Cardiac assist devices with high frequency operation

Abstract

A cardiac assist device with an expandable cup (4) having a shipping state and an operating state, the expandable cup comprising a plurality of inflow openings (5), an outflow nozzle (6) and an inflatable balloon (8) positioned inside the expandable cup (4). A catheter assembly (3) is connected to the inflatable balloon (8) during operation, and a control unit (2) is connected to the catheter assembly (3). The control unit (2) is arranged to operate the inflatable balloon (8) with a frequency of more than 100 beats per minute.

Description

[Technical field]

[0001] The present invention relates generally to circulatory assist devices and systems, and more particularly to cardiac assist devices that can be delivered percutaneously into cardiovascular lumens and are capable of pumping blood at flows high enough to assist patients in cardiogenic shock, acute myocardial infarction, acute heart failure, or during high-risk percutaneous coronary intervention, or other situations requiring hemodynamic support with reduced levels of hemolysis. [Background technology]

[0002] 2. Background of the Invention For patients suffering from cardiogenic shock or undergoing high-risk percutaneous coronary intervention (PCI), the patient's cardiac function may be compromised and therefore the use of a circulatory assist device may be required to maintain adequate blood flow through the circulatory system. Although some variability exists depending on the patient's size and condition, circulatory assist devices for patients undergoing high-risk PCI typically must produce at least 3 L / min of blood flow to maintain adequate circulation, while for patients in cardiogenic shock, a minimum of 5 L / min is generally considered necessary.

[0003] The most common types of circulatory assist devices are the intra-aortic balloon pump (IABP), the extracorporeal membrane oxygenation (ECMO) system, and the impeller-based blood pump. The IABP is a catheter with an inflatable balloon that is placed in the descending aorta and can be periodically inflated to displace blood. The ECMO system includes a venous catheter to remove deoxygenated blood from the venous system, an extracorporeal oxygenator and pump, and an arterial catheter to return blood to the arterial system, thus bypassing the heart. The impeller pump system has a rotating impeller that is placed in the heart's ventricles or major blood vessels and can be rotated at relatively high speeds to propel blood through the circulatory system.

[0004] Although providing some benefits in increasing blood flow and reducing the load on the heart, currently available circulatory support devices suffer from certain drawbacks. IABPs cannot adequately improve flow to support patients when the heart is significantly compromised, such as during cardiogenic shock. ECMO systems may have higher morbidity rates associated with multiple catheter procedures, including bleeding, thrombosis, and infection, and problems associated with membrane oxygenation, including cognitive impairment and stroke. In addition, they increase afterload, which is generally considered counterproductive. Impeller pump systems, when operated at higher speeds to produce higher flows, may result in excessive hemolysis, and furthermore, when impeller pumps are made larger to produce higher flows, the profile of such devices may become undesirably large, inhibiting percutaneous delivery, increasing the risk of damage to cardiovascular structures, and / or causing limb ischemia. As a result, current impeller-type pumps capable of providing the high flows necessary for patients in cardiogenic shock are often too large for intravascular delivery, therefore requiring surgical placement, and furthermore, producing undesirable levels of hemolysis.

[0005] What is needed, therefore, is a circulatory support device capable of producing blood flow of at least 3 L / min for high-risk PCI procedures and at least 5 L / min to support patients in cardiogenic shock, which also has a compact delivery profile to enable percutaneous introduction and intravascular placement, has a small size during operation to reduce space requirements and trauma to cardiovascular structures, and minimizes hemolysis and other complications.

[0006] Background technology U.S. Patent Publication No. US-B-5,169,378 describes an intraventricular assist pump having an expandable outer chamber and an internal balloon that can be sequentially inflated and deflated to generate a pumping action.

[0007] International Patent Publication No. WO 2008 / 113785 (Patent Document 2) discloses a device for circulating bodily fluids including a catheter device having at least one inlet section for receiving bodily fluid at a first location, an outlet section located a distance away from the inlet section for discharging the bodily fluid at a second location located a distance away from the first location, and a pump device for directed transport of the bodily fluid between the inlet and outlet sections of the catheter device.

[0008] International Patent Publication No. WO2015 / 131879 (Patent Document 3) discloses a catheter device for directing bodily fluids in a directed manner. The catheter device includes a sleeve having an interior space, a frame, and at least three openings, the sleeve being configured as a conduit for fluids between the first and second openings. A balloon of a balloon catheter is placed through the third opening in the sleeve and can be inflated and deflated to transport bodily fluids through the sleeve. [Prior art documents] [Patent documents]

[0009] [Patent Document 1] U.S. Pat. No. 5,169,378 [Patent Document 2] International Publication No. 2008 / 113785 [Patent Document 3] International Publication No. 2015 / 131879 Summary of the Invention [Means for solving the problem]

[0010] The present invention seeks to provide a cardiac assist device that can be easily inserted into an operating location and is capable of providing cardiac assist functions in an efficient and reliable manner.

[0011] According to the present invention, cardiac assist devices, systems, and methods are provided that produce high blood flow while having a low device profile for improved intravascular deliverability, less hemolysis, and reduced trauma to cardiovascular structures. The cardiac assist device may include a pumping device positionable within a cardiovascular lumen and including a pumping mechanism, an inlet, and an outlet. The pumping device may have a shipping state with a delivery profile suitable for insertion through a peripheral vessel and may be expandable within the cardiovascular lumen to an operating state, which minimizes engagement with cardiovascular tissue during delivery and operation. In the operating state, the pumping device is configured to produce a blood flow of at least 3 L / min, preferably at least 5 L / min. At the same time, the pumping device is collapsible to a delivery profile of 18 Fr or less, and in some embodiments, 14 Fr or less. Also, in some embodiments, the pumping device is capable of producing blood flows of 6-10 L / min or more, either in one or more short bursts or over a sustained period, which can be extremely beneficial for cardiogenic shock patients whose organs are in severe distress.

[0012] In preferred embodiments, the pumping device comprises an expandable cup having a pumping chamber, and the pumping mechanism comprises a volume displacement member within the pumping chamber configured to cyclically move between low and high volume states at a frequency significantly higher than the patient's natural heart rate, preferably 2-10 times higher than the natural heart rate, and in some embodiments up to 100 times higher.

[0013] The term "cardiac assist device" or "pumping device" as used herein, except as may be otherwise described with respect to any particular embodiment, is intended to include various types of internal pumps for use in various cardiovascular lumens, including, but not limited to, percutaneous ventricular assist devices (pVADs), transvalvular pVADs, and intravascular and intraventricular balloon pumps. Use in the left side of the heart and the arterial system, and in the right side of the heart and the venous system are contemplated.

[0014] A volume displacement member according to the present invention may comprise any of a variety of types of mechanisms capable of displacing a volume of fluid in a cyclically repeating manner. In a preferred embodiment, the volume displacement member comprises an inflatable balloon that can be inflated with fluid to a high volume state and partially or fully deflated to a low volume state. In other embodiments, a piston, bellows, accordion-style expandable body, or other type of volume displacement member may be used. The volume displacement member would be capable of cyclically moving at frequencies disclosed herein between a low volume state in which it occupies a smaller portion of the pumping chamber and a high volume state in which it occupies a substantially larger portion of the pumping chamber, thus displacing blood therefrom.

[0015] In some embodiments, the pumping device includes an expandable cup defining a pumping chamber and a pumping mechanism disposed within the pumping chamber. The expandable cup may have at least one, and preferably multiple, inlet openings in its walls that allow blood to flow into the pumping chamber. Additionally, an outlet nozzle in communication with the pumping chamber may be coupled to the expandable cup. In certain embodiments, one or both of the inlet opening and the outlet nozzle may include a one-way valve.

[0016] In preferred embodiments, the pumping mechanism includes a volume displacement member, which may be an expandable member such as a balloon. The balloon is preferably configured to be cyclically inflated and deflated at a frequency substantially higher than the patient's natural heartbeat. In some embodiments, the balloon is expandable from a collapsed configuration to an expanded configuration and is substantially non-distensible beyond the expanded configuration. The expandable cup is configured to be expandable from a low-profile delivery configuration in a shipping state to a deployed configuration in an operating state and is preferably substantially non-distensible beyond the deployed configuration.

[0017] In preferred embodiments, the expandable cup will have a maximum diameter in the operating state of less than 25 mm, more preferably less than 18 mm, and in some cases between 3 and 12 mm. In some embodiments, the internal volume of the expandable cup in the operating state (i.e., pumping chamber volume) is between 0.3 and 20 ml, and the volume displacement member may be periodically alternated between low and high volume states at a frequency of 100 to 10,000 beats per minute. In other embodiments, the internal volume of the expandable cup in the operating state is between 1 and 20 ml, and the volume displacement member may be alternated between low and high volume states at a frequency of 100 to 5,000 beats per minute. In yet other embodiments, the internal volume of the expandable cup in the operating state is between 5 and 10 ml, and the volume displacement member may be alternated between low and high volume states at a frequency of 200 to 2,000 beats per minute.

[0018] It should be understood that the term "pulsation" as used herein refers to cycles of movement of the volume displacement member from a low volume state to a high volume state and back again, e.g., cycles of balloon inflation and deflation. Thus, for example, operating a cardiac assist device at a frequency of 1,000 beats per minute means that the volume displacement member is cyclically moved between low and high volume states 1,000 times per minute. These "pulses" or "cycles" may or may not be synchronously timed with the natural beating of the patient's heart.

[0019] In a preferred embodiment, a tubular outflow nozzle is coupled to the expandable cup in communication with the pumping chamber, and blood flows into the outflow nozzle through an outflow opening in the pumping chamber. The outflow nozzle may extend to direct blood away from the pumping chamber to a downstream cardiovascular location. The outflow nozzle may also include or be a one-way valve that allows blood to flow out of the pumping chamber and nozzle and prevents backflow of blood through the nozzle back into the pumping chamber. The term "nozzle" as used herein may include a tubular structure having a reduced diameter or constriction relative to the pumping chamber of the expandable cup, or in some cases may refer to a conduit that is not substantially smaller in diameter than the pumping chamber.

[0020] In an embodiment, the expandable cup, the outflow nozzle, the volume displacement member, or a combination thereof, is configured to create a Venturi effect in the blood during operation. In some cases, this is due to a reduced diameter of the outflow nozzle relative to the pumping chamber. A pressure gradient is created between the pumping chamber and the outflow nozzle, which causes blood to flow at a higher rate out of the pumping chamber, which in turn draws more blood into the pumping chamber through the inflow opening. This supplements the flow created by the volume displacement member, such that the actual flow exiting the device exceeds that which would be achieved by blood displacement alone by the volume displacement member. Thus, if the volume displacement member is operating at a frequency F, a volume of blood V is drawn in each cycle (e.g., each inflation of the balloon). d When displacing F, blood flows out of the outlet nozzle at an exit flow rate R, where R>F×V d It is.

