Foldable impeller and blood pump thereof

Abstract

The present application aims to provide a foldable impeller with good resilience and folding effect and a blood pump thereof, comprising a hub and blades arranged in the circumferential direction of the hub, the blade comprises an action surface, the action surface is cut along the axial direction to form a cutting gap, the action surface comprises a concave surface and a convex surface, the cutting gap extends from the convex surface to the middle part of the blade, and the blade is folded along the cutting gap and wrapped outside the hub periphery under the constraint force. The structure is compressed by the collapsible shell when being contracted, folded at the cutting gap, and the overall impeller is folded and contracted. When the overall folded heart blood pump enters the left ventricle, the folded shell will expand due to the loss of constraint, and after the rotation shaft is started to rotate, the hub rotates, and due to the centrifugal force and the resilience of the embedded part, the folded impeller will automatically unfold and be unfolded completely when the gap on the back of the impeller flow passage is closed. The impeller flow passage has no gap and is a complete and smooth flow surface, and hemolysis is small.

Description

Technical Field

[0001] This invention relates to the field of medical device technology, specifically to a foldable impeller and its blood pump. Background Technology

[0002] Various heart diseases, such as heart failure, myocardial infarction, and myocardial injury, can impair the pumping function of the ventricles. Currently, the primary treatment for these patients involves ventricular assist devices (VADs), which not only help the heart pump blood, reducing the burden on the myocardium and aiding in its recovery, but also prevent damage to vital organs such as the brain and kidneys due to ischemia when the heart's pumping function declines or even disappears.

[0003] For interventional cardiac blood pumps, there are three ideal performance requirements: longer clinical support time, smaller size, and better hemodynamics. However, these three points are often difficult to balance simultaneously. Smaller size makes it difficult to achieve greater flow support, and longer support time brings a greater risk of hemolytic events. Therefore, in the latest research direction of blood pumps, foldable blood pumps have become the most popular direction. As a key component, the shape and folding method of the foldable impeller are crucial to hemolytic performance and pumping volume. Currently, most foldable impellers are generally made of thermoplastic elastomer through one-piece injection molding. Relying on the resilience of thermoplastic elastomer, it enters the heart after being folded under force and then expands with the expansion of the outer shell. However, this material is extremely difficult to achieve 100% resilience, and there will be bending gaps at the root of the impeller during compression and folding, resulting in poor compression and folding effect. Summary of the Invention

[0004] One of the objectives of this invention is to provide a foldable impeller with good resilience and folding effect.

[0005] To achieve the above objectives, the technical solution adopted by the present invention is as follows: a foldable impeller, including a hub and blades arranged circumferentially thereon, wherein the blades include a working surface, and a cutting slit is formed on the working surface along its axial direction. The working surface includes a concave surface and a convex surface. The cutting slit extends from the convex surface to the middle of the blade. When the blade is under a constraint force, it folds along the cutting slit and wraps around the outer periphery of the hub.

[0006] The blade has a pre-embedded shape memory alloy, which is in the form of a sheet or chain. The shape memory alloy extends outward from the hub, and the cutting slit extends from the convex surface to the middle of the blade to the shape memory alloy.

[0007] The distance between the shape memory alloy and the concave surface is d, where d ranges from 0.03 mm to 0.06 mm.

[0008] The cutting slits are arranged in multiple spaces along the radial direction of the blade, and the spacing between adjacent cutting slits is different.

[0009] There are a total of 3 cutting gaps, including the first gap, the second gap and the third gap from the wheel hub outwards. The distance between the first gap and the outer circumference of the wheel hub is 0.5mm, the distance between the second gap and the outer circumference of the wheel hub is 1.5mm, and the distance between the third gap and the outer circumference of the wheel hub is 2.5mm.

[0010] The contour line of the working surface includes an outer edge profile away from the hub. The endpoint of the outer edge profile near the inlet end is the profile start point, and the endpoint of the outer edge profile near the outlet end is the profile end point. The angle between the axial plane where the profile start point is located and the axial plane where the profile end point is located is the blade deflection angle θ. The blade deflection angle θ is 90° to 180°.

[0011] The angle between the tangent of the impeller arc at the point where the blade extends from the hub and the tangent of the hub is the blade inflow angle α, which is 90° to 180°.

[0012] Shape memory alloys are selected from one of the following: nickel-titanium alloys, titanium-nickel-copper alloys, titanium-nickel-iron alloys, and titanium-nickel-chromium alloys.

