CROSS-REFERENCE TO RELATED APPLCIATIONS

This application claims the benefit of and priority to U.S. Provisional Application No. 63/065,876 filed August 14, 2020, which is hereby incorporated by reference in its entirety.

FIELD

The present invention relates to the field of devices providing pulsatile fluid movement or circulation.

BACKGROUND

Heart failure, the final common pathway of all forms of heart disease, has a high prevalence domestically and globally, and is on the rise [1]. Mechanical circulatory support (MCS) devices, such as ventricular assist devices (VADs) and the total artificial heart (TAH), [2, 3, 4, 5], have emerged as approved frontline therapy, providing restoration of failing circulation, as either a bridge to transplantation, or an alternative to transplantation as destination therapy. Mechanical circulatory support systems, such as ventricular assist devices, are the last option for the preservation of life in cases of manifested cardiac insufficiency where a patient “needs a heart pump”. Heart assistance systems take over a part, or all, of the pumping work and thereby stabilize circulation until a donor organ is available.

Initial MCS designs largely relied on pulsatile membrane-based displacement systems, for both VAD and TAH design. These systems required a drive mechanism which was initially pneumatic with cumbersome hoses and air supplies. Analogous to a real human heart to produce pulsatile physiologic flow to the circulatory system, the first generation of VADs (e.g. Berlin Heart EXCOR, Thoratec HeartMate, Abiomed AB 5000) and the only FDA approved TAH, SynCardia TAH [20], adopts fixed volume displacement fashion that incorporates sacs, diaphragms or pusher plates actuated pneumatically, electrically or mechanically. Blood enters and is pulled into a flexible chamber from left or right ventricle and pushed out into the ascending aorta or pulmonary artery while uni-directional blood now is guaranteed by prosthetic valves. Although they could generate pulsatile physiologic blood flow, these pumps 1) inherently are bulky which makes them difficult to fit into many patients [21], 2) have large blood contacting surfaces which require frequent anti-coagulation/antithrombotic use including: warfarin, aspirin and dipyridamole to maintain high international standard ratio (INR) [22], and have potential mechanical failures caused by flexible membranes or diaphragms fatigue [23].

To circumvent the disadvantages mentioned above, second generation VADs were developed (e.g. Thoratec HeartMate2, Reliant Heart HeartAssist5, Berlin Heart INCOR), which are small profile hydrodynamic blood pump that generate continuous flow instead of pulsatile flow. These second generation VADs normally include fast spinning impellers (5000- 10000 RPM), flow straighteners, and diffusers. Due to their working principle, the flowrate depends on the pressure difference across the VAD, which requires precise sensors and cardiovascular models and control algorithms to generate desired blood flows. However, the dramatically high velocity at the impeller edge, and via other geometric features of these devices contributes high shear stress to blood, inducing hemolysis and platelet activation [23]. Also, thrombus may form in regions of recirculation or stagnation, such as the stationary flow straightener [23].

The third generation of VADs (e.g. Thoratec HeartMate3, HeartWare HVAD) and some TAHs (e.g. BiCACOR TAH, Cleveland SmartHeart TAH), utilizing centrifugal pumping architecture, use lower pump speeds (-5000 RPM) due to higher hydraulic efficiency. Generally, blood enters into the rotor and is driven outward centrifugally to the aorta or pulmonary artery without the need for flow straighteners at the inlet or diffusers at the outlet, thus lowering the probability to induce hemolysis and platelet activation. Still, noticeably high shear stresses, thus leading to blood damage, are generated inside these pumps [15, 16, 17].

Compared with normal rotary VADs, rotary piston pumps generate pulsatile flow under a dramatically reduced motor speed, which theoretically results in reduced complications and lowered shear stress. A wankel-like VAD [14], a wobbling disk VAD [25], a spherical gerotor TAH [26], a spherical rolling disk TAH [27], and the commercially available Torvad VAD are examples of rotary piston pumps. Despite, the capability of these designs to generate a pulse, they suffer from the same issues as the above detailed rotary continuous flow and centrifugal VADs in that they impart high shear stress, turbulence, sound, heat and other activation forces to blood; and contain areas of stagnancy as well - all of which drive thrombosis.

There remains a need for devices capable of moving critical fluids, biological fluids, including blood, in pulsatile flow and with reduced shear stress and effects.

It is an object of the invention to provide improved pumps and systems for pulsatile fluid flow.

It is another object of the invention to provide improved methods for achieving pulsatile fluid flow, particularly with biological fluids.

SUMMARY

Described are pumps, devices, and systems structured for moving fluids with pulsatile flow. The pumps lack valves and pistons. Generally, the pumps provide pulsatile flow through the gentle rotation of a single, one- piece, sliding vane. The vane is located inside a slot of a rotary cylinder positioned in a hollow central opening of a housing. Generally, a gap is positioned between a majority of the surface area of an end of the vane and the walls of the housing to ensure an opening is located between the vane and the walls. The movement of the vane pushes the fluid in a pulsatile fluid flow through the pump with minimal shear stress on the fluid.

Typically, the valveless pumps for pulsatile fluid flow include: (i) a housing with a central circular opening defined by a peripheral wall; and (ii) a single vane within the central opening. Each of a fluid inlet and a fluid outlet are in fluid communication with the central opening. The vane is typically configured to rotate about an axis offset from a central axis of the central circular opening without substantially contacting the peripheral walls of the housing. Typically, the vane is a single, unitary piece with a first end and a second end located at opposite sides of the vane. A gap separates at least a portion and generally a majority of the surface area of each of the vane ends from the peripheral wall. The gap may be about 10 μm to 5 mm long, as measured from the portion of the proximal vane end that is the greatest distance from the peripheral wall along a horizontal line to the peripheral wall.

A motor can be operably connected to the vane. The motor may be an electric motor, an electromagnetic motor, a passive magnetic motor, an active magnetic motor, a hydraulic motor, acoustic motor, or any other propulsive means utilized to drive the motor, operably connected to the vane. The motor is typically operably connected to a rotary cylinder. The rotary cylinder includes a slot configured to slidably receive the vane in the slot.

Also described are devices and systems containing the pumps. The systems include the pumps and an inlet port and an outlet port. Generally the inlet port is reversably connected to and in fluid communication with a fluid supply tube, and the outlet port is reversably connected to and in fluid communication with a fluid exit tube.

The pumps, devices, and systems are configured to pump at rotation speeds between about 10 and 500 rotations per minute with an output between about 0.1 LPM (liters per minute) and about 12 LPM and a pressure difference between about 1 mmHg and 220 mmHg. During pumping, as the vane rotates in the clockwise or counterclockwise direction, the flowrate varies, such as in the range of about 1 LPM to about 4 LPM (clockwise rotation) or in the range of about -1 LPM to about 4 LPM (counter clockwise rotation).

The pumps, devices and systems are typically configured to provide constant fluid displacement volume in the range between about 5 mL per revolution and about 70 mL per revolution at variable flow rates and variable fluid pressure. The system may be a closed system configured to provide circulatory fluid flow.

The system may include two or more pumps, where each pump is configured to provide a fluid displacement at variable flow rate and variable fluid pressure.

Also described are methods of operating the pumps and systems.

BRIEF DESCRIPTION OF THE DRAWINGS

Figures 1A-1D are diagrams showing the working principle of the single vane pump. Left arrow: inlet port, right arrow: outlet port. Diagonal shaded area: low pressure, small checkered area: high pressure, large checkered area: middle pressure.

Figures 2A and 2B are diagrams showing the displacement calculation (Figure 2A) and the instantaneous flow rate calculation (Figure 2B).

