Source: http://www.freepatentsonline.com/8870552.html
Timestamp: 2018-08-15 10:49:27
Document Index: 301308761

Matched Legal Cases: ['art 2001', 'artz\n5275580', 'art 3', 'art 4', 'art 100', 'art 100', 'art 19', 'art 20', 'art 20', 'art 13', 'art 13']

Rotary blood pump and control system therefor - Thoratec Corporation
United States Patent 8870552
Ayre, Peter Joseph (Crows Nest, AU)
Tansley, Geoffrey Douglas (Mt. Colah, AU)
Watterson, Peter Andrew (West Ryde, AU)
Woodard, John Campbell (Thornleigh, AU)
13/594691
415/111, 415/900, 600/16, 623/3.13
F01D11/00; A61M1/10; A61N1/362
417/423.1, 417/423.14, 415/104, 415/106, 415/900, 415/110, 415/111, 600/16, 623/3.13
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JP8284841 October, 1996
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JP2002224066A 2002-08-13 CARDIAC FUNCTION EVALUATING DEVICE
JP2004278375A 2004-10-07 AXIAL FLOW PUMP
WO1991019103A1 1991-12-12 HYDRODYNAMICALLY SUSPENDED ROTOR AXIAL FLOW BLOOD PUMP
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WO2001005023A1 2001-01-18 METHOD AND APPARATUS FOR CONTROLLING BRUSHLESS DC MOTORS IN IMPLANTABLE MEDICAL DEVICES
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WO2004028593A1 2004-04-08 PHYSIOLOGICAL DEMAND RESPONSIVE CONTROL SYSTEM
JPH08284841A 1996-10-29
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Barletta et al. “Design of a bearingless blood pump”, Proceedings from Third Int. Symposium on Magnetic Suspension Technology, Ed. By Nelson J. Groom and Colin P. Britcher Jul. 1996, pp. I-XIII and 265-274.
This application is a continuation of U.S. application Ser. No. 12/538,824, filed Aug. 10, 2009 (now U.S. Pat. No. 8,282,359 issued) which is a continuation application of U.S. application Ser. No. 10/921,662, filed Aug. 19, 2004 now abandoned, which is a divisional application of U.S. application Ser. No. 09/980,682 filed Aug. 15, 2002 (now U.S. Pat. No. 6,866,625 issued Mar. 15, 2005) which is a national phase application of the International Application PCT/AU00/00355 filed Apr. 20, 2000 and claims the priority benefits of U.S. application Ser. No. 09/299,038 filed Apr. 23, 1999 (issued Jun. 26, 2001 as U.S. Pat. No. 6,250,880) and Australian PP 9959 filed Apr. 23, 1999, all of which are incorporated herein by reference in their entireties.
1. A rotary blood pump, comprising: a pump housing forming a pump cavity; an impeller located within the pump cavity; an impeller drive including a commutator and windings, wherein the commutator is configured to generate an oscillating current within the windings to thereby produce a rotating magnetic field for driving permanent magnets in the impeller; and a porous metal substrate disposed on a housing surface and comprising one of amorphous carbon based materials or microcrystalline carbon based materials; wherein when the pump is in use the porous metal substrate is in communication with a fluid flowing through the housing.
2. The rotary blood pump of claim 1, wherein in use a hydrodynamic bearing forms between inner walls of the pump cavity and an opposing surface of the impeller.
3. The rotary blood pump of claim 1, wherein the substrate has a porosity such that a pseudoneointimal cell lining is capable of growing into pores of the substrate when the pump is in use.
4. The rotary blood pump of claim 1, wherein the porous metal substrate comprises a surface coating disposed on inner walls of the pump cavity to reduce wear on the housing surface in the event of a touchdown event.
5. The rotary blood pump of claim 1, wherein a surface of the impeller opposing the housing surface is made from a different material than the porous metal substrate.
6. The rotary blood pump of claim 5, wherein the impeller surface is Titanium Nitride.
7. A rotary blood pump, comprising: a pump housing forming a pump cavity; an impeller located within the pump cavity; an impeller drive including a commutator and windings, wherein the commutator is configured to generate an oscillating current within the windings to thereby produce a rotating magnetic field for driving permanent magnets in the impeller; and a porous metal substrate comprising a surface coating disposed on inner walls of the pump cavity to reduce wear on the housing surface in the event of a touchdown event, wherein the coating thickness is about 1 micron and the material is titanium nitride; and wherein when the pump is in use the porous metal substrate is in communication with a fluid flowing through the housing.