[0013] The blades are one or more of polyester, polyamide (PA), polyetheretherketone (PEEK), and polyurethane (TPU) with a hardness greater than 70D.

[0014] One of the objectives of this invention is to provide a blood pump with good resilience and folding effect.

[0015] To achieve the above objectives, the technical solution adopted by the present invention is as follows: a blood pump, comprising a catheter and a foldable impeller disposed within the catheter.

[0016] In the above scheme, the structure is compressed by the retractable outer shell during contraction, folding at the cut gap to achieve the overall folding and contraction of the impeller. When the folded heart pump enters the left ventricle, the folded outer shell expands due to the release of its restraints. After the rotating shaft is started, the hub rotates, and due to the centrifugal force and the rebound force of the embedded parts, the folded impeller will automatically unfold, completing the unfolding when the gap on the back of the impeller flow channel closes. The impeller flow channel has no gaps, forming a complete and smooth flow surface, resulting in minimal hemolysis. Attached Figure Description

[0017] Figure 1 This is a schematic diagram of the impeller in its deployed state;

[0018] Figure 2 This is a schematic diagram of the impeller in its folded state;

[0019] Figure 3 for Figure 1 The main view;

[0020] Figure 4 for Figure 2 The main view;

[0021] Figure 5 for Figure 1 Top view;

[0022] Figure 6 for Figure 2 Top view;

[0023] Figure 7a The simulation results of wall shear stress are shown when the blade inflow angle α is 120° and the blade deflection angle θ is 180°.

[0024] Figure 7b The images show the simulated velocity vector effects when the blade inflow angle α is 120° and the blade deflection angle θ is 180°, respectively.

[0025] Figure 7c The images show simulation results at various points in time, corresponding to the start of timing when blood flows through the inlet, with the inflow angle α being 120° and the deflection angle θ being 180°.

[0026] Figure 8a The simulation results of wall shear stress are shown when the blade inflow angle α is 140° and the blade deflection angle θ is 180°.

[0027] Figure 8b The images show the simulated velocity vector effects when the blade inflow angle α is 140° and the blade deflection angle θ is 180°, respectively.

[0028] Figure 8c The simulation results are shown at various time points corresponding to the start of timing when blood flows through the inlet with a blade inflow angle α of 140° and a blade deflection angle θ of 180°.

[0029] Figure 9a The simulation results of wall shear stress are shown when the blade inflow angle α is 160° and the blade deflection angle θ is 180°.

[0030] Figure 9b The images show the simulated velocity vector effects when the blade inflow angle α is 160° and the blade deflection angle θ is 180°, respectively.

[0031] Figure 9c The images show simulation results at various points in time, corresponding to the start of timing when blood flows through the inlet, with the inflow angle α being 160° and the deflection angle θ being 180°.

[0032] Figure 10a The simulation results of wall shear stress are shown when the blade inflow angle α is 140° and the blade deflection angle θ is 120°.

[0033] Figure 10b The images show the simulated velocity vector effects when the blade inflow angle α is 140° and the blade deflection angle θ is 120°, respectively.

[0034] Figure 10c The images show simulation results at various points in time, corresponding to the start of timing when blood flows through the inlet, with a blade inflow angle α of 140° and a blade deflection angle θ of 120°.

[0035] Figure 11a The simulation results of wall shear stress are shown when the blade inflow angle α is 140° and the blade deflection angle θ is 150°.

[0036] Figure 11b The images show the simulated velocity vector effects when the blade inflow angle α is 140° and the blade deflection angle θ is 150°, respectively.

[0037] Figure 11c The simulation results are shown at various time points corresponding to the start of timing when blood flows through the inlet with a blade inflow angle α of 140° and a blade deflection angle θ of 150°.

[0038] Figure 12a The simulation results of wall shear stress are shown when the blade inflow angle α is 140° and the blade deflection angle θ is 180°.

[0039] Figure 12b The images show the simulated velocity vector effects when the blade inflow angle α is 140° and the blade deflection angle θ is 180°, respectively.

[0040] Figure 12c The images show simulation results at various points in time, corresponding to the start of timing when blood flows through the inlet, with the inflow angle α being 140° and the deflection angle θ being 180°. Detailed Implementation

[0041] The hydraulic and hemolytic properties of a blood pump are closely related to the structure of its impeller.

[0042] Hydraulic performance refers to the flow rate and head that a blood pump can achieve under target volume and operating conditions.