Figure 3 is a graph showing change in flow rate ([m ³ /sec]) as a relationship of shaft angle ([rad]). Instantaneous flowrate and an inherent pulsatile flow at a constant rotary speed is shown.

Figures 4A-4E are diagrams showing an exemplary pump 100 and its components. Figure 4A is a side perspective view of pump 100 showing motor 90 above motor case 80 contacting the housing 70 and port 60 with inlet port 62 and outlet port 64. Figure 4B is a front perspective view of pump 100. Figure 4C is an exploded view of the pump. The pump 100 includes motor 90, motor case 80, and housing 70 formed of upper case 40, lower case 30, and flow straightener 50. The housing 70 includes a central opening 42 and peripheral walls 44. A rotary cylinder 20 with vane 10 are housed in the central opening 42. The rotary cylinder 20 and the vane 10 are operably connected to the motor via shaft 92. The housing 70 and contacts the port 60 with inlet port 62 and outlet port 64 through flow straightener 50. Figure 4D is a top elevation view of the pump housing and the inlet and outlet ports. Figure 4E is a perspective view of the rotary cylinder 20 with vane 10 positioned in a slot 22 of the rotary cylinder. The rotary cylinder 20 also includes a contact opening 24 passing through a portion of the rotary cylinder 20 for operably connecting with motor 90 via shaft 92.

Figure 5 is a diagram of the system used to test fluid flow. Flow direction: clockwise: pump 1, flow meter 2, pressure sensor 3 at outlet, pressure sensor 4 at inlet, reservoir 5, motor controller 6, and data acquisition (DAQ) board 7.

Figures 6 is a graph of flowrate (LPM) versus Pressure (mmHg) showing HQ curves at five different pump speeds (300 RPM, 250 RPM, 200 RPM, 150 RPM, and 100 RPM). The open boxes are data from experimental measurements. Dashed lines are linear data fitting curves.

Figures 7A-7F are diagrams showing geometric details of the pump that result in leakage pathways. Figure 7A - gap between the rolling cylinder and flow straightener's flow separator (see box B), Figure 7B - because the sliding vane has a thickness and rounded ends, the locations where the two ends of the vane touch the wall vary along the rounded end as the vane rotates. The length of the sliding vane is shorter than the design value d ₁ +d ₂ Figure 7C - and this results in a gap between the vane and the housing wall as the vane is near θ=90º position. Figure 7D shows the vane at vertical position ( θ=0° position). Figures 7E and 7F show the different chambers 41a, 41b, and 41c in the central opening 42, which move based on the location of the rotating vane 10.

Figure 8 is a graph of flow rate ([LPM]) versus speed ([RPM]) and a comparison between theoretical (solid line) and experimental displacements of the exemplary pump. Experimental data (open boxes) and linear data fitting curve (dashed line).

Figures 9A and 9B are graphs showing theoretical flowrates versus experimental flowrates. Top: w = 101.9[RPM], P = 26.34[mmHgJ; bottom: w = 101.5[RPM], P = 98.45[mmHg]. Dashed horizontal lines are average flowrates.

Figures 10A and 10B are graphs showing theoretical flowrates versus experimental flowrates. Top: w = 196.2[RPM], P = 27.17 [mmH g]; bottom: w = 195.6[RPM], P = 105.5 [mmHg]. Dashed horizontal lines are average flowrates.

Figures 11A and 11B are graphs showing theoretical flowrates versus experimental flowrates. Top: bottom: w = 294.1[RPM], P = 121.7[mmHg]. Dashed horizontal lines are average flowrates.

Figures 12A-12D are front elevation views of single unitary vanes. The gap defined by the maximum distance horizontal between the peripheral wall and the proximal vane end, its length is identified as X. Figure 12D is a diagram of a vane 10 with wipers 16.

Figure 13 is a diagram of a portion of an exemplary system 200 containing two pumps 100 stacked together. The pumps 100 may share the same motor or have separate motors (not shown in figure).

DETAILED DESCRIPTION

I. Pumps

The pumps described herein address fluid mechanical issues, which contribute to observed adverse events limiting overall pump and system safety and effectiveness. Specifically, the pumps include a sliding vane, driven electrically and/or magnetically with minimal or no contact between the vane and pump walls to provide fluid movement with reduced shear stress.

The pumps impart a lower shear stress to fluids by one or more of the following characteristics: reducing pump speed, eliminating the use of prosthetic valves, and/or including a gap between the vane and pump walls. The pumps provide direct control of desired fluid flow. The pumps are compact. The pumps are particularly useful for medical devices that are implanted in a patient, as the compact pump structure and few operating elements reduces risk of infection in a patient.

A. Valveless pumps

Described are pumps for moving fluids with pulsatile flow. The pumps lack valves and pistons and provide pulsatile flow through the rotation of a single, one-piece, sliding vane. Pulsatile now is achieved even at constant rotational speeds.

1. Components

The valveless pumps typically include a housing with a central opening and a single, one-piece sliding vane in the central opening. The housing typically includes a peripheral wall enclosing the central opening. The vane may be positioned within a rotary cylinder positioned in the central opening and containing a slot for slidably receiving the vane. The housing typically includes an inlet and an outlet in fluidic communication with the central opening. A motor is operably connected to the vane and configured to regulate the vane rotation speed.

The structure and operability of the pumps provide gentle pulsatile flow and laminar flow to fluids minimizing shear stress on fluids. Typically, the pumps operate at rotation speeds between about 10 and 500 rotations per minute, optionally in a range from 10 to 400 rotations per minute, from 50 to 400 rotations per minute, or from 100 to 300 rotations per minute. The pumps provide pulsatile flow at flow rates between about 0.1 LPM and about 12 LPM, optionally in a range from about 0.1 LPM to about 10 LPM, from about 1 LPM to about 5 LPM, from about 1 LPM to about 4 LPM, from about 1 LPM to about 3 LPM, or from 1.5 LPM to about 3 LPM with no more than 500 rotations per minute (RPM), optionally no more than 400 RPM or no more than 300 RPM of rotor speed.

An exemplary pump includes the housing 70, the rotary cylinder 20 and the vane 10 operably connected to the motor 90. Figure 4A is a side perspective view of an exemplary pump 100 showing motor 90 above motor case 80 contacting the housing 70 and port 60 with inlet port 62 and outlet port 64. The pump can have a length (L), a width (W), and a height (H). Each of the dimensions can be adjusted, as needed, for specific applications. By way of example, exemplary pump 100 has a length (L) of about 90 mm, a width (W) of about 60 mm, and a height (H) of about 70 mm. Figure 4B is a front perspective view of pump 100. Figure 4C is an exploded view of the exemplary pump. The pump 100 includes motor 90, motor case 80, and housing 70 formed of upper case 40, lower case 30, and flow straightener 50. The housing 70 includes a central opening 42 and peripheral walls 44. A rotary cylinder 20 with vane 10 are housed in the central opening 42. The rotary cylinder 20 and the vane 10 are operably connected to the motor via shaft 92. The housing 70 and contacts the port 60 with inlet port 62 and outlet port 64 through flow straightener 50. Figure 4D is a top elevation view of the device housing components and the port. Figure 4E is a perspective view of the rotary cylinder 20 with vane 10 positioned in a slot 22 of the rotary cylinder. The rotary cylinder 20 also includes a contact opening 24 passing through a portion of the rotary cylinder 20 for operably connecting with motor 90 via shaft 92. a. Housing with Central Opening

The housing may be a once-piece hollow structure with a central opening. The housing may be formed by assembling two or more components, such as an upper case 40, a lower case 30, and a flow straightener 50, to form a hollow structure with a central opening. Typically, the central opening is surrounded by peripheral walls 44 of the housing 70.

