the user.

As shown inFIG.26B, the locking unit typically comprises at least one internal seal configured to surround the delivery tube and to inhibit backflow of blood of the subject, from the delivery catheter, through the locking unit, while the delivery tube passes through the locking unit. For example, the locking unit may comprise a duckbill seal282and an additional seal284.

Typically, as shown inFIG.26B, the locking unit further comprises a fluid port281and is shaped to define a channel283in fluid communication with fluid port281and delivery catheter143. Thus, channel224(FIG.12A) between the delivery catheter and the delivery tube may be purged with a purging fluid. (Fluid port281and channel283may be collectively referred to as purging fluid inlet port89, as indicated inFIG.26Aand as described above with reference toFIG.7D.)

Reference is now made toFIGS.27A and27B, which are schematic illustrations of, respectively, an oblique view and a cross-sectional view of a clip286for securing delivery tube142to delivery catheter143, in accordance with some alternative applications of the present invention.

In some embodiments, the clip is configured to remain in its open state, following a placement of the clip in its open state, unless the clip is returned to its closed state. In other words, the fixation mechanism can be disengaged by opening the clip, and the user does not then need to actively maintain disengagement of the fixation mechanism.

In other embodiments, the clip is configured to return to its closed state, following a placement of the clip in its open state, unless the clip is held in its open state. In other words, the user needs to actively apply force to the clip in order to maintain disengagement of the fixation mechanism. Clip286is an example of such a clip.

Reference is now made toFIGS.28A and28B, which are schematic illustrations of, respectively, an oblique view and a cross-sectional view of another locking unit287for securing delivery catheter143to an introducer sheath290, in accordance with some applications of the present invention.

Typically, delivery catheter143is inserted from outside the subject's body into the patient's vasculature via introducer sheath290, which is configured to extend from outside the body of the subject into the body of the subject. For example, the introducer sheath may be used to provide access to the subject's femoral artery, and the delivery catheter may be advanced from the femoral artery to the subject's left ventricle as described hereinabove.

For some applications, locking unit287is configured to couple to the proximal end of introducer sheath290, and delivery catheter143is configured to pass through locking unit287and introducer sheath290. Locking unit287comprises a clip288, which is used to fix the position of the delivery catheter with respect to the introducer sheath (typically, once the delivery catheter has been positioned in a desired intraprocedural position). Clip288is configured such that it has only two states (or “positions”): a closed state, in which the clip grips the delivery catheter so as to inhibit movement of the delivery catheter relative to the introducer sheath, and an open state, in which the clip does not grip the delivery catheter. This differs from a fixation unit such as that shown inFIG.7D, which typically has varying degrees of closedness depending on how tightly the screw mechanism is activated. Typically, clip288is an external component of locking unit287, such that the state of the clip is readily visible to the user.

Typically, locking unit287comprises at least one internal seal289configured to surround the delivery catheter and to inhibit backflow of blood of the subject, from the introducer sheath, through the second locking unit, while the delivery catheter passes through the second locking unit.

Typically, locking unit287further comprises a fluid port, and is shaped to define a channel291in fluid communication with the fluid port and introducer sheath290. Thus, the space between the delivery catheter and the introducer sheath may be purged with a purging fluid.

In some embodiments, clip288is configured to remain in its open state, following a placement of the second clip in its open state, unless the clip is returned to its closed state. In other embodiments, the clip is configured to return to its closed state, following a placement of the clip in its open state, unless the clip is held in its open state.

Reference is now made toFIG.29, which is a schematic illustration of a left-ventricular assist device, in accordance with some applications of the present invention.

In some embodiments, the proximal end of pump-outlet tube24is folded inwardly so as to define one or more surfaces322configured to direct the blood through blood-outlet openings109by virtue of being oblique with respect to the longitudinal axis324of the pump-outlet tube.

In some embodiments, surfaces322define a projection325, such as a frustoconical projection, having a width that decreases moving distally. Blood is directed by projection325, at an angle, through the blood-outlet openings. In such embodiments, typically, projection325is distally coupled to the delivery tube.

Reference is now made toFIG.30, which is a schematic illustration of a ventricular assist device deployed within the body of a subject, in accordance with some applications of the present invention.