[0043] To enable blood pumps to meet the blood circulation needs of different patients, they can be classified into various types based on their blood output per minute, such as 2.5L / min, 3.5L / min, and 5.0L / min blood pumps.

[0044] Hemolysis refers to the rupture of red blood cells, causing the hemoglobin within them to leak out and dissolve in the blood. Hemolysis leads to changes in the morphology and biochemical properties of red blood cells, shortens their lifespan, and can even result in complete rupture, thereby reducing their ability to transport oxygen to tissues and organs. Furthermore, hemolysis increases the concentration of free hemoglobin in the plasma; the excess free hemoglobin needs to be excreted through the kidneys, potentially leading to kidney damage and multiple organ failure.

[0045] Therefore, preventing hemolysis of the pumped blood is crucial when performing blood pumping on medical patients. Hemolysis during pumping can endanger the patient's life.

[0046] To ensure the safe use of blood pumps, hemolytic performance needs to be considered. Hemolytic performance refers to the probability that hemolysis will occur due to blood cell breakage and damage after the blood flows through the blood pump.

[0047] To meet the requirements of hydraulic performance and hemolytic performance, some related technologies employ micro axial flow structures for their impellers.

[0048] The applicant's research found that due to the special nature of blood pump usage scenarios, there are still two key technical challenges in the design and manufacture of the impeller: First, in order to reduce the impact of implantation on normal human physiology, the outer diameter of the impeller is limited to within 7mm, which is suitable for passing through blood vessels. This severely limits the volume of the impeller and the blades on the impeller, making it difficult for the impeller to achieve the required hydraulic performance. Second, in order to meet the pressure requirements of human blood circulation, it is necessary to increase the impeller speed to improve the pumping performance. However, the high-speed rotation of the impeller will increase the shear stress of the blood flow field, leading to blood cell breakage and mechanical hemolysis.

[0049] Therefore, the structure of a blood pump impeller needs to balance hydraulic performance and hemolytic performance. The main considerations for hydraulic performance are the output flow rate and the pressure difference between the pump inlet and outlet, while the main considerations for hemolytic performance are the shear stress on the blood during pumping and the smoothness of blood flow. By optimizing and improving the impeller structure, hydraulic performance can be enhanced while reducing blood damage caused by hemolysis.

[0050] For the reasons mentioned above, this application provides a foldable impeller, including a hub 10 and blades 20 arranged circumferentially thereon. Each blade 20 includes an action surface 21, on which axial cuts are made to form a cutting slit 22. The action surface 21 includes a concave surface 21a and a convex surface 21b. The cutting slit 22 extends from the convex surface 21b towards the center of the blade 20. When under constraint, the blade 20 folds along the cutting slit 22 and wraps around the outer periphery of the hub 10. During contraction, this structure is compressed by a retractable outer shell and folds at the cutting slit 22, achieving the folding and contraction of the blade 20. When the folded blood pump enters the ventricle, the folded outer shell expands due to the loss of constraint. After the rotating shaft is started, the hub 10 rotates, and due to centrifugal force, the folded blades 20 automatically unfold, and the cutting slit 22 disappears. The impeller flow channel has no gaps and is a completely smooth flow surface, which will not have a negative effect on hemolysis.

[0051] Shape memory alloy 23 is pre-embedded inside the blade 20. The shape memory alloy 23 is in the form of sheets or chains and extends outward from the hub 10. The cutting slit 22 extends from the convex surface 21b to the middle of the blade 20 and reaches the shape memory alloy 23. During impeller injection molding, sheet-like or chain-like shape memory alloy 23 is left at the fixed position of the blade cavity in the impeller mold. After injection molding, the blade 20 wraps the pre-embedded part, and then the molded impeller is cut with a precision cutting machine to cut the plastic part on the back of the impeller flow channel to the surface of the pre-embedded alloy.

[0052] Furthermore, the distance between the shape memory alloy 23 and the concave surface 21a is d, where d ranges from 0.03mm to 0.06mm. The shape memory alloy 23 enhances the strength of the blade 20 on the one hand, and provides driving force for the unfolding of the blade 20 on the other hand, enabling the blade 20 to fully unfold and ensuring the pumping effect.

[0053] In order to completely wrap the blade 20 around the outer periphery of the hub 10, multiple cutting slits 22 are provided at radial intervals along the blade 20, and the spacing between adjacent cutting slits 22 is different.