[0021] In some embodiments, the cup element is configured for placement in the left ventricle (LV) and the outflow nozzle is configured to extend through the aortic valve into the aorta so that blood flows out of the cup and into the aortic lumen. The outflow nozzle may have a length selected such that an outlet port at its proximal (downstream) end is located in the ascending aorta, the aortic arch, or the descending aorta. To avoid any flow restrictions resulting from the angle of the aortic arch relative to the left ventricle, the outflow nozzle may be positionable or pre-shaped at a non-zero angle or curve with respect to the longitudinal axis of the expandable cup to align the outflow nozzle with the aortic lumen. Preferably, the configuration of the outflow nozzle is selected to minimize disturbance of the flow of blood exiting the pumping chamber, minimize turbulence, and maintain laminar flow. The outflow nozzle may be flexible to pass through the aortic valve and conform to the shape of the aorta with little trauma to tissue, directing blood flow directly downstream in alignment with the longitudinal axis of the aorta. The exterior of the outflow nozzle may be configured to provide an atraumatic surface against which the aortic valve leaflets may close and seal, at least during diastole, and in some cases, during systole as well. Other than passage of the outflow nozzle through the aortic valve, the cup element is preferably configured to leave the aortic and mitral valves of the heart unobstructed in the operating state.

[0022] In a preferred embodiment, the cup element may include a self-expanding resilient support member. The support member preferably comprises a material, geometry, and other structural properties selected to resist expansion in diameter beyond the deployed configuration when the pumping chamber fills with blood and the balloon is inflated, producing increased pressure within the pumping chamber. Such non-distensibility allows the spacing between the cup element and the ventricular wall to be maintained, minimizing trauma to the cardiac tissue and increasing pump efficiency. In addition, the support member may be configured to resist collapse when the pumping chamber is under negative pressure during balloon deflation. At the same time, the support member should be configured to be collapsible or crimpable to a lower profile shipping state when exposed to sufficient external force to enable intravascular delivery and retrieval. In a preferred embodiment, the support member may comprise a resilient metal skeleton, such as a nickel-titanium alloy, which may be in the form of an expandable woven wire, mesh, basket, or monolithic tube having an array of openings, slits, or cells that allow expansion in at least one dimension from a shipping state to a delivery state.

[0023] In embodiments configured for placement within the left ventricle, the support member is preferably configured to extend from within the ventricle to a location within the ascending aorta to support the outflow nozzle. Alternatively, a nozzle support member separate from that of the cup may be coupled to the outflow nozzle and provide support thereto. The nozzle support member may be directly or indirectly coupled to the cup support member, or it may not be attached thereto. The nozzle support member may be tubular so as to completely surround the outflow nozzle, or it may be partially cylindrical or generally flat so as to extend along the lateral side of the outflow nozzle on one side of the aortic lumen, and may transition from one shape to another along its longitudinal axis, i.e., from generally circumferential to partially cylindrical to generally flat. The proximal (downstream) end of the support member may be shaped and configured to facilitate retrieval of the device, for example, having a tapered or rounded end. The nozzle support member and / or cup support member may include a retrieval coupling, such as a loop, knob, or hook, which may be coupled to or captureable by a retrieval device, such as a wire, snare, sheath, or catheter. Alternatively, one or more retrieval wires may be coupled to an end of the support member and configured to remain coupled thereto throughout the procedure, such that following the procedure, the wires may be retracted to retract the support member, along with the cup and balloon, into the sheath or other capture means, collapse it to a shipping state, and withdraw it from the patient.

[0024] A blood impermeable membrane preferably extends over at least a portion of the inner and / or outer surface of the support member. In some embodiments, the support member is embedded in the membrane or sandwiched between the inner and outer membranes. Preferably, the support member / membrane combination is substantially non-distensible in the operating state under pressures of up to 400 mmHg, more preferably up to 800 mmHg or more in the pumping chamber. The support member and / or membrane may include at least one inflow opening in communication with the pumping chamber through which blood flows from the ventricle into the chamber during balloon deflation, and an outflow nozzle through which blood is directed from the chamber during balloon inflation. The inflow openings may each include a one-way valve for allowing blood to flow into the pumping chamber and preventing blood flow out of the chamber through the inflow opening, which may be formed from the same material as the membrane, or may be a different material that is joined to the membrane by welding, bonding, adhesives, or mechanical fasteners. The outlet nozzle may comprise a polymeric tube of the same polymer as the membrane or a different polymer, which may be monolithically formed with the membrane or joined thereto by welding, gluing, or other means.

[0025] The cardiac assist device may further comprise a catheter assembly coupled to the volume displacement member and / or the expandable cup, which is configured to operate the volume displacement member to alternate between a low volume state and a high volume state. The catheter assembly and the volume displacement member may be permanently attached to the expandable cup, or may comprise a separate subsystem that is removable from the cup and allows the cup to be placed in the cardiovascular lumen by itself, and then inserts the volume displacement member into the cup using the catheter assembly. The catheter assembly would be configured to extend from a location outside the patient through the vascular system to a location of the pumping device, for example, from the femoral artery in the patient's groin area, through the aorta, and into the left ventricle. The catheter assembly is configured to be coupled to a control unit (described below) located outside the patient.

[0026] In the inflatable balloon embodiment, the catheter assembly is fluidly coupled to the inflatable balloon and has an inflation lumen for delivering inflation fluid to the balloon. The catheter assembly may be configured to optimize delivery of the inflation fluid and enable high frequency inflation of the balloon. Preferably, the inflation lumen has a diameter of at least 1 mm. In a specific embodiment, the inflation lumen has a diameter of between 1 and 20 mm. 2 , in some cases, 2-7 mm 2 3 mm. Additionally, the first portion of the catheter assembly and / or inflation lumen may have a larger diameter than the second portion of the catheter assembly / inflation lumen, the second portion being closer to the inflatable balloon than the first portion. The catheter assembly may also be partially or completely made of a relatively rigid material or combination of materials and / or have a wall thickness of, for example, 0.1-0.3 mm, selected to minimize expansion, collapse, or distortion of the inflation lumen under the pressure of the inflation fluid. The inflation fluid may also be cooled by the control unit, and in some embodiments, the catheter assembly may have a thermal isolation coating on its exterior to maintain the temperature of the inflation lumen at a level to enhance high velocity flow through the inflation lumen. Optionally, the catheter assembly may further include a guidewire lumen extending axially therethrough to a guidewire port at the distal end to allow the catheter assembly, together with the cup and balloon, to be slidably advanced over a guidewire to a desired location within the cardiovascular system. The catheter assembly may further include a pressure lumen having an opening at or near its distal end to allow blood pressure in or around the cup to be measured during the procedure. Alternatively, the catheter assembly may include a pressure transducer or other sensor coupled to or near its distal end for such pressure measurement. Other sensors may also be provided on the catheter assembly to sense cardiovascular or pumping parameters in or around the cup.

[0027] The cardiac assist device of the present invention provides a unique combination of compact structure in both shipping and operating states to improve deliverability while reducing potential negative interactions with the heart and blood vessels, but maintaining high-pressure pumping capabilities. Advantageously, the cardiac assist device pumps blood at flow rates of at least 3 L / min, preferably 5 L / min or more, without causing excessive hemolysis, making it suitable for use with a wider variety of cardiovascular procedures and capable of addressing a wider range of patient conditions than known cardiac assist devices. In specific embodiments, by adjusting the cup volume, frequency and net volume change of the volume displacement member, and other parameters, a flow rate of at least 3 L / min is possible for high-risk PCI support, while a flow rate of at least 5 L / min, preferably at least 6 L / min, and in some embodiments up to 10 L / min or more is possible for treating cardiogenic shock.

[0028] A cardiac assist system according to the present invention may comprise a cardiac assist device as described herein and a control unit coupled thereto and configured to operate the volume displacement member at a frequency up to 10 times, and in some embodiments up to 100 times, the patient's natural heart rate, i.e. up to 1,000 beats per minute, and in some cases between 1,000 and 10,000 beats per minute.

[0029] In embodiments employing an inflatable balloon, the control unit is configured to deliver the inflation fluid and adjust pumping parameters to provide the desired high blood flow rate from a highly compact pump. The control unit may be configured to deliver a selected inflation fluid, including a very low viscosity fluid, e.g., an inert gas such as helium, at a pressure and temperature selected to enable cyclic expansion of the volume displacement member, e.g., the inflatable balloon, at a frequency up to 10 times, and in some embodiments up to 100 times, the patient's natural heartbeat, i.e., up to 1,000 beats per minute, and in some cases, 1,000 to 10,000 beats per minute.

[0030] Preferably, the control unit allows for user adjustment or tuning of the frequency of the volume displacement member so that an appropriate frequency can be selected by the user for a particular patient and procedure, or the frequency can be changed during the procedure according to the needs of the patient. In addition, the control unit may allow for adjustment of the volume displaced by the volume displacement member, i.e., its volume in a low volume state, in a high volume state, or both.

[0031] In another embodiment, the volume change of the volume displacement member is generated by changing the pressure of the volume enclosed by the volume displacement member, e.g., an inflatable balloon. This can be accomplished by pressurizing and depressurizing the enclosed volume through a connecting inner lumen of the catheter shaft.

[0032] In embodiments using an inflatable balloon as the volume displacement member, the control unit may further have the overall system configuration, component layout, fluid circuit design, system volumes, and duty cycles selected to allow for balloon inflation at such frequencies. In an exemplary embodiment, the control unit compresses a reservoir containing inflation fluid that is connected to the catheter assembly. In this embodiment, compression or pressurization of the reservoir results in inflation fluid being delivered to the volume displacement member at the distal end of the catheter.

[0033] An exemplary embodiment comprising a safety diaphragm comprises a high pressure source, a low pressure source, and a switching arrangement connected to the high pressure source, the low pressure source, and the catheter assembly, the switching arrangement being arranged to alternately connect the high pressure source and the low pressure source to a fluid reservoir connected to the catheter assembly. In a specific embodiment, the control unit is arranged to connect the high pressure source to the catheter assembly during an inflation period and the low pressure source to the catheter assembly during a deflation period, the inflation period being shorter than the deflation period. In a preferred configuration, the duty cycle of the inflation period and the deflation period is between 30% and 80%. In an exemplary embodiment, the high pressure source is arranged to provide a maximum air pressure of at least 300 millibars relative to atmospheric pressure and an exhaust pressure of less than -100 millibars. In some embodiments, the control unit is configured to pressurize the fluid in the reservoir to a maximum pressure of at least 100 mmHg and a minimum pressure of less than -50 mmHg relative to atmospheric pressure, preferably a maximum pressure of at least 200 mmHg and a minimum pressure of less than -200 mmHg.