In some embodiments, an elongation-resistant fiber, e.g., an aramid fiber292(which may be equivalent to fiber159shown inFIG.13) runs axially along the wall of delivery tube142. The elongation-resistant properties of aramid fiber292cause the delivery tube to adopt an orientation in which the length of the aramid fiber is minimized. Thus, aramid fiber292biases the orientation of the delivery tube, typically at least by biasing the roll angle of the delivery tube, while the delivery tube traverses the aortic arch, thereby biasing the orientation of the blood pump. Typically, aramid fiber292biases the roll angle of the delivery tube such that the aramid fiber is disposed at the inside of the curve of the aortic arch, such that the length of the aramid fiber is minimized. Typically, aramid fiber292extends to the distal end of the delivery tube.

In some embodiments, as shown in the distal-end view294of the delivery tube, aramid fiber292is disposed within the wall of the delivery tube (e.g., as inFIG.13). For example, the wall of the delivery tube may comprise a braid, and the aramid fiber may be threaded through the braid or run adjacently to the braid.

As described above, pump-outlet tube24is coupled to delivery tube142. Thus, in addition to biasing the orientation of the blood pump, which is disposed at least partly within the pump-outlet tube, aramid fiber292biases the orientation of the pump-outlet tube.

Typically, as shown inFIG.30, the pump-outlet tube is configured to curve within the left ventricle while the blood pump pumps blood from the left ventricle through the pump-outlet tube. For example, the blood pump may be disposed at least partly within the distal portion of the pump-outlet tube, and the proximal portion of the pump-outlet tube may curve.

For example, the delivery tube may be configured to curve (even without aramid fiber292), and the pump-outlet tube may curve by virtue of the delivery tube curving. Alternatively or additionally, the pump-outlet tube may be configured to curve by virtue of being pre-shaped. Alternatively or additionally, the pump-outlet tube may be shaped to define multiple openings (e.g., blood-inlet openings and/or blood-outlet openings) arranged in a non-axisymmetric arrangement, and the pump-outlet tube may curve by virtue of blood flowing through the openings. For some applications, the pump-outlet tube is configured to curve by virtue of one or more features described with reference to FIGS. 19A-F of US 2022/0226632 to Tuval, which is incorporated herein by reference. Alternatively or additionally, as described below with reference toFIG.31, the pump-outlet tube may curve by virtue of one or more bands being bonded to the outer wall of the pump-outlet tube.

Due to the curvature of the pump-outlet tube, the orientation-biasing properties of the aramid may be particularly helpful. For example, the aramid fiber may bias the orientation of the delivery tube such that, while the delivery tube traverses the aortic arch (and the aramid fiber biases the roll angle of the delivery tube such that the aramid fiber is disposed at the inside of the curve of the aortic arch), the pump-outlet tube curves away from the posterior wall and/or septal wall338of the left ventricle. Alternatively or additionally, the aramid fiber may bias the orientation of the delivery tube such that, while the delivery tube traverses the aortic arch (and the aramid fiber biases the roll angle of the delivery tube such that the aramid fiber is disposed at the inside of the curve of the aortic arch), the pump-outlet tube curves toward apex342and/or free wall334of the left ventricle.

In some embodiments, the circumferential angle between the aramid fiber and the circumferential position on the pump-outlet tube at the inside of the curve of the pump-outlet tube is close to zero, e.g., between −10 and 10 degrees. In other words, the aramid fiber is positioned at or near the inside of the curve, such that the delivery tube curves in the same direction as the pump-outlet tube. This feature may further help orient the blood pump in the desired orientation.

As described hereinabove, for some applications, distal-tip portion is configured to curve within a plane. For example, within a given plane, the distal tip portion may have a curvature such as that described hereinabove with reference toFIGS.17A-C. For some applications, the plane in which distal-tip portion120curves is disposed at a roll angle theta with respect to a (hypothetical) line extending from a distal end292dof the aramid fiber (which extends out of the page in distal-end view294). For some applications, the angle theta is between 25 and 65 degrees, such as between 35 and 55 degrees. Typically, the angle theta is selected such that while the delivery tube traverses the aortic arch (and the aramid fiber biases the roll angle of the delivery tube such that the aramid fiber is disposed at the inside of the curve of the aortic arch), the plane in which the distal-tip portion curves is disposed at an orientation with respect to the apex of the left ventricle that is such that the distal-tip portion provides the functionalities of the distal-tip portion that are described hereinabove with reference toFIGS.17A-C.