[0054] Preferably, there are three cutting gaps 22, including a first gap, a second gap and a third gap extending outward from the hub 10. The distance between the first gap and the outer circumferential surface of the hub 10 is 0.5 mm, the distance between the second gap and the outer circumferential surface of the hub 10 is 1.5 mm, and the distance between the third gap and the outer circumferential surface of the hub 10 is 2.5 mm.

[0055] The contour line of the working surface 21 includes an outer edge profile away from the hub 10. The endpoint of the outer edge profile near the inlet end is the profile start point, and the endpoint of the outer edge profile near the outlet end is the profile end point. The angle between the axial plane where the profile start point is located and the axial plane where the profile end point is located is the blade deflection angle θ. The blade deflection angle θ is 90° to 180°. The blade deflection angle θ can be more intuitively understood as the angle formed by the line connecting the profile start point and the axis of the hub 10, and the line connecting the profile end point and the axis of the hub 10, in the top view projection of the impeller. The size of the blade deflection angle θ reflects the length of each blade 20 extending circumferentially on the hub 10. If the blade deflection angle θ is too small, the flow rate may not meet the requirements under high pressure; if the blade deflection angle θ is too large, the flow rate may be too small under low pressure.

[0056] The angle between the tangent of the impeller arc at the position where the blade 20 extends from the hub 10 and the tangent of the hub 10 is the blade inflow angle α, which is 90° to 180°.

[0057] According to research, hemolysis is mainly related to shear stress and red blood cell exposure time. When the critical stress value exceeds 1000 Pa, even with extreme exposure time, red blood cells will be destroyed. Conversely, when the shear stress is in the range of 150 Pa to 1000 Pa, prolonged exposure time will also cause red blood cell rupture. Therefore, we mainly simulate and calculate the shear stress and exposure time, and then calculate the hemolysis rate based on the hemolysis formula obtained experimentally by GIERSIEPEN et al. This model indicates that the degree of hemolysis has a power function relationship with the flow field shear stress and the exposure time of red blood cells under shear stress.

[0058] ;

[0059] In the formula: Total mass of hemoglobin; This refers to the amount of free hemoglobin; denoted as the shear stress scalar in the flow field; t is the exposure time of the red blood cells. To allow for a margin in the hemolysis calculation, the maximum shear stress is used in this application for calculation.

[0060] For the impeller's shape design, we used Unigraphics NX to design the impeller model and used CFD simulation to simulate data such as the impeller's wall shear stress and the time it takes for blood to flow through the blood pump, analyze the hemolytic performance, and optimize the impeller structure to obtain the optimal impeller.

[0061] This application first determines that when the blade deflection angle θ is 180°, the blade inflow angle α is adjusted for simulation to simulate the impeller operation state. Under the conditions of achieving a flow rate of 5 L / min and a fixed pressure difference between the impeller outlet and inlet, the surface shear stress of the impeller, the velocity distribution of blood flowing through the impeller, and the blood flow time through the impeller are obtained to calculate the hemolysis results and find the optimal blade inflow angle α. In the design, the blade inflow angle α is simulated from 90° to 180°. This application uses simulation results of 120°, 140°, and 160° for demonstration and comparison (example labeling: LR120PZ180 is used to label impellers with a blade inflow angle α of 120° and a blade deflection angle θ of 180°). The statistical data from the simulation results are shown in Table 1.

[0062] Table 1

[0063] Serial Number Data categories LR120PZ180 LR140PZ180 LR160PZ180 1 Maximum shear stress (Pa) 762 726 763 2 The longest time (s) that blood flows through the impeller. 0.0027 0.0025 0.0026 3 Hemolysis value calculation 0.0323 0.027 0.0314

[0064] After obtaining the optimal impeller extension angle, a 140° impeller extension angle was set, and the blade deflection angle was adjusted for simulation to simulate the impeller's operating state. The surface shear stress of the impeller and the time it takes for blood to flow through the impeller were obtained to calculate the hemolysis results and find the optimal blade deflection angle θ. In the design, the blade deflection angle θ was simulated from 90° to 180°. This application uses simulation results of 120°, 150°, and 180° for comparison and demonstration. Statistical data from the simulation results are shown in Table 2.