The housing typically includes at least one inlet 46 and at least one outlet 48, although more than one inlet and/or more than one outlet may be formed for letting the fluid in and out of the pump. The inlet and the outlet are openings in the peripheral wall to provide fluid communication between the inlet and the central opening and between the outlet and the central opening. The designations inlet and outlet are functional in nature, where the inlet allows fluid to flow into the pump and the outlet allows fluid to flow out of the pump. However, whether an opening in the peripheral wall functions as an inlet or an outlet depends on the direction that the rotary cylinder is rotating.

The housing may include additional one or more openings for accommodating connections between the motor and the rotary cylinder. The connections may be wire connections.

The housing may be coated with a sealant on its interior surface to further reduce stress on fluid and reduce shear stress. The central opening of the housing typically has a height between about 10 and 220 mm, such as between about 10 mm and about 100 mm, between about 10 mm and about 25 mm, or about 18 mm. The central opening may be circular in cross-section and have a radius (R in Figure 2A) between about 10 mm and about 50 mm, between about 10 mm and about 25 mm, or about 18 mm. These dimensions are particularly relevant for implantable medical devices. b. Rotary cylinder

The rotary cylinder is typically positioned within the central opening of the housing. The central axis of the rotary cylinder is offset from the central axis of the central opening. The rotary cylinder generally includes a contact opening and a slot, such as a through hole.

The contact opening is configured to operably connect the rotary cylinder to a motor. The contact opening typically mates with the motor shaft of the motor. The contact opening may be an indentation, or a shaped hole, such as a D-shaped hole, a star-shaped, oval-shaped, tooth-shaped hole, which is configured to mate with a respectively- shaped motor shaft. The contact opening is typically positioned on or about the central axis of the rotary cylinder, does not contact the slot. The contact opening is separate from the slot.

The slot, such as a through hole, is positioned perpendicular to the central axis of the rotary cylinder and runs from one side of the peripheral wall to the opposite side of the peripheral wall of the cylinder and through the central axis of the rotary cylinder. The slot typically has a shape that conforms to the shape of the vane, while allowing the vane to slide within the slot.

Typically, the rotary cylinder has a radius (r in Figure 2A) between about 5 mm and about 30 mm, such as between about 5 mm and about 15 mm, or about 10 mm. The central axis of the rotary cylinder is offset from the central axis of the central opening by a value e (eccentricity, e in Figure 2A). Typically, e is a value between about 2 mm and about 45 mm, between about 2 mm and about 20 mm, or about 8 mm. These dimensions are particularly relevant for implantable medical devices.

A rotary cylinder 20 with a contact opening 24 and slot 22, with a single, one-piece sliding vane 10 shown therein, is depicted in Figure 4E. c. Vane

Typically, the pump contains a single vane. The vane is a one-piece vane, positioned in and in slidable relation with the through-hole of the rotary cylinder. The vane typically has two opposing ends positioned proximal to the peripheral wall of the housing, and two opposing lateral surfaces. The opposing ends may have different geometries along a longitudinal plane, but generally are rounded at the edges along a transverse plane.

The lateral surfaces are typically smooth and slide through the through-hole contacting the inner walls of the through-hole.

The lateral surfaces may include a raised section proximal to the ends of the vane to act as stoppers. The raised sections may be in any shape that protrudes from the lateral surfaces. The raised sections may be positioned proximal to the ends of the vane to prevent the vane from sliding out of the through hole. The presence of the raised sections is particularly useful in those in embodiments where the ends of the vane do not contact the peripheral wall of the housing.

Typically, the vane has a length (d ₁ +d ₂ of Figurse 2A and 2B) between about 10 mm and about 60 mm, between about 20 mm and about 40 mm, such as about 36 mm. These dimensions are particularly relevant for implantable medical devices. However, the vane may have any suitable length and shape selected to provide a desired level of leakage between chambers of the housing. For example, the vane shape and length may be selected to provide more or less leakage between housing chambers. As noted below, the vanes may be selected to have a gap with a distance 1 from about 10 μm to 5 mm, optionally from 10 μm to 100 μm, from 100 μm to 1 mm, or from 1 mm to 5 mm. In some embodiments, the vane may be configured to have an adjustable length within the housing. The length of the vane may be adjusted with springs, screws, stoppers, or gears. In these embodiments, the vane is not a single, one-piece vane, and may include two or more pieces.

The vane is generally operably connected to the motor via the rotary cylinder.

Minimal Contact with the Wall

The vane can have a variety of different geometries. Exemplary different geometries to the vane 10 are presented in Figures 4E and 12A- 12D. Typically, the vane makes minimal contact with the peripheral wall of the housing when in a resting position and in use, operates without substantial contact with the peripheral wall.

The gap between the peripheral wall 44 and the proximal vane end 12 is shown as region 18. The maximum distance measured along a horizontal line between the peripheral wall 44 and the proximal vane end 12 is shown by the distance λ (see, e.g. FIGs. 12A-12C). The distance 1 may be from about 10 μm to 5 mm, optionally from 10 μm to 100μm, from 100 μm to 1μm, or from 1 mm to 5mm, long.

The minimal contact surface is typically equivalent to between about 0.01% and 5% of vane's surface area, such as from about 0.01% to 1%, from about 0.01% to 0.5%, or from about 0.01% to 0.1%. Thus, when rotating in use, the vane operates without substantial contact with the peripheral wall. The minimal contact surface between the vane and the peripheral wall is sufficient to guide the vane as it rotates, such as at rotation speeds between about 10 and 500 rotations per minute, optionally in a range from 10 to 400 rotations per minute, from 50 to 400 rotations per minute, or from 100 to 300 rotations per minute.

Figure 7D shows the different chambers 41a, 41b, and 41c in the central opening 42 defined by the location of the vane 10 and the peripheral wall 44. The fluid in chamber 41a is able to mix with fluid in chambers 41b and 41c due to the gaps between the vane ends 12 and 14 and the peripheral wall 44 as well as due to the gap between the rotary cylinder 20 and the peripheral wall 44 showing in Figure 7B.

Wipers

In some embodiments, the vane includes one or more wipers at each of its ends. In these embodiments, the one or more wipers contact the peripheral wall of the housing, while allowing fluid to flow through the gap between the end of each vane and the peripheral wall. The wipers may be of any elongated shape. The wipers are typically flexible and able to bend and move along the wall as the vane rotates.

An exemplary vane 10 with wipers 16 is shown in Figure 12D. d. Inlet and Outlet

The pump includes an inlet and an outlet in fluid communication with the central opening. The inlet and the outlet may be openings in the flow straightener that are connected to the central opening. Figure 7C shows the inlet 46 and the outlet 48 when the vane rotates clockwise. The designations inlet and outlet are functional. The direction of pump flow may be reversible. For example, if the direction of the vane rotation reverses and the vane rotates in a counterclockwise direction, the inlet may function as an outlet, and the outlet may function as an inlet.