FIG.30shows an embodiment in which the aramid fiber is parallel to the longitudinal axis296of the delivery tube. In other embodiments, at least a portion of the aramid fiber is not parallel to longitudinal axis296. Thus, in addition to biasing the roll angle of the delivery tube, the aramid fiber may twist the delivery tube. For some applications, this portion of the aramid fiber is configured to twist a portion of the delivery tube that passes through the ascending aorta in the vicinity of the aortic valve, such as to conform with the anatomical twist of this portion of the aorta. For some applications, the length of this non-parallel portion is between 5 and 20 cm (e.g., between 8 and 12 cm), and/or the circumferential span of the portion is between 30 and 60 (e.g., between 35 and 55) degrees.

It is noted that aramid fiber292may be used to bias the orientation of delivery tube142regardless of the type of device coupled to the distal end of the delivery tube. In fact, the aramid fiber may be used even if no device at all is coupled to the distal end of the delivery tube. Moreover, the delivery tube may extend into any chamber of the heart. Furthermore, it is noted that although some embodiments have been described with reference to an aramid fiber, the scope of the present disclosure includes using other types of elongation-resistant fibers (e.g., polymeric and/or natural fibers) and/or elongation-resistant wires to provide the functionalities described hereinabove both with reference to fiber159(shown inFIG.13) and with reference to fiber292(shown inFIG.30).

Reference is now made toFIG.31, which is a schematic illustration of a left-ventricular assist device comprising a curved pump-outlet tube24, in accordance with some applications of the present invention.

As noted above, frame34is disposed at least partly within the distal portion of pump-outlet tube24and is configured to hold the distal portion of the pump-outlet tube open. The impeller is disposed within frame34and is configured to pump blood of the subject, through the pump-outlet tube, from the left ventricle into the aorta, thereby maintaining the proximal portion of the pump-outlet tube, which is proximal to the frame and traverses the aortic valve, in an open state.

In some embodiments, the left-ventricular assist device comprises one or more (e.g., 1-8) bands298, each of which is bonded to the outer wall of the proximal portion of the pump-outlet tube, without extending around the full circumference of the pump-outlet tube. For example, each of the bands may extend around 20-80% of the full circumference of the pump-outlet tube. Thus, while the proximal portion of the pump-outlet tube remains open by virtue of the blood flowing proximally through the tube, the proximal portion of the pump-outlet tube curves at respective locations of bands298.

Advantageously, the curvature of the proximal portion of the pump-outlet tube helps orient the distal portion of the pump-outlet tube. For example, by virtue of the proximal portion of the pump-outlet tube curving at the locations of the bands, the distal portion of the pump-outlet tube may point toward apex342or toward free wall334(FIG.30).

In some embodiments, as described above with reference toFIGS.9A-B, bands298couple blood-pressure-measurement tube222and/or optical fiber228to the outer wall of the proximal portion of the pump-outlet tube.

In some embodiments, two or more bands298are spaced from one other, along the length of the pump-outlet tube, by 1-5 mm. In some embodiments, two or more bands298are displaced circumferentially from one another, e.g., by 20-120 degrees. Advantageously, this circumferential displacement may provide a twist to the pump-outlet tube upon blood being pumped through the pump-outlet tube, thus orienting the pump-outlet tube even more precisely. For some applications, the axial spacing between adjacent bands is non-uniform along the length of the pump-outlet tube, in such a manner that it causes different regions of the pump-outlet tube to assume respective degrees of curvature upon blood being pumped through the pump-outlet tube.

Reference is now made toFIGS.32A-B, which collectively show an assembly of part of a left-ventricular assist device, in accordance with some applications of the present invention.

In some embodiments, as shown inFIG.32A, prior to the assembly of the left-ventricular assist device, a first section24aof pump-outlet tube24is separate from a second section24bof the pump-outlet tube. The proximal end of first section24acomprises multiple tabs302, and is shaped to define respective gaps304between successive ones of tabs302. Likewise, second section24bcomprises multiple tabs306and is shaped to define respective gaps308between successive ones of tabs306.