[0034] The present invention further comprises: delivering the pumping device in a lower profile transport state through a peripheral blood vessel into a cardiovascular lumen; Expanding an expandable cup of a pumping device to an operative state within a cardiovascular lumen, the expandable cup having a pumping chamber therein and at least one inflow opening and at least one outflow nozzle communicating with the pumping chamber; cyclically alternating a volume displacement member disposed within the pumping chamber between a low volume state and a high volume state to allow blood to flow into the pumping chamber through the inlet opening and to direct blood out of the pumping chamber through the outlet nozzle; Including, The volume displacement member provides a method of providing cardiac assistance to a patient in which the volume displacement member is cyclically alternated between low volume and high volume states at a frequency of 100 to 10,000 beats per minute.

[0035] In another embodiment, a method for providing cardiac assist to a patient's heart includes the steps of: delivering the pumping device in a delivery state through a peripheral blood vessel into a cardiovascular lumen; Expanding an expandable cup of a pumping device to an operative state within a cardiovascular lumen, the expandable cup having a pumping chamber therein and at least one inflow opening and at least one outflow nozzle communicating with the pumping chamber; periodically alternating a volume displacement member disposed within the pumping chamber between a low volume state and a high volume state to allow blood to flow into the pumping chamber through the inlet opening and direct blood out of the pumping chamber through the outlet nozzle to produce a blood flow of at least 5 L / min; Including, The expandable cup has a maximum volume of less than 20 ml in an operating condition.

[0036] In a specific embodiment, the expandable cup has a maximum volume of less than 10 ml, more preferably less than 5 ml.

[0037] In yet another embodiment, a method of providing cardiac assistance to a patient includes: delivering the pumping device in a delivery state through a peripheral blood vessel into a cardiovascular lumen; Expanding an expandable cup of a pumping device to an operative state within a cardiovascular lumen, the expandable cup having a pumping chamber therein and at least one inflow opening and an outflow nozzle communicating with the pumping chamber; periodically alternating a volume displacement member disposed within the pumping chamber between a low volume state and a high volume state to allow blood to flow into the pumping chamber through the inlet opening and direct blood out of the pumping chamber through the outlet nozzle to produce a blood flow of at least 5 L / min; Including, The expandable cup has a delivery profile of less than 18 Fr in the shipping state.

[0038] Preferably, the expandable cup has a delivery profile in the shipping state of 14 Fr or less, more preferably 12 Fr or less.

[0039] In yet another embodiment, a method of providing cardiac assistance according to the present invention comprises: Delivering a pumping device in a delivery state through a peripheral blood vessel into a left ventricle of the heart, the pumping device having a delivery profile of 14 Fr or less in the delivery state; expanding at least a portion of a pumping device into an operative state within the left ventricle, the pumping device having a pumping mechanism, an inlet, and an outlet, the outlet being positioned within a lumen of the aorta downstream of the left ventricle; and operating the pumping mechanism to allow blood to flow into the inlet and direct blood out the outlet at a flow rate of at least 5 L / min.

[0040] In other embodiments, lower flows can be achieved with lower profile devices, hi some embodiments, the pumping device produces at least 3 L / min and the delivery profile of the pumping device in the transport state is 12 Fr or less, more preferably 10 Fr or less, desirably 8 Fr or less.

[0041] In another embodiment, a method of providing cardiac assist to a patient's heart includes: delivering the pumping device in a delivery state through a peripheral blood vessel into a cardiovascular lumen; Expanding at least a portion of a pumping device to an operative state within a cardiovascular lumen, the pumping device having a pumping mechanism, an inlet, and an outlet; operating the pumping mechanism so that blood flows into the inlet and out of the outlet at a flow rate of at least 5 L / min; Including, Red blood cells in the blood are exposed to a maximum shear stress of less than 400 Pa between the inlet and the outlet.

[0042] In yet another embodiment, a method of providing cardiac assist to a patient's heart includes the steps of: delivering the pumping device in a delivery state through a peripheral blood vessel into a cardiovascular lumen; Expanding at least a portion of a pumping device to an operative state within a cardiovascular lumen, the pumping device having a pumping mechanism, an inlet, and an outlet nozzle; a volume displacement member disposed inside the pumping chamber is cyclically moved between a low volume state and a high volume state at a frequency F, the volume displacement member displacing a volume V of blood from the pumping chamber during each cycle; d Displacing Including, Blood exits the outlet nozzle at an exit flow rate R, where R>F×V d It is.

[0043] In any of the methods of the invention, the pumping device may comprise a cardiac assist device according to any of the embodiments described herein. In a preferred embodiment, the volume displacement member comprises an inflatable balloon that is cyclically alternated between low and high volume states by cyclically inflating and deflating the balloon, either partially or fully. In addition, the pumping device may be part of a cardiac assist system that includes a control unit as described elsewhere herein.

[0044] Other aspects of the nature and advantages of the present invention will become apparent from the following detailed description considered in conjunction with the drawings. [Brief description of the drawings]

[0045] The invention will be discussed in more detail below with reference to the accompanying drawings.

[0046] [Figure 1A] 1A and 1B show partial side views of an embodiment of a cardiac-assist device according to the present invention in which an inflatable balloon of the cardiac-assist device is inflated and deflated, respectively. [Figure 1B]1A and 1B show partial side views of an embodiment of a cardiac-assist device according to the present invention in which an inflatable balloon of the cardiac-assist device is inflated and deflated, respectively.

[0047] [Diagram 2] FIG. 2 shows a schematic diagram of a cardiac assist device, according to an embodiment of the present invention.

[0048] [Figure 3A] FIG. 3A illustrates the cardiac assist device of FIG. 3 positioned within the left ventricle of a patient's heart.

[0049] [Figure 3B] FIG. 3 shows a perspective view of a cardiac assist device according to a further embodiment of the present invention.

[0050] [Figure 4] FIG. 4 shows a schematic diagram of a control unit of a cardiac assist device, according to an embodiment of the present invention.

[0051] [Figure 5A] 5A and 6A are side elevational views of a distal portion of a cardiac assist device according to the present invention, showing its outflow valve in closed and open positions, respectively. [Figure 5B] 5B and 6B are cross-sections of the cardiac assist device of FIGS. 5A and 6A, respectively.

[0052] [Figure 6A] 5A and 6A are side elevational views of a distal portion of a cardiac assist device according to the present invention, showing its outflow valve in closed and open positions, respectively. [Figure 6B] 5B and 6B are cross-sections of the cardiac assist device of FIGS. 5A and 6A, respectively.

[0053] [Figure 7A]7A and 8A are cross-sections of an expandable cup of a cardiac assist device according to the present invention with an inflatable balloon in inflated and deflated configurations, respectively. [Figure 7B] 7B and 8B are cross-sectional views of a portion of the wall of the expandable cup and the inlet valve shown in FIGS. 7A and 8A, respectively.

[0054] [Figure 7C] 7C and 8C are side elevational views of the expandable cup of the cardiac assist device shown in FIGS. 7A-B and 8A-B, respectively.

[0055] [Figure 7D] 7D and 8D are cross-sections of the outflow nozzle and outflow valve of the cardiac assist devices of FIGS. 7A-C and 8A-C, respectively.

[0056] [Figure 8A] 7A and 8A are cross-sections of an expandable cup of a cardiac assist device according to the present invention with an inflatable balloon in inflated and deflated configurations, respectively. [Figure 8B] 7B and 8B are cross-sectional views of a portion of the wall of the expandable cup and the inlet valve shown in FIGS. 7A and 8A, respectively.

[0057] [Figure 8C] 7C and 8C are side elevational views of the expandable cup of the cardiac assist device shown in FIGS. 7A-B and 8A-B, respectively.

[0058] [Figure 8D] 7D and 8D are cross-sections of the outflow nozzle and outflow valve of the cardiac assist devices of FIGS. 7A-C and 8A-C, respectively.

[0059] [Figure 9] 9A-C are side elevational views showing a cardiac assist device, according to the present invention, installed within a delivery sheath in a shipping state, a partially deployed state, and an operational state, respectively.

[0060] [Figure 10A] FIG. 10A is a side elevational view of an expandable cup in a cardiac assist device according to the present invention, in a further embodiment thereof.

[0061] [Figure 10B] 10B and 10C are close-up views of the proximal end of the outflow nozzle of the cardiac assist device of FIG. 10a, in two different embodiments thereof. [Figure 10C] 10B and 10C are close-up views of the proximal end of the outflow nozzle of the cardiac assist device of FIG. 10a, in two different embodiments thereof.

[0062] [Figure 11A] FIG. 11A is a side elevational view of a cardiac assist device of the present invention showing blood flow within its expandable cup.

[0063] [Figure 11B] FIG. 11B is an enlarged view of a portion of the expandable cup of the cardiac assist device of FIG. 11A.

[0064] [Figure 11C] FIG. 11C is a cross-section of the expandable cup of the cardiac assist device of FIG. 11A, and FIG. 11D is a partial wall cross-section thereof. [Figure 11D] FIG. 11C is a cross-section of the expandable cup of the cardiac assist device of FIG. 11A, and FIG. 11D is a partial wall cross-section thereof.

[0065] [Figure 12A] FIG. 12A is a side elevational schematic view of a cardiac assist device, in accordance with the present invention, being placed through a delivery sheath.

[0066] [Figure 12B] 12B-F are cross-sections of the cardiac assist device of FIG. 12A at various points along its length. [Figure 12C]12B-F are cross-sections of the cardiac assist device of FIG. 12A at various points along its length. [Figure 12D] 12B-F are cross-sections of the cardiac assist device of FIG. 12A at various points along its length. [Figure 12E] 12B-F are cross-sections of the cardiac assist device of FIG. 12A at various points along its length. [Figure 12F] 12B-F are cross-sections of the cardiac assist device of FIG. 12A at various points along its length. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0067] Description of the embodiments The present invention provides intraluminal cardiac assist devices, systems, and methods that are effective in supporting blood circulation in a patient. Cardiac assist devices according to embodiments of the present invention can function to provide circulatory assistance in a patient by pumping blood from a cardiovascular lumen at higher flow rates, reduced hemolysis, and improved deliverability compared to known devices. A "cardiovascular lumen" as used herein includes the interior of a blood vessel lumen in either the arterial or venous system, a cardiac chamber such as the left or right ventricle or atrium, or any other organ or vessel in the cardiovascular system.

[0068] As shown in the side views of Figures 1A and 1B and the schematic diagram of Figure 2, a cardiac assist system 100 according to the present invention includes a cardiac assist device (or pumping device) 1 and a control unit 2. The cardiac assist device 1 comprises an expandable cup 4 having an internal pumping chamber 30, a plurality of inflow openings 5 ​​and outflow nozzles 6 communicating with the pumping chamber 30, a volume displacement member 8 positioned within the pumping chamber 30 inside the expandable cup 4, and a catheter assembly 3 connected to the volume displacement member 8 during operation. The volume displacement member 8 is cyclically movable between a low volume state and a high volume state. In this embodiment, the volume displacement member 8 comprises an inflatable balloon that is inflatable to a high volume state and deflated to a low volume state. The control unit 2 is connected to the catheter assembly 3 and supplies inflation fluid thereto for inflation of the balloon 8.