The inlet and the outlet may be of any suitable shape and dimension to provide desired volume of fluid flow into and out of the pump, respectively. The dimensions of the inlet and the outlet may vary to provide a pressure difference for the fluid pressure at the inlet and at the outlet. The typical pressure differences achieved at the inlet and the outlet are between about 1 mmHg and about 220 mmHg, such as between about 5 mmHg and 200 mmHg, between about 5 mmHg and about 50 mmHg, between about 50 mmHg and about 200 mmHg, or about 5 mmHg, about 10 mmHg, about 15 mmHg, about 20 mmHg, about 60 mmHg, about 90 mmHg, about 100 mmHg, about 120 mmHg, about 150 mmHg, about 200 mmHg, or about 220 mmHg. For example, the pump may generate a pressure head between 60 mmHg and 140 mmHg at 6LPM or between 10 mmHg and 40 mmHg at 6LPM.

Typically, the inner diameter of the inlet and the outlet are in the range between about 5 mm and about 20 mm, between about 10 mm and about 15 mm, such as about 12.7 mm. The cross-sectional area of the inlet and the outlet is typically in the range between about 18 mm ² and about 320 mm ² , such as about 75 mm ² and about 180 mm ² , such as about 126.68 mm ² . These dimensions are particularly relevant for implantable medical devices.

The pump may also include a flow straightener positioned about the inlet and outlet and forming a portion of the housing. A port may enclose the flow straightener and contact the housing. The port may include an inlet port in fluid communication with the pump inlet and an outlet port in fluid communication with the pump outlet. e. Motor

In some embodiments, the pump includes a motor positioned proximal to the housing. The motor is operably connected to the rotary cylinder. The term “operably connected to the rotary cylinder” refers to a direct connection or an indirect connection between the motor and the rotary cylinder. A direct connection may include a motor shaft configured to mate with the contact opening of the rotary cylinder. The indirect connection may include the motor contacting one or more elements other than the rotary cylinder, such that the one or more elements transfer the motor-generated rotary motion to the rotary cylinder.

The motor, which is operably connected to the rotary cylinder, is thus also operably connected to the vane. The rotation of the rotary cylinder controls the rotation of the vane.

The motor may be an electric motor, an electromagnetic motor, a passive magnetic motor, or an active magnetic motor, hydraulic motor, acoustic motor, or involve other propulsive means for generating motion, operably connected to the vane. The motor may include magnetic bearings and receive electrical signals for controlling the magnetic bearings. Additional magnetic bearing(s) may be positioned on rotary cylinder. The motor may control axial force of a magnetic axial bearing of the rotor.

The motor may be in feedback connection with one or more sensors.

Sensors for various physiological variables, such as fluid temperature and pressure, may be arranged in the region of the pump. Also, the motor may be in feedback connection with a sensor for the acceleration measurement, which gives information with regard to a movement of the patient, and a sensor that detects the angular position of the rotor in the pump.

The rotary position sensor may be useful with use of a magnetic bearing in order to determine the axial rotor position. This may form an input variable for the control of the bearing and thus the magnetic mounting of the rotor.

2. Scalability

The dimensions of pump components may vary based on pump’s use and purpose. Pumps designed for pulsatile flow of biological fluids may have dimensions given herein. Pumps designed for pulsatile of non-biological may have dimensions of elements scaled up to accommodate their use in industrial settings. For example, pump and component dimensions as described may be scaled up by a factor of 1.5, 2, 3, 4, 5, 6, 7, 8, 9, 10 or more to accommodate pulsatile flow of fluids in food and beverage industry.

For example, pumps with output of 60 LPM (instead of 6 LPM) with a gentle pulsatile flow to fluids may be formed by scaling the pump and its component dimensions by a factor of 10.

3. Function

Typically, the pumps are useful for moving fluids by pulsatile flow. The pumps are gentle on fluids, minimizing fluid shear stress. Typically, the pumps operate at a rotation speeds between about 10 and 500 rotations per minute, optionally in a range from 10 to 400 rotations per minute, from 50 to 400 rotations per minute, or from 100 to 300 rotations per minute, and flow rate between about 0.1 LPM and about 12 LPM, optionally in a range from about 0.1 LPM to about 10 LPM, from about 1 LPM to about 5 LPM, from about 1 LPM to about 4 LPM, from about 1 LPM to about 3 LPM, or from 1.5 LPM to about 3 LPM. The pumps can achieve a pressure different at the inlet and the outlet of between about 1 mmHg and 200 mmHg, optionally between about 5 mmHg and 200 mmHg, between about 5 mmHg and about 50 mmHg, or between about 50 mmHg and about 200 mmHg. Typically, the pumps provide fluid displacement at variable flow rates and variable fluid pressure. The fluid displacement volumes of the pumps range between about 5 mL per revolution and about 70 mL per revolution, optionally from about 5 mL per revolution to about 50 mL per revolution, from about 10 mL per revolution to about 50 mL per revolution, or from about 10 mL per revolution to about 30 mL per revolution.

Exemplary fluids include incompressible fluids, fluids that have a density and viscosity of biological fluids, colloids, crystalloids, or emulsions. Preferably, the fluid is blood.

The fluid shear may be measured by methods known in the art. The methods include mathematical modeling and computational fluid dynamics (CFD) simulations, which may be used to predict fluid shear stress by utilizing digitally modeled fluid domain geometry, flow rate and viscosity parameters. The shear stress may be expressed as force per unit area of the pump’s internal surface (e.g. Pa, dynes/cm ² ). The shear stress in pump may be compared to shear stress in tubing with smooth surfaces, and the difference expressed in percentage increase.

In embodiments where the fluid to be pumped is blood, in vitro hemolysis tests and/or shear-mediated platelet activation (SMPA) assay may be used to assess the level of fluid shear (Berk et al., Artificial Organs, 43:870-879 (2019); Li et al., Artificial Organs, 44:E226-E237 (2020)). Assessing SMPA by platelet activity state (PAS) assay can utilize gel-filtered platelets (GFP) isolated by filtration of platelet-rich plasma. GFP is typically diluted and passed through the pump at a desired flow rate and samples are taken at timed intervals. A prothrombinase-based platelet activation state assay may be used to measure platelet activation in the sample(s). The value of platelet activation in the test sample(s) may be compared to the value of platelet activation obtained from a negative control sample (an undisturbed sample), as well as to the value of platelet activation obtained from a positive control sample (fully sheared sample).

Shear-mediated platelet activation may be measured utilizing phosphatidylserine externalization [PSE] and Annex V binding, or other means as detailed by Roka-Moiia et al., identifying shear vs. other means of activation (Roka-Moiia et al., Thrombosis and Haemostasis, 120:776-792 (2020)).

Another method that can be used to assess the level of fluid shear includes measuring plasma-free hemoglobin (PFH) and obtaining Normalized Index of Hemolysis (NIH) in blood samples obtained at time intervals from blood passed through the pump (Berk et al., Artificial Organs, 43:870-879 (2019)).

Typically, the pumps and systems pump the fluid in a manner that achieves no more than 15% increase in fluid shear over fluid shear observed for a negative control, or no greater increase in fluid shear than 15% increase over fluid shear at baseline (a measurement obtained without pumping initiation). Preferably, the fluid shear obtained with the disclosed pumps is between about 0.1% and about 15%, between about 0.1% and about 10%, or between about 0.1% and about 5%, greater than fluid shear observed at baseline. Expressed in terms of shear stress, overall shear stress by the pump may be configured to not induce activation of the traversing fluid. For example, when fluid is blood, the shear stress is no more than optimal of 0 - 30 dynes/cm ² , and typically is less than 70 dynes/cm ² . Further, the design and control may also modulate and control the level of shear stressaccumulation. For example, accumulation of shear stress over time (shear stress x time), can be controlled to achieve non-activating passage, pumping and traverse of fluid.