8. The rotary blood pump of claim 7, wherein in use a hydrodynamic bearing forms between the inner walls of the pump cavity and an opposing surface of the impeller.
9. The rotary blood pump of claim 7, wherein the substrate has a porosity such that a pseudoneointimal cell lining is capable of growing into pores of the substrate when the pump is in use.
10. The rotary blood pump of claim 7, wherein a surface of the impeller opposing the coating is made from a different material than the coating.
11. The rotary blood pump of claim 10, wherein the impeller surface is Titanium Nitride.
12. A rotary blood pump, comprising: a pump housing forming a pump cavity; an impeller located within the pump cavity; an impeller drive including a commutator and windings, wherein the commutator is configured to generate an oscillating current within the windings to thereby produce a rotating magnetic field for driving permanent magnets in the impeller; and a porous metal substrate comprising a surface coating disposed on inner walls of the pump cavity to reduce wear on the housing surface in the event of a touchdown event, wherein the coating thickness is about 1 micron; and wherein when the pump is in use the porous metal substrate is in communication with a fluid flowing through the housing.
13. The rotary blood pump of claim 12, wherein the coating comprises one of amorphous carbon based materials or microcrystalline carbon based materials.
14. The rotary blood pump of claim 12, wherein in use a hydrodynamic bearing forms between the inner walls of the pump cavity and an opposing surface of the impeller.
15. The rotary blood pump of claim 12, wherein the substrate has a porosity such that a pseudoneointimal cell lining is capable of growing into pores of the substrate when the pump is in use.
16. The rotary blood pump of claim 12, wherein a surface of the impeller opposing the coating is made from a different material than the coating.
17. The rotary blood pump of claim 16, wherein the impeller surface is Titanium Nitride.
This invention relates to rotary pumps adapted, but not exclusively, for use as artificial hearts or ventricular assist devices and, in particular, discloses in preferred forms a seal-less shaft-less pump featuring open or closed (shrouded) impeller blades with at least parts of the impeller used as hydrodynamic thrust bearings and with electromagnetic torque provided by the interaction between magnets embedded in the blades or shroud and a rotating fixed relative to the current pattern generated in coils pump housing.
Contact or pivot bearings, as exemplified by U.S. Pat. No. 5,527,159 to Bozeman et al. and U.S. Pat. No. 5,399,074 to Nose et al., have potential problems due to wear, and cause very high localised heating and shearing of the blood, which can cause deposition and denaturation of plasma −3 proteins, with the risk of embolisation and bearing seizure.
Magnetic bearings, as exemplified by U.S. Pat. No. 5,350,283 to Nakazeki et al., U.S. Pat. No. 5,326,344 to Bramm et al. and U.S. Pat. No. 4,779,614 to Moise et al., offer contactless suspension, but require rotor position measurement and active control of electric current for stabilisation of the position in at least one direction, according to Eamshaw's theorem. Position measurement and feedback control introduce significant complexity, increasing the failure risk. Power use by the control current implies reduced overall efficiency. Furthermore, size, mass, component count and cost are all increased.
Accordingly, in one broad form of the invention there is provided a rotary blood pump for use in a heart assist device or like device, said pump having an impeller suspended in use within a pump housing exclusively by hydrodynamic thrust forces generated by relative movement of said impeller with respect to and within said pump housing; and wherein at least one of said impeller of said housing includes at least a first deformed surface lying on at least part of a first face and a second deformed surface lying on at least part of a second face which, in use, move relative to respective facing surfaces on the other of said impeller or said housing thereby to form at least two relatively moving surfaces pairs which generate relative hydrodynamic thrust between said impeller and said housing which includes everywhere a localized thrust component substantially and everywhere normal to the plane of movement of said first deformed surface and said second deformed surface with respect to said facing surfaces; and wherein the combined effect of the localized normal forces generated on the surfaces of said impeller is to produce resistive forces against movement in three translational and two rotational degrees of freedom.