[0069] Although volume displacement member 8 may be shown and described as an inflatable balloon in various exemplary embodiments disclosed herein, it should be understood that other types of volume displacement members may be substituted for such balloons without departing from the scope of the present invention.

[0070] Note that the expandable cup 4 has a low profile shipping state configured for intravascular delivery and a larger profile operating state, the operating state being shown in Figures 1A and 1 B. The inflow opening 5 may, for example, include a one-way inflow valve and the outflow nozzle 6 may include a one-way outflow valve 7, which is described further below.

[0071] In an embodiment of the present invention, the control unit 2 is arranged to operate the inflatable balloon 8 with a frequency of more than 100 beats per minute (bpm). The inflatable balloon 8 is deflated via the catheter assembly 3 to allow blood to enter the pumping chamber 30 of the expandable cup via the inflow opening 5 (fully deflated state 8B shown in dotted line in FIG. 1B), and subsequently, the inflatable balloon 8 is inflated, thereby expelling blood out of the outflow nozzle 6 (fully inflated state 8A shown in dotted line in FIG. 1A). Having a high frequency of deflation and inflation (>100 bpm) allows having a smaller pumping volume (internal volume of the expandable cup 4) and thus smaller dimensions and improved deliverability of the cardiac assist device compared to prior art cardiac assist devices, but maintaining a sufficiently high throughput to effectively operate as a cardiac assist device.

[0072] In a group of embodiments, the internal volume of the expandable cup 4 in the operating state is 0.3-20 ml. This makes it possible to design the cardiac assist device 1 with minimal dimensions, making it easier to position it, for example, in the left ventricle of the patient's heart. Also, transport of the cardiac assist device in a transport state (with the entire expandable cup 4 folded over the inflatable balloon 8 and the end of the catheter assembly 3) through the patient's arteries is thus very well enabled.

[0073] The control unit 2 can be arranged to operate the inflatable balloon 8 with a predetermined frequency above 100 bpm when using the higher range of indicated volumes (20 ml), or with a frequency up to 10,000 bpm when using the lower range of indicated volumes (0.3 ml) in order to obtain sufficient flow of blood through the cardiac assist device 1.

[0074] In an exemplary embodiment, the internal volume of the expandable cup 4 in an operating state is 1-20 ml, and the control unit 2 is arranged to operate the inflatable balloon with a frequency of 100-5,000 beats per minute. This sufficiently high frequency of operation combined with a relatively low pumping volume allows the cardiac assist device to maintain a sufficient blood flow rate.

[0075] In a further exemplary embodiment, the internal volume of the expandable cup 4 in an operating state is 5-10 ml and the control unit 2 is arranged to operate the inflatable balloon with a frequency of 200-1,000 beats per minute. The smaller size and associated operating frequency range allows for sufficient assisted flow of blood to provide proper functioning of the cardiac assist device 1.

[0076] The dimensions of the cardiac assist device 1, primarily the internal volume of the expandable cup 4, together with the size and position of the inflatable balloon 8 (combined with the extent to which the inflatable balloon 8 is inflated by the control unit 2), determine the (maximum) cardiac output. Combined with the frequency, this determines the blood flow through the cardiac assist device 1.

[0077] An internal volume of 20 ml can be obtained, for example, with a maximum outer diameter of the cardiac assist device 1 of 25 mm. Lower volumes may be obtained by reducing the diameter, for example, to a value of 12 mm, more preferably 10 mm, which can be offset by a higher pumping frequency, for example, of 600 bpm, in order to obtain a sufficiently high blood flow through the cardiac assist device 1. Even smaller cardiac assist devices 1, for example with an internal volume of 1 ml or even 0.3 ml, may be obtained, with a sufficient flow of blood being maintained by increasing the pumping frequency to 2,000 bpm or even 5,000 or 10,000 bpm. The diameter of the cardiac assist device 1 may therefore be between 0.3 and 10 mm.

[0078] A lower pumping frequency may be advantageous, especially in view of the long operating life of the internal balloon 8. Exemplary embodiments in this group have a pumping frequency of 200-600 bpm and an internal volume of 5-20 ml. A higher pumping frequency, for example in the range of 600-5,000 or even 10,000 bpm, combined with an internal volume in the range of 0.3-5 ml, may be advantageous in terms of the external dimensions of the cardiac assist device 1 (both in the transport state and in the operating state).

[0079] The control unit 2 may also be arranged to vary the pumping frequency, for example by operator input / adjustment or by automatic control using sensor data and control algorithms. The actual blood flow through the cardiac assist device 1 can therefore (also) be controlled using only the pumping frequency as a control parameter.

[0080] FIG. 3 shows a perspective view of a cardiac assist device 1 according to a further exemplary embodiment of the present invention, showing details of the expandable cup 4, the inflatable balloon 8, and a portion of the catheter assembly 3. FIG. 3A shows the cardiac assist device 1 positioned in the left ventricle LV of the patient's heart. In this group of embodiments, the expandable cup 4 comprises a support structure (or skeleton) 4a of a material that is expandable from a low-profile shipping state to an expanded operating state, and that preferably remains non-distensible beyond its size in the operating state and collapsible for removal. The expandable cup 4 further comprises an outer wall 4b of semi-rigid or flexible material that is connected to, surrounds, or extends across the outer surface of the support structure 4a. In addition to or instead of the outer wall 4b, an inner wall (not shown) may be provided on the inner side of the support structure 4a, which may have similar properties as the outer wall 4b and may be made of the same or a different material. The support structure 4a can be embodied, for example, using resilient metal wires woven or formed into a mesh, laser cut tubes, tubular baskets or stent-like structures, meshes or a series of ribs or fibers made from non-expanding polymeric materials such as nylon, or expandable tubing of nylon or other polymers. In a preferred embodiment, the support structure 4a is a monolithic tube or basket-like structure of a resilient metal, for example, nickel-titanium alloy (nitinol), cobalt-chromium, chromoly, or stainless steel, consisting of a plurality of struts or cells arranged in a radially expandable geometry. Preferably, the support structure 4a is collapsible to a low-profile operating state and resiliently biased to an expanded operating state. The support structure 4a may further be embedded in either the outer wall 4b or the inner wall, or sandwiched between the inner and outer walls.

[0081] The expandable cup 4 is preferably semi-rigid in the operating state due to either the support structure 4a and the outer wall 4b (and / or inner wall, if provided) or any combination of those elements, even during operation when the inflatable balloon 8 is continuously inflated and deflated inside the expandable cup 4 to increase the fluid pressure therein. By "semi-rigid" we mean that the expandable cup 4 is substantially non-distensible or distensible only by a known or predictable amount beyond the diameter to which it expands in the operating state when the balloon 8 is inflated to increase the pressure in the pumping chamber 30. This allows the expandable cup 4 to be appropriately sized to limit excessive engagement or force against the ventricular wall. Furthermore, the expandable cup 4 is configured to resist contraction or collapse under negative pressure in the pumping chamber 30 in the operating state. Preferably, the expandable cup 4 will not expand more than about 3% or 0.5 mm in diameter under internal fluid pressures up to 500 mmHg, while any reduction in diameter should be no more than 2% or 0.3 mm during deflation. The expandable cup may still have the flexibility to deform into non-circular shapes or bend to conform to cardiovascular lumens (e.g., the left ventricle and aorta) and would be collapsible to a shipping state for removal when exposed to sufficient radially inwardly directed force.

[0082] In the operating state, the expandable cup 4 directs the kinetic energy provided to the blood by the inflatable balloon 8. The inflatable balloon 8 is an active component of the cardiac assist device 1 and is cyclically inflated and deflated. During inflation, the inflatable balloon 8 expands and displaces volume, thereby generating kinetic energy that also increases the velocity of the displaced blood and, as a result, reduces its viscosity. To direct the accelerated blood towards the aortic system, the expandable cup 4 (comprising a membrane / outer wall 4b, an inflow valve provided in the inflow opening 5, an outflow valve 7 provided in the outflow nozzle 6, and a support structure 4a) directs the volume displacement radially and / or downstream, i.e. towards the outflow nozzle 6, which is, for example, placed through the aortic valve or, alternatively, placed in the ventricle and directed towards the aortic valve. During contraction, the outflow valve 7 closes and prevents backflow, while the inflow valve opens and allows the inflow of fresh blood from the surrounding lumen into the internal volume of the expandable cup 4.

[0083] When the expandable cup is placed in the LV, as shown in FIG. 3A, the outflow nozzle 6 will direct the fluid out of the heart. In a preferred embodiment, the outflow nozzle 6 extends from the cup 4 across the aortic valve into the ascending aorta. The outflow nozzle 6 has an outlet 31 at its proximal (downstream) end with an outflow valve 7, which in various embodiments is configured to be positioned in the ascending aorta AA, in the aortic arch, or in the descending aorta. In right-heart embodiments, the outflow nozzle 6 may be configured to extend from the expandable cup 4 positioned in the right atrium or right ventricle through the pulmonary valve into the pulmonary artery. For the outflow valve 7 to be optimal in performance, the outflow resistance should be as low as possible while being fast to open and close, with good sealing ability in the closed position. For this, the inner diameter of the nozzle 6 and outflow valve 7 depends on the flow (velocity) out of the device, which is preferably below 4 m / s for a device pumping about 5 L / min in a pulsatile manner. In such an embodiment, the outlet nozzle 6, including the outlet valve 7, has an inner diameter of about 5 to 17 mm, preferably about 12 mm or less.