[0065] Table 2

[0066] Serial Number Data categories LR140PZ120 LR140PZ150 LR140PZ180 1 Maximum shear stress (Pa) 886 779 726 2 The longest time (s) that blood flows through the impeller. 0.0027 0.0028 0.0025 3 Hemolysis value calculation 0.0465 0.035 0.027

[0067] Based on the impeller simulation results (such as...) Figures 7a to 12c Based on the results of hemolysis calculations (as shown in Tables 1 and 2), the impeller structure with the lowest hemolysis value and the best effect was found to have a blade inflow angle α of 140° and a blade deflection angle θ of 180°.

[0068] Shape memory alloys are materials composed of two or more metallic elements that exhibit shape memory effects through thermoelasticity and martensitic phase transformation and its inverse. Shape memory alloy 23 is selected from one of the following: nickel-titanium alloy, titanium-nickel-copper alloy, titanium-nickel-iron alloy, and titanium-nickel-chromium alloy.

[0069] The blade 20 is made of one or more of polyester, polyamide (PA), polyetheretherketone (PEEK), and polyurethane (TPU) with a hardness greater than 70D. These materials have high hardness and will not deform during impeller operation.

[0070] A blood pump includes a conduit and a foldable impeller disposed within the conduit. The blood pump of the above embodiment relies on the corresponding impeller in any of the foregoing embodiments to pump blood and has the beneficial effects of the corresponding impeller embodiments, which will not be described in detail here.

[0071] It should be noted that the above description describes some embodiments of this application. Other embodiments are within the scope of the appended claims. In some cases, the actions or steps recorded in the claims can be performed in a different order than that shown in the above embodiments and still achieve the desired result. Furthermore, the processes depicted in the drawings do not necessarily require a specific or sequential order to achieve the desired result. In some embodiments, multitasking and parallel processing are also possible or may be advantageous.

Claims

1. A foldable impeller, comprising a hub (10) and blades (20) arranged circumferentially thereon, characterized in that: The blade (20) includes an action surface (21), on which a cutting slit (22) is formed by cutting along its axial direction. The action surface (21) includes a concave surface (21a) and a convex surface (21b). The cutting slit (22) extends from the convex surface (21b) toward the middle of the blade (20). When the blade (20) is under constraint, it folds along the cutting slit (22) and wraps around the outer periphery of the hub (10). When the constraint is released and the blade (20) rotates, the cutting slit (22) disappears. The blade (20) is pre-embedded with a shape memory alloy (23), which is in the form of a sheet or a chain. The shape memory alloy (23) extends outward from the hub (10), and the cutting slit (22) extends from the convex surface (21b) to the middle of the blade (20) to the shape memory alloy (23).

2. The foldable impeller according to claim 1, characterized in that: The distance between the shape memory alloy (23) and the concave surface (21a) is d, and the value of d is 0.03mm to 0.06mm.

3. The foldable impeller according to claim 1, characterized in that: The cutting slits (22) are arranged in multiple spaces along the radial distance of the blade (20), and the spacing between adjacent cutting slits (22) is different.

4. The foldable impeller according to claim 3, characterized in that: There are three cutting gaps (22), including a first gap, a second gap and a third gap from the hub (10) outward. The distance between the first gap and the outer circumference of the hub (10) is 0.5 mm, the distance between the second gap and the outer circumference of the hub (10) is 1.5 mm, and the distance between the third gap and the outer circumference of the hub (10) is 2.5 mm.

5. The foldable impeller according to claim 1, characterized in that: The contour of the working surface (21) includes an outer edge profile away from the hub (10). The endpoint of the outer edge profile near the inlet end is the profile start point, and the endpoint of the outer edge profile near the outlet end is the profile end point. The angle between the axial plane where the profile start point is located and the axial plane where the profile end point is located is the blade deflection angle θ. The angle of the blade deflection angle θ is 90° to 180°.

6. The foldable impeller according to claim 1, characterized in that: The angle between the tangent of the impeller arc at the position where the blade (20) extends from the hub (10) and the tangent of the hub (10) is the blade inflow angle α, which is 90° to 180°.

7. The foldable impeller according to claim 1, characterized in that: Shape memory alloy (23) is selected from one of nickel-titanium alloy, titanium-nickel-copper alloy, titanium-nickel-iron alloy, and titanium-nickel-chromium alloy.

8. The foldable impeller according to claim 1, characterized in that: The blade (20) is injection molded from one or more of polyester, polyamide (PA), polyether ether ketone (PEEK), and polyurethane (TPU) with a hardness greater than 70D.

9. A blood pump, characterized in that: Includes a conduit and a foldable impeller as described in any one of claims 1 to 8 disposed within the conduit.

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