To assemble the left-ventricular assist device, impeller50(and frame34) are passed through the proximal end of first section24a, such that impeller50(and frame34) are disposed within first section24a. Next, as shown inFIG.32B, second section24bis bonded to first section24asuch that tabs306overlap tabs302(either inside or outside tabs302) and gaps308are continuous with gaps304so as to define one or more blood-outlet openings109that are laterally facing (by virtue of being defined by the central cylindrical portion of pump-outlet tube24). Advantageously, the overlap312between the tabs may help maintain the structural integrity of the pump-outlet tube, even as blood flows through blood-outlet openings109. In some embodiments, the first section and second section overlap one another by 0.25-1 mm.

Following the assembly of the left-ventricular assist device, the pump-outlet tube may be inserted, through the aorta of the subject, into the left ventricle of the subject's heart such that the pump-outlet tube traverses the aortic valve of the subject with blood-outlet openings109being disposed within the aorta. While the pump-outlet tube is positioned in this manner, the impeller may pump blood of the subject, through blood-outlet openings109, from the left ventricle into the aorta.

In some embodiments, the blood-outlet openings occupy 20-80% of the circumference of the pump-outlet tube. Advantageously, although overlap312between the tabs adds to the minimum diameter to which this region of the pump-outlet tube can be radially constrained, within this region the blood-outlet openings occupy much of the circumference of the pump-outlet tube. Thus, overall this region does not increase the minimum diameter to which the pump-outlet tube can be radially constrained. More generally, typically, overlap312between the tabs provides the pump-outlet tube with the right balance between strength and crimpability.

As noted above, the pump-outlet tube may be configured to curve proximally to the impeller, e.g., by virtue of being pre-shaped, by virtue of blood-outlet openings109being arranged in a non-axisymmetric arrangement, by virtue of blood-inlet openings108being arranged in a non-axisymmetric arrangement, and/or by virtue of bands298(FIG.31) being bonded to the outer wall of the pump-outlet tube. Alternatively or additionally, any of the embodiments described with reference toFIGS.32A-Bmay be combined with any of those described with reference toFIGS.19A-B,FIGS.20A-D, orFIG.29.

Reference is now made toFIGS.33A-C, which are schematic illustrations of pump-outlet tube24, in accordance with different respective applications of the present invention.

Pump-outlet tube24may be configured in various ways so as to provide various directions of blood flow from the pump-outlet tube into the aorta.

InFIG.33A, pump-outlet tube24is shaped to define blood-outlet openings109that are laterally facing (by virtue of being defined by the central cylindrical portion of pump-outlet tube24), such that, as indicated by blood-flow arrows318, blood flows laterally from the tube into the aorta while the blood-outlet openings are disposed within the aorta.

InFIGS.33B-C, on the other hand, pump-outlet tube24comprises a narrower section24c, in which the impeller is disposed, and a wider section24d, which is proximal to and wider than narrower section24c. Wider section24dis shaped to define at least a portion of each blood-outlet opening109such that a normal vector319to the portion of the blood-outlet opening has a distally-facing component. Thus, as indicated by blood-flow arrows318, the direction of blood flow from the tube into the aorta has a distal component. Given that the pumping of blood produces a distal thrust on the pump-outlet tube, this distal component may be advantageous in that, by virtue of the distally-facing component, the flow of blood through the blood-outlet openings may produce a proximal thrust on the pump-outlet tube that at least partially cancels the distal thrust.

In some embodiments, each of the blood-outlet openings spans the interface between narrower section24cand wider section24d. In other embodiments, each of the blood-outlet openings is positioned entirely within the wider section.

In some embodiments, as shown inFIG.33B, the angle α between normal vector319and longitudinal axis324of the pump-outlet tube at the wider section is between 20 and 80 degrees. In other embodiments, as shown inFIG.33C, the normal vector is parallel to the longitudinal axis of the pump-outlet tube at wider section24d.

In some embodiments, wider section24dis folded inwardly so as to define one or more surfaces322configured to direct the blood through blood-outlet openings109, as described above with reference toFIG.29. Alternatively or additionally, the embodiments shown inFIGS.33B-Cmay be combined with those shown inFIGS.32A-B, i.e., first section24amay be identical to narrower section24c, and second section24bmay be identical to wider section24d.