[0084] As shown in Figures 5A-5B and 6A-6B, the outflow nozzle 6 comprises a support structure 32 and a dynamic structure 34. The support structure 32 may be a monolithic extension of the cup's support structure (skeleton) 4a or a separate structure that may be directly or indirectly connected to the cup's support structure 4a. The dynamic structure 34 should be connected to the support structure 32 at least in the distal region (closer to the ventricle) or in multiple locations, e.g., in the distal region, across the spine on one or more lines. In some embodiments, the entire length of the dynamic structure 34 up to its outflow end is attached to the support structure 32. The support structure 32 may be tubular to surround the dynamic structure 34 over all or part of its length, or it may be a partial cylinder or elongated flat structure that is only partially tubular in the more distal region, e.g., closer to the ventricle, and the remaining portion is configured to be positioned only on one side of the aortic lumen. In some embodiments, the supportive structure 32 has a gradual transition from cylindrical to flat. The dynamic structure 34 may be a flexible polymer and in some embodiments is the same as the membrane of the cup. The dynamic structure 34 is movable between an open configuration (shown in FIGS. 6A-B) in which blood may flow out of the outflow nozzle 6, and a closed configuration (shown in FIGS. 5A-B) in which blood is blocked from flowing back into the outflow nozzle 6 when the balloon 8 is deflated and creating negative pressure in the pumping chamber 30. The dynamic structure 34 itself may have a gradual stiffness or radial strength. For example, it may be made from a tube with a stiffer inflow end than the outflow or valve end 31 of the structure. The gradual stiffness may be provided by varying the wall thickness, material type, or material durometer hardness along the length of the dynamic structure. The length of the dynamic structure is related to the diameter and is preferably 0.5-3.5 cm, more preferably 2 cm in length, 8-14 mm, preferably 12 mm in diameter, and the wall thickness is 20-120 mu, preferably 80 mu. As illustrated in Figures 5A-B and 6A-B, the outflow valve 7 may form part of the tubular outflow nozzle 6 near the outlet end 31, which is configured to collapse and seal when the pressure therein is exceeded by that outside it.Alternatively, the outlet valve 7 may be a windsock valve, a duckbill valve, a flap valve, or a multi-leaflet valve mounted to the outlet end 31 of the outlet nozzle 6 .

[0085] In certain embodiments, the expandable cup 4, the outflow nozzle 6, the volume displacement member 8, or a combination thereof may be configured to create a Venturi effect in the blood during operation. The increasing volume of the displacement member (e.g., due to expansion in the case of an expandable volume displacement member such as an inflatable balloon) imparts kinetic energy and velocity to the displaced blood. The moving blood will drag surrounding blood in its wake, creating a Venturi effect in addition to the displacement of blood from the pumping chamber due to the increased volume of the volume displacement member. Furthermore, in some embodiments, the outflow nozzle 6 has a reduced diameter relative to that of the pumping chamber 30 such that blood flowing out of the pumping chamber 30 accelerates and causes a pressure drop. In an exemplary embodiment, the diameter of the pumping chamber 30 may range from A to B mm, while the inner diameter of the nozzle 6 may range from C to D mm, resulting in a ratio of the inner diameter of the outflow nozzle 6 to that of the pumping chamber 30 of X to Y. The pressure gradient between the pumping chamber and the outlet nozzle causes blood to flow out of the pumping chamber at a higher rate, which in turn draws more blood into the pumping chamber through the inlet. This supplements the flow generated by the volume displacement member such that the actual flow exiting the device exceeds that which would be achieved by blood displacement alone by the volume displacement member. Thus, if the volume displacement member is operating at a frequency F, a volume of blood V d When displacing F, blood flows out of the outlet nozzle at an exit flow rate R, where R>F×V d Such exit flow rates can be as high as 40-80% or more than the flow rates produced by volumetric displacement alone.

[0086] 7A-D and 8A-D illustrate the inflow valve 5 of the expandable cup 4 in closed and open configurations, respectively. The valve 5 may comprise elongated leaflets coupled to the inner or outer wall of the cup 4 and extending across an elongated inlet opening in the cup wall. In some embodiments, the leaflets are formed from the same material as the membrane of the inner or outer wall of the cup 4. As the balloon 8 expands, as shown in FIGS. 7A-D, the inflow valve 5 is in a closed configuration, which blocks blood from flowing out of the pumping chamber 30 as the pressure in the pumping chamber 30 increases. The outflow valve 7, shown in FIGS. 7C-D, opens, allowing blood to exit. As shown in FIGS. 8A-D, when the balloon 8 contracts, the pressure in the pumping chamber 30 decreases, causing the outflow valve 7 to close. When the pressure in the pumping chamber 30 becomes sufficiently low relative to the ambient pressure, the inflow valve 5 opens, allowing blood to flow into the pumping chamber 30. The inlet valve 5 is designed to have low resistance during the inlet of the device while having high sealing ability during injection. To allow high actuation frequencies, the inlet valve 5 should be capable of having fast response times for both opening and closing. To achieve this, the total inlet area of ​​the valve 5 should be at least 50 mm 2 , preferably 200 to 500 mm 2 The valves should be sized to fit the pumping chamber 30, which may be in a single valve or distributed across 2-500 valves, preferably 10-50 valves. Preferably, the valve geometry, size, number, and configuration will be selected to minimize resistance to blood flow into the pumping chamber 30. The valves may be formed from flexible, impermeable polymers, fabrics, or tissues. For high flexibility, the durometer hardness of the leaflets may be lower compared to the cups. The leaflets will be organized so that the inflow of blood is as laminar as possible. This can be accomplished by placing the valves in one orientation, for example, horizontal, vertical, or oblique planes, or combinations thereof.

[0087] The support structure 4a is constructed from a material that can be folded or collapsed into a shipping state with a small profile for delivery and expanded to a larger operating state with a consistent shape that is resistant to high internal pressures, preferably substantially non-distensible at internal pressures up to 400 mmHg (range 200-600 mmHg). The support structure 4a can be folded, for example, to fit within a 2-8 mm, more preferably 3-5 mm, lumen of a delivery sheath for introduction into a desired vascular lumen. Once in position, the support structure 4a is expanded to its operating state with a diameter of, for example, 18 mm (within a preferred range of 5-24 mm). When expanded, the support structure 4a will prevent collapse of the membrane / outer wall 4b at negative pressures as low as -50mmHg (preferred range of -10 to -100mmHg), while during inflation, a pressure increase of, for example, +250mmHg (preferred range of 100 to 600mmHg) will not result in an increase in diameter of the expandable cup 4. The transportable and operational states, including the inflation and deflation stages, will be discussed below.

[0088] In the shipping state, the support structure 4a is configured such that the maximum stretch for deformation of the material is higher than the final stretch on the material during crimping. The combination of small cup size and volume, low thickness of the cup membrane (preferably 0.01-0.15 mm) and support structure 4a (preferably 0.05-0.3 mm), and low profile shaft of the inflatable balloon (preferably 2.0-3.5 mm) allows to have a low crimping profile (preferably 8 Fr-18 Fr) in the shipping state while maintaining a high pump capacity (preferably 2 L / min-9 L / min). By varying these parameters, the device can be adapted to specific clinical needs. The preferred configuration for cardiogenic shock settings has a delivery profile of 18 Fr or less, more preferably 14 Fr or less, producing a blood flow of at least 6 L / min. For use in high-risk PCI, a delivery profile of 9 Fr or less and producing a blood flow of at least 4 L / min is preferred.

[0089] In the transport state, the device needs to remain flexible enough to pass through difficult anatomical variations in the vasculature. To this end, the support structure 4a should be flexible, for example, by having a low wall thickness, e.g., 0.15-0.25 mm in a preferred configuration. Furthermore, the design of the support structure 4a needs to be configured to be crushed or crimped into the transport state without damage that could weaken or destroy the support structure 4a. This can be achieved by selecting a material, such as a nickel-titanium alloy, that is less susceptible to microfracture, and by designing the support structure 4a to minimize microfracture. Importantly, membrane selection is critical in determining the scaffold material and geometry, e.g., the scaffold should not stretch the membrane material beyond its yield point for plastic deformation, which would compromise the effectiveness of the membrane. The scaffold 4a may include a woven wire mesh or may be formed from a monolithic tube into multiple struts in a closed or open cell configuration. The construction of the support structure 4a may be accomplished using superelastic Nitinol with heat curing up to 35 degrees Celsius. Alternatively, the cup can be made from expandable struts, for example from a nylon material, which can be pressurized (e.g., using a fluid) to provide conformity to the outer wall 4b. In one embodiment, the support member / skeleton 4a has a diamond-shaped cell design with 4.2 mm struts, which in the expanded state are about 6 mm long and in the collapsed (shipped) state are 8 mm long, resulting in a 100% elongation. In one embodiment, the membrane comprises a TPU such as Technotan, Pellatan, or Tecotan at a durometer hardness of about 72D, which tolerates these stresses without plastic deformation.

[0090] The operating or expanded state of the cup is designed not to interfere with the natural blood flow in the cardiovascular lumen in which it is located. As shown in FIG. 3A, when used in the left ventricle LV, the size and shape of the cup 4 in the operating state (expanded) is such that it leaves maximum room for the natural blood flow in and out of the left ventricle. During inflow through the mitral valve MV there should not be any obstruction, and during outflow obstruction of the aortic valve is minimized so that blood can flow out of the ventricle around the outflow nozzle 6. Preferably, the outflow nozzle 6 is configured to extend through the aortic valve AV and allow the aortic valve leaflets to seal against its outer surface during diastole. Disruption of the mitral valve MV, chordae tendineae CT, papillary muscles, or other structures of the ventricle is also minimized. Furthermore, the cup 4 may have a hydrodynamic bullet / ellipsoid or cigar-like shape that prevents compartmentalization of the flow inside the heart. By maintaining the natural blood flow, there will not be any areas of stagnant flow. Thus, with a cup diameter of less than 18 mm, preferably less than 15 mm, with a length (not including nozzle 6) of less than 60 mm in its expanded configuration, the hydrodynamic shape of the cup and the distribution of the inflow valves will help to reduce the possibility of intraventricular thrombosis formation. In addition, given the angle of the aortic valve relative to the axis of the left ventricular lumen, the cardiac assist device 1 may have a flexible, curved or angular attachment of the outflow nozzle 6 to the main body of the cup 4 to reduce resistance while maintaining laminar flow in the outflow of blood.

[0091] In operation, the surface of the cup 4 is preferably smooth to allow contact with tissue without damaging it and to prevent areas of thrombus formation. The overall smoothness is achieved by capturing the support structure 4a / scaffolding between two membranes or layers of polymer within the outer wall 4b. These layers may be connected to form one layer in which the scaffolding is embedded. The end effect is preferably a fully coated support structure 4a on both the interior and exterior of the cup 4 and nozzle 6. Typically the polymer layer is sufficiently smooth, but an additional hydrophilic coating may be added for additional smoothness. Anti-thrombogenic coatings may also be employed to reduce clot formation. The outflow nozzle 6 may further have a smooth exterior surface configured to allow the aortic valve leaflets to seal against the nozzle surface without trauma.

[0092] During balloon deflation, the cup preferably maintains its size and shape with limited to no deformation. The support structure 4a, e.g., Nitinol skeleton, in combination with the membrane, gives the cup radial strength to resist such deformation. The radial strength of the cup, i.e., the combined support structure 4a and outer wall membrane 4b, must withstand the pressure gradient between the ventricle and the pumping chamber 30 during blood inflow. The magnitude of this gradient depends on the inflow area (resistance) of the inflow valve and the speed and volume of the balloon deflation. In one embodiment, the radial strength is provided by the material properties of the membrane, e.g., Tecotan 72D, in a layer of 60 mu combined with the diamond- or hexagon-shaped cell design of the Nitinol skeleton with a wall thickness of 0.2 mm. Higher radial stiffness can be achieved by adhesion of the membrane to the support structure 4a.