B. Pumps and Devices

1. Mechanical circulatory support (MCS) devices

Mechanical circulatory support (MCS) devices, i.e. ventricular assist devices (VADs), ventricular replacement devices, and total artificial hearts (TAHs), while effective and vital in restoring hemodynamics in patients with circulatory compromise in advanced heart failure, remain limited by significant adverse thrombotic, embolic, and/or bleeding events. Many of these complications are due to chronic exposure, via the existing MCS devices, to non-pulsatile flow and high shear stress created by the methods of blood propulsion or use of prosthetic valves.

The valveless pumps may be incorporated into devices for circulatory fluid flow, and may be used as blood flow assist devices. The valveless pumps provide lower shear stress imparted to blood by reduced pump operating speed compared to first, second, and third generation VADs, while achieving the same output. The valveless pumps do not utilize prosthetic valves, thus diminishing the generation of shear stress compared to the shear stress typically observed with first, second, and third generation VADs. The valveless pumps allow direct flowrate control to generate desired blood flowrate and pulsatile flow profile.

The pumps described herein can be sized to fit into human adult or pediatric patients. The prototype described in the Examples, having the dimensions listed in Table 1, can be modified to be more compact, such by as having pumps with a thinner wall (e.g., from 5 mm to 3mm) and a shallower hole (e.g., from 5mm to 3mm). The motor could be modified to have with a longer shaft, and the can counterbore-like hole of the upper case could be shortened, such as from 10 mm to 5 mm.

The pumps may be connected to any closed loop or open fluid flow circuits to provide pulsatile fluid flow with reduced fluid shear.

Exemplary closed loop circulatory fluid flow circuits with pulsatile flow include cardiovascular systems, extracorporeal membrane oxygenation system, cardiopulmonary bypass system, or hemodialysis systems. One or more pumps may be connected to patient’ s circulatory system and function as VADs or TAHs. One or more pumps may be connected to an extracorporeal membrane oxygenation system. One or more pumps may be connected to a cardiopulmonary bypass system. One or more pumps may be connected to a hemodialysis system to aid with fluid flow. The pumps can be incorporated into a fully implantable VAD or TAH system that minimizes the risk of infection for a patient, provides pulsatile fluid flow at reduced operational speeds than what is available for the existing VADs and TAHs, and reduced fluid shear stress.

2. Other uses for Valveless Pumps

Exemplary open fluid flow circuits with pulsatile flow include circuits in dairy milk production, in food manufacturing, such as in processes requiring a step of gentle fluid movement and/or a step of fluid fermentation, in beer and wine production, in biotechnology and fermentation applications, and in chemical manufacturing requiring movement of fragile fluids. Fragile fluids include, for example, colloidal fluids, emulsions, and fluids containing microrganisms, cells, cell aggregates, micro/nanocapsules, liposomes and microparticles and/or nanoparticles.

II. Systems

The pumps may be included in a system containing the pump and one or more flow path elements. The one or more flow path elements are additional elements configured to link the pump to a flow path and to provide a fluid flow path. The elements may include a port with an inlet port section and an outlet port section, a fluid supply tube, a fluid exit tube, flow path tubing, flow stoppers, flow splitters, and flow connectors. The connection between the pump and any one of the additional elements may be reversible or permanent.

Two or more pumps and one or more flow path elements may be combined into a system.

The system may provide pulsatile fluid flow to one or more fluid flow paths. Typically, the two or more pumps in the system are arranged in any suitable manner relative to each other. The pumps may have any suitable proximal arrangement. In this embodiment, the position of the pumps relative to each other may vary, so long as the pumps are in physical contact with one another (e.g., the pumps in Figure 13 are in physical contact with one another). The pumps may have a distal arrangement. In this embodiment, the pumps are not in physical contact with one another. In proximal arrangement or distal arrangement, the pumps in a system are generally operably linked through interconnecting flow paths and one or more motors. The motors may be programmed to operate at the same or different speeds. The speeds may range between about 10 rotations per minute and 500 rotations per minute, optionally in a range from 10 to 400 rotations per minute, from 50 to 400 rotations per minute, or from 100 to 300 rotations per minute. The pumps in the system may be configured to provide an output between about 5 mL per revolution and about 50 mL per revolution controlled by the one or more motors. The pumps in the system may be controlled to provide a time-varying instantaneous output flow rate by actively controlling the motor speed.

III. Materials for Pumps and Devices

Suitable materials for forming pump and system elements include metals and natural and synthetic polymers. The metals include grade 5 titanium, titanium alloys, nickel-titanium alloys, cobalt chromium alloys, stainless steel alloys, copper alloys, iron and/or ferrous alloys, nichrome, zinc and galvanized materials, tantalum, kanthal, or cupronickel.

The polymers typically are biocompatible, and include polydimethylsiloxane (PDMS), polysulfone (PSF), and other materials. PDMS is a versatile elastomer that is easy to mold, and PSF is a rigid, amber colored, machinable thermoplastic. Other suitable materials include biologically stable thermosetting and thermoforming polymers, including polyethylene, polypropylene, polyoxymethylene (POM) - also known as acetal, polyacetal, and polyformaldehyde, polymethylmethacrylate, polyurethane, polysulfones, polyetherimide, polyimide, ultra-high molecular weight polyethylene (UHMWPE), cross-linked UHMWPE and members of the polyaryletherketone (PAEK) family, including polyetheretherketone (PEEK), carbon-reinforced PEEK, and polyetherketoneketone (PEKK). Preferred thermosetting polymers include, but are not limited to, poly etherketoneketone (PEKK) and poly etheretherketone (PEEK).

One or more of the components of the pumps and/or systems can be formed via stereolithography, soft lithography, laser machining, micromachining, curing, bonding, three-dimensional printing, additive manufacturing, molding, micromolding, and/or coating.

IV. Methods of Using

The pumps and systems described herein can be used in a variety of different methods involving the movement of one or more fluids with a pulsatile flow profile along a flow path.

The fluids may be Newtonian or non-Newtonian fluids, and include biological fluids, critical fluids, colloids, crystalloids, emulsions, nutritional fluids, and fluids in dairy, food, pharmaceutical, biotechnology, and beverage industries. In preferred embodiments, the fluids are blood, serum, plasma, colloids, crystalloids, or nutritional fluids.

The methods typically include flowing a fluid of interest through the disclosed pumps or systems. The pumps and systems move the fluids with fluid displacement when subjected to variable flow rates and variable fluid pressures.

Typically, the methods include providing a pulsatile flow to the fluid at a flow rate between about 0.1 LPM and about 12 LPM. For example, the methods include pumping the fluid through the pump with a flowrate in the range from about 0.1 LPM to about 12 LPM, from about 1 LPM to about 4 LPM, from about 1 LPM to about 3 LPM, or from 1.5 LPM to about 3 LPM. As depicted in Figure 3, even at constant rotational speeds, the flowrate varies during the rotation of the vane. Thus, during a single rotation of the vane, the flowrate can vary from about 1 LPM to about 4 LPM, about 1 LPM to about 3 LPM, or from 1.5 LPM to about 3 LPM. Throughout rotation of the vane during the pumping method, the flowrate can vary from about 1 LPM to about 4 LPM, or about 1 LPM to about 3 LPM, or from 1.5 LPM to about 3 LPM.