With regards to all aspects of ventricular assist device20described with reference toFIGS.1A-33C, it is noted that, althoughFIGS.1A and1Bshow ventricular assist device20in the subject's left ventricle, for some applications, ventricular assist device20is placed inside the subject's right ventricle, such that the device traverses the subject's pulmonary valve, and techniques described herein are applied, mutatis mutandis. For some applications, components of device20are applicable to different types of blood pumps. For example, aspects of the present invention may be applicable to a pump that is used to pump blood from the vena cava and/or the right atrium into the right ventricle, from the vena cava and/or the right atrium into the pulmonary artery, and/or from the renal veins into the vena cava. Such aspects may include features of tube24(e.g., the curvature of the tube), impeller50, features of pump-head portion27, drive cable130, etc. Alternatively or additionally, device20and/or a portion thereof (e.g., impeller50, even in the absence of tube24) is placed inside a different portion of the subject's body, in order to assist with the pumping of blood from that portion. For example, device20and/or a portion thereof (e.g., impeller50, even in the absence of tube24) may be placed in a blood vessel and may be used to pump blood through the blood vessel. For some applications, device20and/or a portion thereof (e.g., impeller50, even in the absence of tube24) is configured to be placed within the subclavian vein or jugular vein, at junctions of the vein with a lymph duct, and is used to increase flow of lymphatic fluid from the lymph duct into the vein, mutatis mutandis. Since the scope of the present invention includes using the apparatus and methods described herein in anatomical locations other than the left ventricle and the aorta, the ventricular assist device and/or portions thereof are sometimes referred to herein (in the specification and the claims) as a blood pump.