[0093] During inflation of the balloon, the support structure 4a in combination with the outer and / or inner wall membrane will preferably not have any increase or have a very limited increase in diameter under higher fluid pressures within its interior. The cup should be able to withstand a pressure increase of up to 1,000 mmHg with a limited increase in diameter of about 1-5% or less. To prevent possible outward expansion during inflation, the membrane / outer wall 4b is preferably made of a material with limited or no compliance. The membrane / outer wall 4b is constructed from or coated with a material that has limited to no interaction with blood. The material is selected to be flexible and durable, such as nylon or polyurethane with a high durometer hardness value. This can be achieved by using a polymer with a high tensile modulus, for example, Tecothane 72D. The membrane / outer wall 4b is attached to or encapsulates the support structure 4a. The membrane / outer wall 4b itself has a wall thickness of, for example, 10-100 μm for a single layer. The connection of the membrane / outer wall 4b to the support structure 4a is made such that the expandable cup 4 can be collapsed in the shipping state, for example to a diameter of less than 6 mm, preferably about 3 mm, and expanded in the operating state to a diameter of 10-25 mm, for example about 18 mm. The cup 4 may also include circumferential fibers of a material with high tensile strength, such as Kevlar, Spectra, or other such materials, attached to, embedded in, or woven into the support structure 4a and / or outer wall 4b, which further limits the expandability of the support structure 4a beyond its desired size while still allowing it to be collapsed to the shipping state for insertion and removal.

[0094] The cup preferably has a high elastic recoil, which creates a higher shape consistency during both inflation and deflation, with elastic deformation preferably limited to about 0.1% and -3%, and +0.1 to 10%, with a preferred range of 0.5 to +1%. Higher shape consistency during pumping has multiple positive effects. First, it improves energy transport, making the device more efficient at transporting blood. Second, the higher stiffness allows the balloon to be operated at a higher inflation-deflation frequency. Third, this effect reduces the impact of pressure waves on the cardiovascular environment (i.e., tissues such as the heart and vessels). Fourth, by having smaller pulses, device buoyancy (due to helium) is reduced and the stability of the device is improved. Additionally, a smaller balloon with a higher frequency of inflation results in reduced recoil forces and less oscillation, which may allow it to act more like a continuous pump.

[0095] As illustrated in Figures 9A-C, in order to facilitate intravascular delivery and removal, the auxiliary device 1 is preferably positionable in the inner lumen of the sheath 36 (shown in Figure 9A) in a shipping state. The expandable cup 4 may be crimped into the shipping state using a crimping tool, or in some embodiments, may be collapsed by being advanced into the proximal end of the sheath 36 or retracted proximally into the distal end of the sheath 36. A guidewire GW may slidably extend through a guidewire lumen in at least a distal portion of the catheter assembly 3 within the expandable cup 4, over which the cardiac assist device 1 contained within the sheath 36 may be advanced through the vascular system to a desired location for operation. Once positioned at a desired location within the cardiovascular system, the sheath 36 may be retracted relative to the device 1 such that the cup 4 expands to an operating state shown in Figure 9C. After use, the cardiac assist device 1 needs to be safely removed from the body. This may be done by collapsing the cup 4 to return to the shipping state within the body. The device 1 may be retracted into the sheath 36 to compress the cup 4 so that the sheath 36 returns to the shipping state. To allow the cup 4 to collapse smoothly as it is retracted into the sheath 36, the proximal (downstream) end 38 of the outflow nozzle 6 is preferably tapered, rounded, or ellipsoidal in shape, as shown in Figures 10A-B. The downstream tip 40 of the tapered end 38 may be connected to a retrieval wire or shaft 42, which may extend through the delivery sheath from a point outside the patient, as shown in Figure 10B. Alternatively, the downstream tip 40 may include an eyelet 44, or a knob, loop, hook, or other means, to which a retrieval device, wire, or snare 46 may be connected when the cardiac assist device is to be removed, as shown in Figure 10C.

[0096] In a further embodiment, the cardiac assist device is configured to operate with an inert fluid in the catheter assembly 3 and the inflatable balloon 8 to inflate and deflate the inflatable balloon 8. This assists in operating the cardiac assist device at a relatively high frequency. The inert fluid is a low viscosity fluid, such as, for example, helium or carbon dioxide, to minimize friction with the inner wall of the catheter assembly. Helium has the added advantage of having a low density and therefore a lower mass inertia, allowing for higher inflation frequencies. To achieve a large volumetric displacement, the inflatable balloon 8 is inflated and deflated with a low viscosity medium (e.g., helium) at an adjustable or predetermined frequency. The inflatable balloon 8 is connected to the catheter assembly 3, which is optimized for high flow, as described below. The parts in contact with the inert fluid are the catheter assembly, the driver part of the control unit 2, and the inflatable balloon 8. The total volume of the inert fluid is, for example, about 16 ml (with a range of, for example, 2 ml to 35 ml), which is sufficiently low to allow high frequency operation.

[0097] The inflatable balloon 8 is, in a further embodiment, a directional action balloon arranged to allow fluid inflow into the expandable cup 4 and (directed) outflow of fluid through the outflow nozzle 6 when it is deflated. The balloon may be configured to inflate first at its distal portion (farthest from the outflow nozzle) and sequentially in the downstream direction such that the proximal portion (closest to the outflow nozzle) inflates last. Directional inflation can be achieved by shape, (for example) a conical balloon, different wall thicknesses (higher wall thickness in the area to be inflated last), or different durometer hardness of the balloon (higher durometer material in the area to be inflated last). In an exemplary embodiment, the inflatable balloon 8 can be made of various wall thicknesses along the longitudinal axis of the cardiac assist device 1 to obtain a predetermined inflation pattern, for example starting at the distal end of the expandable cup 4 and moving upwards in a direction towards the outflow nozzle 6.

[0098] The inflatable balloon 8 is designed to prevent any kind of outflow blockage from the pumping chamber 30, for example, by having a diameter smaller than the inner diameter of the expandable cup 4 in the area of ​​the outflow nozzle 6. As shown in Figures 11A-D, the pumping chamber 30 may have an outflow opening 50 through which blood flows out of the pumping chamber 30, and the shape and size of the cup 4 and balloon 8 are selected to maintain a space between the inflated balloon 8 and the outflow opening 50 and minimize resistance to the flow of blood out of the chamber (indicated by arrow B). In some embodiments, the inflatable balloon may be attached to the expandable cup at a distal end 51 of the expandable cup remote from the outflow opening 50, which helps to anchor the balloon in a fixed position relative to the cup, minimize movement of the balloon relative to the cup (other than from inflation) and reduce vibration. The inflatable balloon 8, for example, has a volume of about 8 ml (preferably in the range of 0.3-20 ml) when inflated. The material of the inflatable balloon 8 is a durable, thin, ideally 20 μm (10-60 μm) single wall material such as, for example, polyurethane or nylon. In an exemplary embodiment, the balloon 8 is made from Pellethan 55D or a material with similar mechanical properties and has a single wall thickness of 10-60 microns, ideally about 20 microns.

[0099] In a specific embodiment, the cardiac assist device 1 is configured to provide a desired blood flow rate for a particular procedure and patient condition while limiting hemolysis below an acceptable threshold. Hemolysis can be caused by the application of excessive shear stress to red blood cells for an excessively long period of time. The cardiac assist device 1 of the present invention is configured to limit the amount of shear stress to which red blood cells are exposed and to minimize the time that red blood cells are exposed to such shear stress as they pass through the cardiac assist device 1. For assistance during high-risk PCI procedures, the cardiac assist device 1 may produce a flow of at least 3 L / min while limiting shear stress to 100 Pa or less. For treating cardiogenic shock, the cardiac assist device 1 may produce a flow of at least 5 L / min while limiting shear stress to 400 Pa or less. In a particularly high-flow embodiment, the cardiac assist device 1 may produce a flow of at least 6 L / min while limiting shear stress to 3,000 Pa or less.

[0100] Preferably, the expandable cup 4 and the inflatable balloon 8 are configured to maximize flow and minimize hemolysis. As shown in Figures 11A-D, when the balloon 8 is in its inflated configuration and the cup 4 is in an operating state, a spacing 52 of at least about 0.05-1.5 mm is preferably maintained between the balloon 8 and the inner wall of the cup 4, which helps to limit hemolysis during high frequency balloon inflation. The balloon 8 is configured such that its maximum inflated diameter is below the inner diameter of the cup, e.g., about 0.1-3 mm below in diameter, when the pressure inside the balloon 8 is at least about 600 mmHg. This size difference avoids cavitation, which can occur when the balloon 8 comes into direct contact with the inner wall or membrane of the cup 4. Additionally, the spacing between the balloon 8 and the cup 4 avoids potentially trapping blood (compartmentalization) within the pumping chamber 30. Finally, the clearance between the balloon 8 and the cup 4 minimizes the shear stress on the red blood cells during the last part of the ejection phase when the balloon 8 is close to the cup wall, thereby reducing hemolysis. A clearance 52 of about 0.1 mm is sufficient to keep the shear stress low enough (less than 100 Pa at 3 L / min or less than 400 Pa at 5 L / min). Also, the exposure time to this shear rate is low (less than 5 ms) so as not to reach the threshold for hemolysis.

[0101] The catheter assembly 3 of various embodiments of the present invention also includes specific features to enable or enhance high frequency operation of the cardiac assist device and / or optimize the flow of the inert fluid (helium). In a first exemplary embodiment shown in Figures 12A-F, where the catheter assembly 3 is shown extending through a delivery sheath 36, the catheter assembly 3 is 1-20 mm in diameter. 2The balloon shaft 54 ​​has an inflation lumen 56 with a cross-sectional flow area of ​​0.5 mm. This generally corresponds to an inner diameter of 0.5-5 mm, which allows for a good balance between flow resistance with respect to the inert fluid and further properties of the catheter portion such as bend radius, kink resistance, etc. In a preferred embodiment, the inflation lumen 56 will not be subdivided, since resistance would rise when the cross-sectional area is distributed over different channels. The shape of the inflation lumen 56 should be configured to have the lowest resistance possible while leaving room in the catheter assembly for potential guidewires, pull wires for device retrieval, and potential sensors.

[0102] 12B-E, the balloon shaft 54 ​​preferably comprises three or more lumens, including an inflation lumen 56, which will usually be largest in diameter, a guidewire lumen 58, and one or more wire lumens 60 for pull wires for retrieval. The catheter assembly 3 may further comprise one or more sensors at or near its distal end, such as a pressure transducer for measuring pressure within the pumping chamber 30 of the cup 4 or within the heart outside the cup 4, a heart rate sensor, or other sensor, and the balloon shaft 54 ​​may include lumens for wires to such sensors.