The flowrates and flowrate ranges listed above correspond with the vane rotating in one direction, e.g. clockwise direction. However, if desired, the vane can reverse direction, and rotate in the counterclockwise direction,. When the flow moves in the reverse direction, e.g. due to the counterclockwise rotation of the vane, the fluid can be pumped through the pump with a flowrate in the range from about -0.1 LPM to about -12 LPM, from about -1 LPM to about -4 LPM, from about -1 LPM to about -3 LPM, or from - 1.5 LPM to about -3 LPM. Similarly, due to the variation in the flowrate during the rotation of the vane, during a single rotation of the vane in the counterclockwise direction, the flowrate can vary from about -1 LPM to about -4 LPM, about -1 LPM to about -3 LPM, or from -1.5 LPM to about -3 LPM. Throughout rotation of the vane during the pumping method for a vane rotating in the reverse (e.g. counterclockwise) direction, the flowrate can vary from about -1 LPM to about -4 LPM, or about -1 LPM to about -3 LPM, or from -1.5 LPM to about -3 LPM.

As the direction of the rotation of the vane is able to change from a first, clockwise direction to a second, counterclockwise direction, the overall flowrate for the pump can range for example from about -12 LPM to about 12 LPM, from about -8 LMP to about 8 LPM, from about -4 LMP to about 4 LPM, from about -3 LPM to about 3 LPM.

The methods include pumping the fluid through the pump with a pressure difference of between about 5 mmHg and 200 mmHg, such as with a pressure difference between about 5 mmHg and about 50 mmHg, or between about 50 mmHg and about 200 mmHg.

The methods for providing pulsatile flow of fluids provides gentle changes in fluid flow rate resulting in low levels of fluid shear stress. Low level of fluid stress includes fluid stress at between about 0 dynes/cm ² and about 70 dynes/cm ² , such as between about 0 dynes/cm ² and about 60 dynes/cm ² , between about 0 dynes/cm ² and about 50 dynes/cm ² , between about 0 dynes/cm ² and about 40 dynes/cm ² , between about 0 dynes/cm ² and about 30 dynes/cm ² .

The disclosed pumps, devices, systems, and methods can be further understood through the following numbered paragraphs.

1. A valveless pump for pulsatile fluid flow comprising:

(i) a housing with a central circular opening defined by a peripheral wall; an inlet in fluid communication with the central circular opening; and an outlet in fluid communication with the central circular opening; and

(ii) a single vane within the central circular opening; wherein the vane is configured to rotate about an axis offset from a central axis of the central circular opening without substantially contacting the peripheral wall of the housing.

2. The valveless pump of paragraph 1, wherein the vane is a one-piece vane.

3. The valveless pump of paragraph 1 or 2, wherein the vane comprises two ends and wherein the valveless pump comprises a gap between each of the two ends and the peripheral wall.

4. The valveless pump of any one of paragraphs 1-3, wherein the gap is about 10 μm to 5 mm long, optionally from about 10 μm to 100 μm, from about 100 μm to 1 mm, or from about 1 mm to 5 mm long, as measured from the proximal vane end to the peripheral wall.

5. The valveless pump of any one of paragraphs 1-4, comprising a motor operably connected to the vane.

6. The valveless pump of any one of paragraphs 1-5, comprising an electric motor, an electromagnetic motor, a passive magnetic motor, or an active magnetic motor, hydraulic motor, or acoustic motor, operably connected to the vane.

7. The valveless pump of any one of paragraphs 1-6, comprising a motor operably connected to a rotary cylinder comprising a slot wherein the vane is slidably received in the slot, wherein a central axis of the rotary cylinder is offset from the central axis of the central circular opening.

8. The valveless pump of paragraph 7, the slot that runs from one side of the peripheral wall to the opposite side of the peripheral wall of the cylinder and through the central axis of cylinder and has shape that conforms to the shape of the vane. 9. The valveless pump of any one of paragraphs 1-6, comprising a motor operably connected to a rotary cylinder.

10. The valveless pump of paragraph 9, wherein the rotary cylinder is configured to rotate about the central axis of the rotary cylinder.

11. The valveless pump of any one of paragraphs 1-10, wherein the vane has two ends and wherein each end is rounded.

12. A system comprising the pump of any one of paragraphs 1-11.

13. The system of paragraph 12, further comprising an inlet port and an outlet port, wherein the inlet port is reversably connected to and in fluid communication with a fluid supply tube and the outlet port is reversably connected to and in fluid communication with a fluid exit tube.

14. The system of paragraph 13, wherein the system is a closed system configured to provide circulatory fluid flow.

15. The system of any one of paragraphs 12-14, wherein in use the system is configured to provide constant fluid displacement volume in a range from about 5 mL to about 70 mL per revolution, optionally in a range from about 5 mL per revolution to about 50 mL per revolution, from about 10 mL per revolution to about 50 mL per revolution, or from about 10 mL per revolution to about 30 mL per revolution.

16. The system of any one of paragraphs 12-15, wherein in use the system is configured to provide fluid displacement at variable flow rates and variable fluid pressure.

17. The system of any one of paragraphs 12-16, wherein the system comprises two or more pumps of any one of paragraphs 1-11.

18. The system of paragraph 17, wherein each of the two or more pumps is configured to provide fluid displacement at variable flow rates and variable fluid pressure.

19. A method for pulsatile flow of an incompressible fluid comprising pumping the fluid through the pump of any one of paragraphs 1-11 or through the system of any one of paragraphs 12-18. 20. The method of paragraph 19, wherein the fluid has a density and viscosity of biological fluids, colloids, crystalloids, or emulsions, optionally wherein the fluid is blood.

21. The method of paragraph 19 or 20, wherein the pump or system pumps the fluid into and out of the pump at rotation speeds between 10 and 500 rotations per minute, optionally in a range from 10 to 400 rotations per minute, from 50 to 400 rotations per minute, or from 100 to 300 rotations per minute.

22. The method of any one of paragraphs 19-21, wherein the pump or system pumps the fluid through the pump at a flow rate between about -12 LPM and about 12 LPM, optionally in a range from about -8 LMP to about 8 LPM, from about -4 LMP to about 4 LPM, from about -3 LPM to about 3 LPM.

23. The method of any one of paragraphs 19-22, wherein the step of pumping the fluid through the pump or system comprises rotating the vane in the clockwise and/or counter clockwise direction within the central circular opening.

24. The method of paragraph 23, wherein during rotation of the van, the fluid flowrate varies, such as from about -12LPM to about 12 LPM, from about -8 LMP to about 8 LPM, from about -4 LPM to about 4 LPM, or from about -3 LPM to about 3 LPM.

25. The method of any one of paragraphs 19-24, wherein the system is a closed circulatory system and wherein the pump pumps the fluid in a circulatory fluid flow path at a pressure difference of between about 1 mmHg and 220 mmHg, optionally between about 10 mm and about 100 mm or between about 10 mm and about 25 mm.

26. A mechanical circulatory support device comprising a pump for pulsatile fluid flow comprising:

(i) a housing with a central circular opening defined by a peripheral wall; an inlet in fluid communication with the central opening; and an outlet in fluid communication with the central opening; and

(ii) a single vane within the central opening; wherein the vane is configured to rotate about an axis offset from a central axis of the central circular opening without substantially contacting the peripheral walls of the housing.

27. The mechanical circulatory support device of paragraph 26, wherein the device is a ventricular assist device, heart pump, ventricular replacement device, or a total artificial heart.

28. The mechanical circulatory support device of paragraph 26 or 27, wherein the device is a heart pump, and wherein the pump generates a pressure head between 60 mmHg and 140 mmHg at 6LPM or between 10 mmHg and 40 mmHg at 6LPM.