The scope of the present invention includes combining any of the apparatus and methods described herein with any of the apparatus and methods described in one or more of the following applications, all of which are incorporated herein by reference:U.S. application Ser. No. 18/447,025 to Tuval, entitled “Ventricular assist device with motion-cushioning spring,” which is a continuation of International Application PCT/IB2022/058101 to Tuval (published as WO 23/062453), entitled “Ventricular assist device,” filed Aug. 30, 2022, which claims priority from:U.S. Provisional Patent Application 63/254,321 to Tuval, entitled “Ventricular assist device,” filed Oct. 11, 2021, andU.S. Provisional Patent Application 63/317,199 to Tuval, entitled “Ventricular assist device,” filed Mar. 7, 2022;US 2023/0226342 to Tuval, entitled “Ventricular assist device,” which is the US national phase of International Application No. PCT/IB2022/051990 to Tuval (published as WO 22/189932), entitled “Ventricular assist device,” filed Mar. 7, 2022, which claims priority from:U.S. Provisional Patent Application 63/158,708 to Tuval, entitled “Ventricular assist device,” filed Mar. 9, 2021, andU.S. Provisional Patent Application 63/254,321 to Tuval, entitled “Ventricular assist device,” filed Oct. 11, 2021;US 2023/0137473 to Zipory, entitled “Centrifugal and mixed-flow impellers for use with a blood pump,” which is the US national phase of International Application No. PCT/IB2021/052590 to Zipory (published as WO 21/198881), entitled “Centrifugal and mixed-flow impellers for use with a blood pump,” filed Mar. 29, 2021, which claims priority from U.S. 63/003,955 to Zipory, entitled “Ventricular assist device,” filed Apr. 2, 2020;US 2022/0226632 to Tuval, entitled “Ventricular assist device,” which is the US national phase of PCT Application No. PCT/IB2021/052857 (published as WO 21/205346), filed Apr. 6, 2021, which claims priority from:U.S. Provisional Patent Application 63/006,122 to Tuval, entitled “Ventricular assist device,” filed Apr. 7, 2020,U.S. Provisional Patent Application 63/114,136 to Tuval, entitled “Ventricular assist device,” filed Nov. 16, 2020, andU.S. Provisional Patent Application 63/129,983 to Tuval, entitled “Ventricular assist device,” filed Dec. 23, 2020;US 2020/0237981 to Tuval, entitled “Distal tip element for a ventricular assist device,” filed Jan. 23, 2020, which claims priority from:U.S. Provisional Patent Application 62/796,138 to Tuval, entitled “Ventricular assist device,” filed Jan. 24, 2019,U.S. Provisional Patent Application 62/851,716 to Tuval, entitled “Ventricular assist device,” filed May 23, 2019,U.S. Provisional Patent Application 62/870,821 to Tuval, entitled “Ventricular assist device,” filed Jul. 5, 2019, andU.S. Provisional Patent Application 62/896,026 to Tuval, entitled “Ventricular assist device,” filed Sep. 5, 2019;US 2019/0209758 to Tuval, which is a continuation of International Application No. PCT/IB2019/050186 to Tuval (published as WO 19/138350), entitled “Ventricular assist device,” filed Jan. 10, 2019, which claims priority from:U.S. Provisional Patent Application 62/615,538 to Sohn, entitled “Ventricular assist device,” filed Jan. 10, 2018,U.S. Provisional Patent Application 62/665,718 to Sohn, entitled “Ventricular assist device,” filed May 2, 2018,U.S. Provisional Patent Application 62/681,868 to Tuval, entitled “Ventricular assist device,” filed Jun. 7, 2018, andU.S. Provisional Patent Application 62/727,605 to Tuval, entitled “Ventricular assist device,” filed Sep. 6, 2018;US 2019/0269840 to Tuval, which is the US national phase of International Patent Application PCT/IL2017/051273 to Tuval (published as WO 18/096531), filed Nov. 21, 2017, entitled “Blood pumps,” which claims priority from U.S. Provisional Patent Application 62/425,814 to Tuval, filed Nov. 23, 2016;US 2019/0175806 to Tuval, which is a continuation of International Application No. PCT/IL2017/051158 to Tuval (published as WO 18/078615), entitled “Ventricular assist device,” filed Oct. 23, 2017, which claims priority from U.S. 62/412,631 to Tuval filed Oct. 25, 2016, and U.S. 62/543,540 to Tuval, filed Aug. 10, 2017;US 2019/0239998 to Tuval, which is the US national phase of International Patent Application PCT/IL2017/051092 to Tuval (published as WO 18/061002), filed Sep. 28, 2017, entitled “Blood vessel tube,” which claims priority from U.S. Provisional Patent Application 62/401,403 to Tuval, filed Sep. 29, 2016;US 2018/0169313 to Schwammenthal, which is the US national phase of International Patent Application PCT/IL2016/050525 to Schwammenthal (published as WO 16/185473), filed May 18, 2016, entitled “Blood pump,” which claims priority from U.S. Provisional Patent Application 62/162,881 to Schwammenthal, filed May 18, 2015, entitled “Blood pump;”US 2017/0100527 to Schwammenthal, which is the US national phase of International Patent Application PCT/IL2015/050532 to Schwammenthal (published as WO 15/177793), filed May 19, 2015, entitled “Blood pump,” which claims priority from U.S. Provisional Patent Application 62/000,192 to Schwammenthal, filed May 19, 2014, entitled “Blood pump;”U.S. Pat. No. 10,039,874 to Schwammenthal, which is the US national phase of International Patent Application PCT/IL2014/050289 to Schwammenthal (published as WO 14/141284), filed Mar. 13, 2014, entitled “Renal pump,” which claims priority from (a) U.S. Provisional Patent Application 61/779,803 to Schwammenthal, filed Mar. 13, 2013, entitled “Renal pump,” and (b) U.S. Provisional Patent Application 61/914,475 to Schwammenthal, filed Dec. 11, 2013, entitled “Renal pump;”U.S. Pat. No. 9,764,113 to Tuval, issued Sep. 19, 2017, entitled “Curved catheter,” which claims priority from U.S. Provisional Patent Application 61/914,470 to Tuval, filed Dec. 11, 2013, entitled “Curved catheter;” andU.S. Pat. No. 9,597,205 to Tuval, which is the US national phase of International Patent Application PCT/IL2013/050495 to Tuval (published as WO 13/183060), filed Jun. 6, 2013, entitled “Prosthetic renal valve,” which claims priority from U.S. Provisional Patent Application 61/656,244 to Tuval, filed Jun. 6, 2012, entitled “Prosthetic renal valve.”