[0103] In a further embodiment, the distal section of the catheter assembly 3 has a wider diameter than the proximal section of the catheter assembly, which is further from the inflatable balloon 8 than the distal section. This allows the inflation lumen in the distal section of the catheter assembly 3 to be larger than in the proximal section, thereby reducing friction to the inert inflation fluid in the larger distal section. The distal section may be configured to be positioned in a larger blood vessel further from the point of introduction toward the heart, such as in the aorta. As an example, the diameter in the distal section of the catheter assembly 1 may be 60 mm long with a diameter of 2.2 mm, and the proximal section (in the aorta, thigh area, and outside the patient) may be 1,200 mm long and 2.5 mm in diameter.

[0104] The catheter assembly 3 may further comprise multiple catheter segments with different diameters. The wider diameter segments are configured for use, for example, in areas where blood flow is not obstructed (e.g., peripheral arteries) or in areas where they remain outside the patient during operation. In one embodiment, the first segment (in the expandable cup 4) is 60 mm long with a diameter of 2.2 mm, the second segment (in the aortic area) is 800 mm long with a diameter of 3 mm, the third segment (in the femoral area) is 400 mm long and 2.5 mm diameter, and the fourth segment (outside the patient) has a length of 750 mm and a diameter of 4 mm.

[0105] In yet further embodiments, the catheter assembly 3 comprises a stiff material selected to provide low flow resistance (i.e., low impedance to inflation and deflation pressures during operation) and kink resistance. In further embodiments, the catheter assembly 3 comprises a nylon material with a wall thickness of 0.1-0.3 mm, e.g., 0.2 mm. The selection of dedicated catheter materials and dimensions allows for radial shape preservation while being sufficiently flexible in the longitudinal direction. The shaft will have high radial stiffness, achieved by a high durometer material, e.g., nylon 12 Pebax or polyimide with a wall thickness of 72D or higher, reinforced with a wire or ribbon braid or coil. The high durometer stiffness aids in rapid transport of helium. The durometer stiffness may be varied over the shaft length to accommodate the curvature of the vasa vasorum or ascending aorta.

[0106] In a further embodiment, the catheter assembly 3 includes a thermally isolating coating layer, e.g., Al, over its exterior. This will help maintain a (relatively) low temperature of the inert fluid, such as helium, thereby giving it a higher density and allowing for higher flow rates. Additionally or alternatively, the control system 2 may include an active cooling subsystem for controlling the temperature of the inert fluid delivered to the catheter assembly 3 during operation. In an exemplary embodiment, the inflation fluid may be cooled and maintained at a temperature between -20°C and 20°C.

[0107] The catheter assembly 3, together with the expandable cup 4, is configured to minimize vibration or oscillation when operated at high frequencies. When blood is ejected from the pumping chamber 30 of the cup 4 by the inflation of the balloon 8, the resulting thrust leads to a force in the opposite direction. The thrust caused by blood exiting the pumping chamber through the outlet nozzle 6 into the aorta leads to a counter force, which can move the expandable cup 4 from its equilibrium position to a deeper position in the left ventricle. Once the pump stroke is completed, the expandable cup 4 will attempt to return to its equilibrium position, driven by the pulling force from the catheter assembly 3 and the pushing force from the distal tip of the device. The size, geometry, and stiffness of the balloon shaft 54 ​​may be selected such that it acts to dampen this movement of the expandable cup 4. Furthermore, by operating the volume displacement member, e.g., the balloon 8, at a sufficiently high frequency, the next pump stroke will occur before the device has time to relax and return to its equilibrium position. In this case, the device will be "trapped" in a position away from its equilibrium position. The higher the frequency, the less time the device has to move back towards its equilibrium position and the more stable the device tip will be.

[0108] FIG. 4 shows a schematic diagram of details of an exemplary control unit 2 of a cardiac assist system 100, according to an embodiment of the present invention. In one embodiment, the control unit 2 comprises a high pressure source 21, a low pressure source 22, and a switching arrangement 23 connected to the high pressure source 21, the low pressure source 22, and the catheter assembly 3, the switching arrangement 23 being arranged to alternately connect the high pressure source 21 and the low pressure source 22 to the catheter assembly 3. This allows the control unit 2 to use commercially available hydraulic / pneumatic control components with reliable and robust (bedside) operation. In the exemplary embodiment of FIG. 4, the high pressure source 21 is implemented as a combination of a high pressure buffer 21, a high pressure compressor 21a, and a regulator 21c (and optionally a high pressure sensor 21b). The low pressure source 22 is implemented as a combination of a vacuum buffer 22, a vacuum pump 22a, and a low pressure sensor 22b. The switching arrangement 23 is implemented, for example, as a combination of a three-way valve unit 23, which is controlled via a switching valve 23a using a function generator 23c. The function generator 23c receives signals from the high pressure sensor 21b, the high pressure sensor 22b and the switching pressure sensor 23b to appropriately drive the switching valve 23a. The connection to the catheter assembly 3 is implemented via a safety drive 24.

[0109] To enable high frequency operation of the three-way valve unit 23, the valve design is selected to have optimal flow through the valve with minimal turbulence, high switching speed, and low leakage rate. The separate high and low pressure sources 21, 22 are connected, for example, with an angle of less than 45 degrees and a small difference in diameter between the tubing and the valve, such that turbulence is minimized and flow towards the safety actuator 24 is optimized.

[0110] In still further embodiments, the control unit 2 further comprises a safety driver 24 having a source side chamber 25 and a catheter side chamber 26 separated by a safety diaphragm 27. This allows the use of the pneumatic / hydraulic part of the control unit 2 to be separated from the inert gas side part of the control unit 2 and connected to the catheter assembly 3, minimizing the volume of inert gas required in the cardiac assist system 100. The safety driver 24 accommodates the actuation of the inert gas circuit by compressing and expanding the inert gas circuit in the catheter side chamber 26 of the safety driver 24. The actuation speed of the safety driver is sufficient to provide a pressure difference of +800 to -800 mmHg in a volume of, for example, 20 ml (a typical range is 5 to 70 ml) within, for example, 5 to 200 milliseconds.

[0111] In a group of embodiments, the control unit 2 is arranged to connect the high pressure source 21 to the catheter assembly 3 during the inflation period and the low pressure source 22 during the deflation period, the inflation period being shorter than the deflation period. In general, the inflation (high pressure) is applied more quickly than the deflation (low pressure) in a dynamic system, which can provide a more efficient operation for this embodiment. The inflation and deflation periods together form one operating cycle, and thus the duty cycle can be defined as the percentage of the inflation period of a complete cycle, which is preferably less than 50%. In further exemplary embodiments, the duty cycle is less than 40% or even 30-50%. In exemplary embodiments, a duty cycle of about 40% has proven to be most efficient for the cardiac assist device 1 of the present invention.

[0112] This duty cycle can be modified with varying resistive pressure: the higher the resistive pressure, the more need there may be for inflation time, while deflation can be faster when the environmental pressure also reduces the balloon volume.

[0113] The control unit 2 may further be arranged to respond to sensor data or user input. For example, the control unit 2 may respond to a sensed resistive pressure by modifying the rate of inflation, or simply generate additional pulsatile flow. Implementations may include, but are not limited to, responding to an ECG trigger adaptation to switch between a lower frequency (e.g., 200 bpm) and a higher frequency (e.g., 700 bpm), e.g., operation of the cardiac assist device 1 over a period of 1 second each, or, e.g., operation only during diastole or only during systole. Or, (short) pauses in inflation or deflation can be triggered at specific moments in the cardiac cycle.

[0114] In the exemplary embodiment shown in FIG. 4, the actuator 24 (or safety actuator) is sized relative to the total volume of the inert gas (helium) circuit. The safety diaphragm 27 may be movable to modify the volume of the inert gas circuit. By moving the safety diaphragm 27 into the source chamber 25, the total volume of the inert gas circuit may be expanded, thus depressurizing the inert gas circuit. Or, by moving it in the opposite direction, the total volume of the inert gas circuit may be reduced, thereby pressurizing it. Using this actuation, a 3 mm actuator may be used, for example, to inflate and deflate the inflatable balloon 8 within 10 milliseconds. 2 In a catheter assembly 3 of 130 cm length with a cross-sectional area of, for example, a helium pressure of -600 mmHg can be obtained in the inert gas circuit.

[0115] To further optimize the translation of pressure from the safety diaphragm 27 to the inflatable balloon 8, the length of the catheter assembly 3 should be minimized. Thus, the safety diaphragm 27 and other components of the control unit 2 are adapted to be contained within the bedside control unit 2, for example, by using external versions of the high pressure source 21 and the low pressure source 22. The bedside control unit 2 can then be mounted on the bed at a distance of 20-100 cm, and in some cases less than 30 cm, from the vascular access site on the patient.

[0116] In the embodiment shown in FIG. 4, the inert gas circuit between the safety diaphragm 27 and the inflatable balloon 8 comprises an inert gas pressure sensor 26b, whose signal can be provided to the function generator 23c for user information, driver control, and / or fault detection functions. For example, in the control unit 2, continuous checks on pressure and waveforms are applied to detect gas leaks or kinks / blockages of the catheter assembly 3. Deviations from the normal expected inert gas pressure signal provide the following information: If the pressure drops, this is an indication of an inert gas leak and the cardiac assist device 1 is stopped or switched to a vacuum / low pressure control mode. Overpressure can be caused by a kink in the catheter assembly. More advanced software may be used in the control unit 2 to infer the ventricular pressure from the inert gas (balloon) pressure.

[0117] In still further embodiments, the safety drive 24 is cooled / heated, for example, to 10 degrees Celsius, to preserve material properties, safety, and inert gas flow rate. As an alternative to this indirect temperature control of the inert fluid within the catheter assembly 3 during operation, the temperature of the inert fluid may be controlled using an active fluid cooling subsystem. Furthermore, the inert gas circuit may further comprise an automatic filling system. For example, to ensure a stable helium concentration, every two hours (or periodically at intervals of 30 minutes to four hours), the helium system is emptied and there is an automatic injection of a new helium capsule. Such capsules can be replaced in the bedside safety drive 24 when needed.

[0118] According to the method of the present invention, the cardiac assist device and system of the present invention may be used to provide cardiac assistance in various procedures and address various patient conditions. For example, the cardiac assist device and system may be used for cardiac assistance during high-risk percutaneous coronary intervention (PCI), including angioplasty and stent placement. In addition, the cardiac assist device and system may be used to provide cardiac support for patients suffering from cardiogenic shock. In addition, the cardiac assist device and system may be used to provide cardiac support for patients suffering from acute myocardial infarction. Generally, for such procedures, the cardiac assist device will be configured for placement in the left ventricle, however, placement in various other cardiovascular luminal sites, including in the ascending or descending aorta, right atrium, right ventricle, or pulmonary artery, is also possible.