29. A method of mechanically supporting a ventricular or heart function in a subject comprising connecting a pump to the cardiovascular system of the subject, the pump comprising:

(i) a housing with a central circular opening defined by a peripheral wall; an inlet in fluid communication with the central opening; and an outlet in fluid communication with the central opening; and

(ii) a single vane within the central opening; wherein the vane is configured to rotate about an axis offset from a central axis of the central circular opening without substantially contacting the peripheral walls of the housing.

30. An extracorporeal membrane oxygenation system comprising a pump of any one of paragraphs 1-11.

31. A cardiopulmonary bypass system comprising a pump of any one of paragraphs 1-11. Examples

Example 1. Design of the Non-Compressible Single Sliding Vane MCS Pump

Non-Compressible Single Sliding Vane MCS Pump. The MCS pump, a non-compressible single sliding vane pump, consists of a rolling cylinder (rotor), a sliding vane, case(s), a flow straightener, and ports. The rotating cylinder has a through-all slot to allow the vane to slide through completely, and the vane separates the chamber into two compartments while pumping blood. In one embodiment, the wall geometry of the case (a portion of the housing) is configured to serve as a guide for one or both of the ends of the vane to slide against. The flow straightener straightens (a portion of the housing) the flow into and out of the pump and also serves as a guide for the vane to slide against.

As the vane has certain thickness, during the sliding motion, the actual contacting point is not at the tip of the end of the vane (Figure7C). Therefore, the actual length of the vane is shorter than the theoretical value obtained by Eqn. 1. When the vane is in a horizontal position (see, e.g. Figure 7C), the vane is closest to the peripheral wall. When the vane is in a vertical position (see, e.g. Figure 7D), there is a larger gap between the end of the vane and the housing wall. At other positions of the van in the housing, a gap exists between the end of the vane and the peripheral wall.

The vane, the rotary cylinder, and/or the motor guide the vane during rotation (Figures 7C-7F and 12A-12D).

Two chambers are sectioned by the sliding vane and blood is pulled in and pumped out cyclically. As depicted in Figures 1A-1D, as the vane rotates, the volume of one chamber decreases and blood is pumped out while the volume of the other chamber increases and blood is drawn in. Throughout this process, blood gently pulsates into and out of the MCS without need of prosthetic valves, the vane effectively functions as a valve would. Geometric Design. A parametric equation describes the trajectory of the sliding vane's end-point, which in turn determines the shape of the wall case. d1(θ) + d ₂ (θ ) — R + r + e (1) where d ₁ and d ₂ are distances from the two ends of the sliding vane to the center of the case respectively; θ is the rotating angle of the sliding vane, measured from the vertical position; R is pseudo-radius of the case; r is the radius of the rolling cylinder; and e is the eccentricity, as shown in Figures 2 A and 2B.

With boundary conditions [d ₁ , d ₂ ] = [R + e, r] at θ = θ, [d _1, d ₂ ] = [(R + e + r)/2 , (R + e + r)/2] at θ = π/2, and [d _1, d ₂ ] = [r, R + r] at θ = π , one smooth and continuous solution to Eqn. 1 is obtained as follows:

Porting. The chamber volume reaches a maximum when θ = TI/2 , and in order to avoid compressing the working fluid, blood, porting starts at this position. Ideally porting ends at because this position mathematically separates the low- and high-pressure sides. Actual design results in a little narrow porting size and the porting ends at θ < π because wall thickness would occupy some space.

Displacement and Instantaneous Flowrate Calculation.

Displacement volume of existing VADs is vital as a trade-off between overall pump size and pump speed always exists. It can be seen from Figures 1A- 2B that one full cycle pumps out four times the shaded volume. The planar area A is calculated as i Thus, the overall displacement volume per revolution is obtained by multiplying planar area A and the chamber height h by four. (4)

The total size of the pump can be obtained similarly: (5) (6)

Another feature of the disclosed MCS pump is inherent pulsatility - even at constant rotational speeds. The instantaneous output flowrate, Q(t), is equal to the change rate of an infinitesimal volume displacement, hdA, and infinitesimal time step, dt:

(7)

Figure 3 plots instantaneous flow rate Q(t) against input shaft angle θ at a constant speed θ = to. Figure 3 shows that the instantaneous flowrate reaches its minimum at θ = (2i + 1) π/2 and its maximum at θ (2i π ), i ∈ N. Therefore, two flow pulses are generated for every revolution (2π ) of the driving shaft.

Kinematics of Sliding Vane. The kinematics of the sliding vane are derived herein. From Eqns. 1-2, the instantaneous locations of the sliding vane's distal points, P ₁ and P ₂ , are

[P _1x (θ)P _1y (θ)P _2x (θ) P _2y (θ)] (8)

= [d ₁ (θ)sinθ d ₁ (θ)cosθ d ₂ (θ)sinθ d2(θ)cosθ ]

The instantaneous location of the center of mass P _M and its trajectory are obtained as follows: where the geometric relation Φ = 20 exists (see Figure 2B). It can be seen from Eqn. 10 that the center of mass of the sliding vane P _M forms a circular trajectory around (0, (R + e — r)/4) with radius (R + e — r)/4. The motion of the sliding vane can be described as a combination of a circular motion of the center of mass at a speed Φ = 2Φ and a rotation of the rigid body about the center of mass at a speed 0.

Design Objectives and Constraints. The first objective is to allow the pump to generate enough output flowrate, to satisfy blood demand under the majority of body conditions. Accordingly, 6 [LPM] is used here [12].

Table 1. The design specifications of a prototype.

Term Variable Value

Pseudo radius R 18 [mm]

Rolling cylinder radius r 10 [mm]

Eccentricity e 8 [mm]

Length of sliding vane d ₁ + d ₂ 36 [mm]

Height of pump chamber h 18 [mm]

Displacement volume (per V _disp 24.8 [ml] revolution)

Inner diameter of port d _port 12.7 [mm]

Cross sectional area of port A _port 126.68[mm ² ]

(each)

Example 2. Pump Construction and Experiment Setup Set up

A prototype having the specifications described in Table 1 was designed, assembled and tested. First Generation Prototype. The design of the MCS pump is illustrated in Figures 4A-4E. All blood contact parts were manufactured by computer numerical control (CNC) machine with polytetrafluoroethylene (PTFE). The lower and upper cases have counterbore-like structures to constrain the rolling cylinder to rotary motion only. The flow straightener was designed to offer a smooth guide for the vane to slide against. It also separates and prevents the flow path from inlet to outlet. Different sizes and orientations of the ports are optional to better fitting with cardiac vessels and here a straight configuration with 0.50” ID was used. Silicone sealant (GE®) was used to provide further sealing between all surfaces. A brushless DC motor (Maxon EC 45) was used as a prime driver, with motor shaft filed to a D-shape profile. A motor case was 3D printed by acrylonitrile butadiene styrene (ABS) to fasten the motor and to couple with the pump. A custom- made motor with longer shaft and thinner rotor size may further compact the overall size of the MCS pump in the future.