[0119] In an exemplary embodiment of the method according to the invention, the cardiac assist device 1 is crimped into a shipping state having a low profile suitable for intravascular delivery. Typically, the cardiac assist device 1 will be placed in the lumen of a tubular delivery sheath having an outer diameter of less than about 18 Fr, more preferably less than 14 Fr, and in some cases less than 10 Fr, as described elsewhere herein. A guidewire may be inserted through the guidewire lumen of the catheter assembly 3 and advanced (typically through an introducer sheath) into a peripheral artery, such as the femoral artery, preferably using a percutaneous technique. The guidewire may be advanced through the femoral and iliac arteries, the aorta, and the aortic valve into the left ventricle. The sheath containing the cardiac assist device 1 is then advanced over the guidewire into the femoral artery and through the vessels and the aortic valve into the left ventricle. The sheath may then be retracted relative to the cardiac assist device 1, allowing the expandable cup 4 to self-expand into an operating state within the left ventricle. The sheath is further retracted until the outflow nozzle 6 is deployed with its outflow end in the ascending aorta. Optionally, the sheath is withdrawn from the patient or it may be left in place downstream of the outflow nozzle 6 .

[0120] The proximal end of the catheter assembly 3 is coupled to a control unit 2 outside the patient. The control unit 2 may then be activated to initiate cyclic inflation and deflation of the balloon 8 within the cup 4. Preferably, the control unit 2 includes a frequency controller that allows the user to vary the inflation frequency, speed, and select the desired inflation frequency setting. In a preferred embodiment, as described above, the balloon 8 is cyclically inflated and deflated at a frequency at least twice, in some cases up to 10 times, and in other cases up to 100 times the natural heart rate of the patient, which may be sensed using a heart monitor, heart rate sensor, ECG, or other means. In some embodiments, the means for sensing the natural heart rate is electronically connected, either by wire or wirelessly, to the control unit 2, which is adapted to automatically vary the inflation frequency according to changes in heart rate. In other embodiments, the inflation frequency may be set to a desired number of beats (inflation / deflation cycles) per minute, which may range from 200 to up to 10,000 cycles per minute, as described elsewhere herein. The inflation frequency may be set to be synchronous with the frequency of the natural heartbeat, to be a desired multiple or fraction of the frequency of the natural heartbeat, or to be completely independent of the natural heartbeat.

[0121] The blood flow rate produced by the cardiac assist device may be varied by varying the inflation frequency of the balloon 8. For high-risk PCI procedures, a flow rate of 3-4 L / min is typically required. For patients in cardiogenic shock, a higher flow rate is usually desirable, at least 5 L / min, preferably at least 6 L / min, and in some cases 8 L / min or more. Advantageously, the cardiac assist device and system of the present invention can produce such high flow rates while having a low delivery profile for percutaneous delivery, a compact operating profile to minimize trauma to cardiac tissue, and a pumping mechanism to minimize hemolysis.

[0122] Upon completion of the procedure or treatment, the control unit 2 may be deactivated so that balloon inflation stops. The delivery sheath may be advanced distally relative to the cardiac assist device 1 to a desired location, typically in the ascending aorta downstream of the outflow nozzle 6. The cardiac assist device 1 may then be retracted by applying traction to the retrieval wire 62 until the cup 4 is retracted into the inner lumen of the sheath. Advantageously, the tapered proximal end of the outflow nozzle 6 captures the cup 4 within the sheath and facilitates collapsing the cup 4 to a shipping state as it is retracted into the sheath. The sheath containing the cardiac assist device 1 may then be removed from the patient and the femoral puncture or incision may be closed.

[0123] The invention has been described above with reference to several exemplary embodiments as shown in the drawings. Modifications and alternative implementations of several parts or elements are possible and fall within the scope of protection as defined in the appended claims.

Claims

**Claim 1**: A heart assist system, comprising: An expandable cup having a transport state and an operating state, the expandable cup comprising a pumping chamber and at least one inlet opening; An outflow nozzle coupled to the expandable cup and in communication with the pumping chamber, the outflow nozzle having a length and shape selected such that when the at least one inlet opening is positioned within the left ventricle of the heart, an outlet of the outflow nozzle is downstream from the aortic valve of the heart; A volume displacement member positioned inside the expandable cup and movable between a low volume state and a high volume state; A catheter assembly connected to the volume displacement member during operation; A control unit connected to the catheter assembly ; The control unit is configured to periodically move the volume displacement member between the low volume state and the high volume state at a frequency of at least 600 beats per minute (bpm). **Claim 2**: The heart assist system according to claim 1, wherein the expandable cup has a length and shape selected to extend from a location within the left ventricle through the aortic valve into the aorta of the patient. **Claim 3**: The heart assist system according to claim 1, wherein the expandable cup comprises a support structure. **Claim 4**: The heart assist system according to claim 3, wherein the support structure is embedded within or covered by a polymer membrane. **Claim 5**: The heart assist system according to claim 3, wherein the support structure extends around at least a portion of the outflow nozzle. **Claim 6**: The heart assist system according to claim 3, wherein the support structure has a substantially the same axial extent as substantially all of the length of the outflow nozzle. **Claim 7**: The heart assist system according to claim 1, wherein the volume displacement member comprises an inflatable balloon. **Claim 8**: The heart assist system according to claim 1, wherein the volume displacement member is configured to displace a volume of blood Vd from the pumping chamber during each cycle, and the control unit is configured to periodically move the volume displacement member at a frequency F and produce a blood flow through the outflow nozzle at an exit flow rate R, where R > F · Vd. **Claim 9**: The cardiac assist system according to any one of claims 1 to 8, further comprising at least one one-way inflow valve associated with the at least one inflow opening. **Claim 10**: A cardiac assist system, an expandable cup having a transport state and an operating state, the expandable cup comprising a pumping chamber and at least one inflow opening, the expandable cup comprising a first support structure coated by or embedded within a polymer membrane; an expandable cup, an outflow nozzle coupled to the expandable cup and in communication with the pumping chamber, the outflow nozzle comprising a second support structure, the outflow nozzle having a length and shape selected such that an outlet of the outflow nozzle is downstream of a heart valve adjacent to the ventricle when at least a portion of the expandable cup is positioned within a ventricle of the heart; an outflow nozzle, a volume displacement member positioned inside the expandable cup and movable between a low volume state and a high volume state, a catheter assembly connected to the volume displacement member during operation, a control unit connected to the catheter assembly, comprising, the control unit being configured to periodically move the volume displacement member between the low volume state and the high volume state at a frequency of at least 600 beats per minute (bpm); a cardiac assist system. **Claim 11**: The cardiac assist system according to claim 10, wherein the second support structure is a monolithic extension of the first support structure and is coated by or embedded within an extension of the polymer membrane. **Claim 12**: The cardiac assist system according to claim 11, wherein the extension of the polymer membrane forms the outflow nozzle. **Claim 13**: The cardiac assist system according to claim 11, wherein the outflow nozzle comprises a monolithic extension of the polymer membrane. **Claim 14**: The cardiac assist system according to claim 10, wherein the control unit is configured to periodically move the volume displacement member between the low volume state and the high volume state at a frequency of 600 bpm to 5,000 bpm. **Claim 15**: The cardiac assist system according to any one of claims 10 to 14, further comprising at least one one-way inflow valve associated with the at least one inflow opening. **Claim 16**: A heart assist device, comprising: A flexible catheter shaft having a proximal end and a distal end; An expandable and collapsible cup coupled to the distal end of the catheter shaft and having a pumping chamber; An inflow section and an outflow section, each in communication with the pumping chamber, the inflow section having an inlet and the outflow section having an outlet, the inflow and outflow sections being sized and shaped such that when the inlet is within a heart chamber upstream of a heart valve, the outlet is positioned within a blood vessel lumen downstream of the heart valve; A one-way valve configured to open to allow blood flow into the pumping chamber through the inlet and further configured to close to block blood flow out of the pumping chamber through the inlet; A pumping mechanism within the cup, the pumping mechanism operating cyclically between an inflow phase and an ejection phase and configured to displace blood from the pumping chamber through the outlet, the one-way valve being open during the inflow phase and closed during at least a portion of the ejection phase; And The pumping mechanism is configured to operate cyclically between the inflow phase and the ejection phase at a frequency of at least 100 cycles per minute. A heart assist device. **Claim 17**: The heart assist device according to claim 16, wherein the cup comprises a support structure embedded within or covered by the membrane. **Claim 18**: The heart assist device according to claim 16, wherein the support structure comprises a metal mesh or a stent-like structure. **Claim 19**: The heart assist device according to claim 16, wherein the inflow section comprises a portion of the cup and the inlet comprises an opening or openings within the cup. **Claim 20**: The heart assist device according to claim 16, wherein the pumping mechanism comprises an expandable member movable between an expanded configuration and a collapsed configuration. **Claim 21**: The heart assist device according to claim 16, wherein the pumping mechanism is configured to increase the velocity of the blood exiting the pumping mechanism so as to entrain blood in its wake. **Claim 22**: The heart assist device according to any one of claims 16 to 21, wherein the pumping mechanism is configured to operate at a frequency of at least 600 cycles per minute. **Claim 23** A heart assist device, comprising: a flexible catheter shaft having a proximal end and a distal end; an expandable and collapsible cup coupled to the distal end of the catheter shaft and having a pumping chamber; an inflow section and an outflow section, each in communication with the pumping chamber, the inflow section having an inlet and the outflow section having an outlet, the inflow and outflow sections being sized and shaped such that when the inlet is positioned within the left ventricle, the outlet is positioned within the vascular lumen downstream of the aortic valve; a one-way valve coupled to the cup, the one-way valve being configured to open to allow blood flow into the pumping chamber through the inflow section and further configured to close to block blood flow out of the pumping chamber through the inflow section; a pumping mechanism within the pumping chamber, the pumping mechanism being configured to operate cyclically between an inflow phase in which blood flows into the pumping chamber through the inlet and an ejection phase in which blood flows out of the pumping chamber through the outlet; and the pumping mechanism is configured to operate cyclically between the inflow phase and the ejection phase at a frequency of at least 100 cycles per minute. A heart assist device. **Claim 24** The heart assist device according to claim 23, wherein the support structure comprises a metal mesh or scaffold embedded within or covered by the membrane. **Claim 25** The heart assist device according to claim 23, wherein the expandable member comprises an expandable balloon. **Claim 26** The heart assist device according to claim 23, wherein the pumping mechanism is configured to operate at a frequency of at least 600 cycles per minute. **Claim 27** The heart assist device according to claim 26, wherein the one-way valve opens during the inflow phase and closes during at least a portion of the ejection phase. **Claim 28** The volume of blood Vd is displaced from the pumping chamber during each ejection phase, and the pumping mechanism is configured to increase the velocity of the blood such that the flow rate of the blood exiting the pumping chamber exceeds the product of the frequency and the volume Vd. The heart assist device according to any one of claims 23 to 27.

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