Experiment Setup. An experimental loop was constructed using 0.5” tubing to characterize the MCS pump performance and to demonstrate its inherent flow pulsatility, as shown in Figure 5. An industrial grade flow meter (McMaster Part No. 4352K51) was placed at the outlet to measure the flowrate. The inlet and outlet pressures were measured by two Honeywell pressure transducers (Part No. ABPDANT005PGAA5). The motor speed was measured with motor's integrated hall-effect sensor. A mixture of water and glycerol with a mass ratio of 54 to 46 was circulated in the loop to represent the density (1.12 [g/mL] ) and viscosity (3.56 [mPaS]) of blood [14]. The MCS prototype was controlled to run at several constant speeds against various pressure differences, which was achieved by a resistance clamp (not shown in the figure). The flowrate and pressure difference were chosen according to human-heart operating conditions and range from 0 to 6 [LPM] and 0 to 120 [mmHg], respectively. All electronic control schemes and data collection were implemented using MATLAB Simulink and an xPC Target machine using a PCI-6229 National Instrument DAQ board. CFD Setup. To solve the fluid- structure interaction (FSI) problem, a parallelized computational framework was employed that was previously developed for simulating biological systems that involve large deformations [28-31]. In this partitioned framework, the flow was solved by using a Cartesian grid based on a sharp-interface direct- forcing immersed-boundary method. While in the current case, the structures were considered as rigid parts and their transient kinematics were prescribed according to Kinematics of Sliding Vane section.

In the simulation, the flow domain was represented by a 79x40x19 cm ³ rectangular bounding box and was divided by a 263x126x60 uniform Cartesian grid with Δx, Ay and Az being about 3.0x10-4 cm. The density and dynamic viscosity of the blood are, p = 1.0 [g/cm3] and p = 0.005 [Pa · s], respectively. Each cycle had a time duration of T = 0.6s and the time step used in the flow solver is Δt = 5.0 x 10 — 5 s. Pressure boundary conditions were prescribed at both the inlet and outlet.

Laminar flow. When fluid is blood, the flow may be laminar flow based on the following: averagely blood density p = 1056kg/m3, blood dynamic viscosity μ = 3.5cP = 3.5e — 3PaS, the diameter of inlet/outlet d = 11e — 3m, average flow rage Q = 5lpm = 1.67 e — 5m3/s,so Reynolds = 4pQ/πμ d = 582 < 2300.

Meanwhile, in order to accelerate the simulation, a two-dimensional domain decomposition was applied and 9 and 5 was used uniform subdomains in y- and z-directions respectively, which yields the deployment of a total number of 45 CPU cores. The simulation used the high- performance computing facility of Wiegand Advanced Visualization Environment at Santa Clara University (SCU WAVE HPC).

Results

MCS Pump Performance Characterization. A series of experiments were conducted with various operating conditions. The motor speed varied from 100 [RPM] to 300 [RPM] while the system pressure difference was adjusted by the clamp from 0 to 120 [mmHg]. The resulting pressure difference versus flowrate curves of the MCS prototype are shown in Figure 6. Non-positive displacement feature can be noted that the flowrate decreases as the load pressure across the pump increases. Compared with prevailing hydrodynamic rotary VADs, such as Heartware, HeartMate2 [18] and CH-VAD [19], the sliding vane MCS pump has flatter HQ curves and can be regarded as positive displacement pump with leakage. The flatness of the HQ curves shown in Figure 6 appear to be due to the positive displacement pumping architecture.

The leakage is represented in Figures 7A-7C by a gap depicted in box B) between rotary cylinder and flow straightener’ s flow separator, which provides a leakage pathway from outlet directly to inlet, and a gap (depicted in box C) between vane and housing wall, which provides leakage from one chamber to the other. Nonetheless, fairly linear relationships between pressure and flowrate are apparent (R ² > 0.9700) and corresponding linear least squares regression fittings are obtained as follows (with Q, w, and AP have units of [LPM], [PPM], and [mmHg] respectively)

Q(w| ₁₀₀ , ΔP) = -1.045 x 10 ^-2 ΔP + 1.9095, R ² = 0.9762 (12)

Q(w| ₁₅₀ , ΔP) = -9.01 x 10 ^-3 ΔP + 3.2570, R ² = 0.9804 (13)

Q(w| ₂₀₀ , ΔP) = -9.34 x 10 ^-3 ΔP + 4.2960, R ² = 0.9783 (14)

Q(w| ₂₅₀ , ΔP) = -9.10 x 10 ^-3 ΔP + 5.4050, R ² = 0.9702 (15)

Q(w| _3oo , AP) = -1.032 x 10 ^-2 ΔP + 6.6086, R ² = 0.9778 (16)

Theoretically, with specifications shown in Table 1, the MCS pump is a positive displacement pump and has a displacement of 24.8 [mL/rev]. The theoretical output flowrate with respect to rotary speed is plotted as a solid line in Figure 8. No-leakage flowrates of the experimental MCS prototype pump under different operating speeds can be acquired by extrapolating the above linear fittings (dashed lines) and intersecting with the y-axis. The intersect points are depicted as open squares in Figure 8. The noleakage flowrates under 100 [PPM] to 300 [PPM] show that the prototype has a fairly linear displacement pumping ability with respect to rotary speed. The mathematical model developed previously for displacement calculation matches well with experimental displacement result (93.10%).

V _{disp-.experimental} /V _{disp : theoretical} = 23.09/24.80 (17)

= 93.10%

Pulsatile Flow. To illustrate the MCS pump's pulsatility, an estimator of the MCS flowrate is needed. A first-order polynomial equation with ΔP [mmHg], and ω [RPM] as input variables was found to match well with the experiment data presented previously (Figure 6) and is used here to estimate flowrate Q (t) [LPM] :

Q _est (t) = α ₁ ΔP(t) + α ₂ ω(t) + α ₃ (18) where parameters (α ₁ , α ₂ , α ₃ ) = (—9.644 x 10 ^-3 , 2.309 x 10 ^-2 , — 3.233 x 10 ^-1 ). Real time pressure sensor data and motor speed data were collected and resampled due to the sensors’ different sampling frequencies before the data were plugged into Eqn. 18. The estimated flowrate, Q _est , was then compared with the theoretical flowrate. These flowrate curves under several operating conditions are plotted in Figures 9A-11B.

As the figures illustrate, the estimated instantaneous flowrate of the prototype MCS pump shows pulsatile behavior. Flowrate varies noticeably as the rotary angle increments. Since the first-order polynomial fitting function to calculate instantaneous flowrate is derived using the data in Figure 6, the discrepancy between the average experimental (fitting) flowrate and the average theoretical flowrate is consistent with the data in Figure 8.

Experimental tests were conducted to characterize the MCS pump prototype and to validate the developed model. H-Q curves of the proposed MCS pump were experimentally measured.

The leakage could happen between the rolling cylinder and the flow straightener (see FIG. 7B). A mechanism like a wiper blade between them could reduce leakage through this pathway. In certain applications, a small degree of leakage may be designed in, to control the overall shear applied to the traversing fluid, balancing net output against shear stress and stress accumulation.

Another possible leakage source is the gap between the vane and the wall (see Fig. 7C). Although other solutions to Eqn. 1 exist to allow constant contact between sliding vane end and the wall, manufacture tolerance still exists, leading to gaps. Smaller gaps will lead to less backflow leakage, thus increasing the pump efficiency. However, this would also contribute to higher shear stresses onto blood, causing blood damage. Alternative vane designs are provided (Figures 12A-12C) to introduce a gap for reducing shear stress on the fluid but to retain the pump performance as a positive displacement pump.

The proposed MCS pump, mathematically and experimentally, showed a more physiologic pulsatile flow generation, compared to the attenuated or non-pulsatile flow generated from prevailing rotary VADs. A flowrate estimator using polynomial fittings was adopted.

Unless defined otherwise, all technical and scientific terms used herein have the same meanings as commonly understood by one of skill in the art to which the disclosed invention belongs. Publications cited herein and the materials for which they are cited are specifically incorporated by reference.

Those skilled in the art will recognize, or be able to ascertain using no more than routine experimentation, many equivalents to the specific embodiments of the invention described herein. Such equivalents are intended to be encompassed by the following claims.

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