In yet a further broad form of the invention there is provided an estimation and control system for a pump; said pump of the type having an impeller located within a pump cavity in a pump housing; said housing having a fluid inlet in fluid communication with said pump cavity; said impeller urged to rotate about an impeller axis so as to cause fluid to be urged from said inlet through said pump cavity to said pump outlet; said impeller urged to rotate by impeller urging means; said impeller supported for rotational movement by impeller support means; said impeller maintained at or near a predetermined speed of rotation by control means acting on said impeller urging means; said control means receiving as input variables a first input variable comprising power consumed by said urging means; said control means receiving a second input variable comprising actual speed of rotation of said impeller; said control means thereby estimating head across the pump and/or rate of flow of said fluid to an approximation of predetermined accuracy relying on signals available from said urging means; said control system adapted to maintain speed of rotation of said impeller within a range whereby said impeller, in use, substantially resists five degrees of freedom of movement with respect to said pump housing predominantly without any external intervention from said control system to control the position of said impeller with respect to said housing.
In yet a further broad form of the invention there is provided a rotary blood pump and an estimation and control system therefor, said pump having an impeller suspended hydrodynamically within a pump housing by thrust forces generated by the impeller during movement in use of the impeller as it rotates about an impeller axis; said estimation and control system of the type described above.
In yet a further broad form of the invention there is provided a rotary blood pump having a housing within which an impeller acts by rotation about an impeller axis to cause a pressure differential between an inlet side of the pump housing of said pump and an outlet side of the pump housing of said pump; said impeller suspended hydrodynamically by thrust forces generated by the impeller during movement in use of the impeller; said pump controlled by the estimation and control system as described above.
In yet a further broad form of the invention there is provided a seal-less, shaft-less pump comprising a housing defining a chamber therein and having a liquid inlet to said chamber and a liquid outlet from said chamber; said pump further including an impeller located within said chamber; the arrangement between said impeller, said inlet, said outlet, and the internal walls of said chamber being such that upon rotation of said impeller about an impeller axis relative to said housing liquid is urged from said inlet through said chamber to said outlet; and wherein thrust forces are generated by relative movement of said impeller with respect to said housing; said pump controlled by the estimation and control system as described above.
In yet a further broad form of the invention there is provided a pump having a housing within which an impeller acts by rotation about an axis to cause a pressure differential between an inlet side of a housing of said pump and an outlet side of the housing of said pump; said impeller suspended exclusively hydrodynamically in at least one of a radial or axial direction by thrust forces generated by the impeller during movement in use of the impeller; said pump controlled by the estimation and control system as described above.
In yet a further broad form of the invention there is provided a method of hydrodynamically suspending and controlling an impeller within a rotary pump for support in at lest one of a radial or axial direction; said method comprising incorporating a deformed surface in at least part of said impeller so that, in use, a thrust is created between said deformed surface and the adjacent pump casing during relative movement therebetween; said method further including the step of maintaining speed of rotation of said impeller within a range whereby said impeller, in use, substantially resists five degrees of freedom of movement with respect to said pump housing without any external intervention.
In yet a further broad form of the invention there is provided an estimation and control system for a pump; said pump of the type having an impeller located within a pump cavity in a pump housing; said housing having a fluid inlet in fluid communication with said cavity; said housing having a fluid outlet in fluid communication with said pump cavity; said impeller urged to rotate about an impeller axis so as to cause fluid to be urged from said inlet through said pump cavity to said pump outlet; said impeller urged to rotate by impeller urging means; said impeller supported for rotational movement by impeller support means; said pump maintained at or near a predetermined operating point by control means acting on said impeller urging means; said control means receiving as input at least a first input variable derived from said urging means; said control means receiving at least a second input variable also derived from said urging means; said control means thereby calculating an estimate of said operating point to an approximation of predetermined accuracy relying on signals available from said urging means; said control means controlling said pump by comparing said predetermined operating point with said estimate of said operating point; and wherein instantaneous pump speed and electrical input power are allowed to be modulated by the heart, in use, by appropriate selection of a control time constant.
Preferably said pump comprises a ventricular assist device adapted to assist operation of a ventricle of a hear and wherein said control means adjusts pump output so that, in alternation fashion, said ventricle in conjunction with said aortic valve is allowed to eject blood over a predetermined number of cardiac cycles and then said ventricle in conjunction with said aortic valve is caused to eject blood over a following predetermined number of cardiac cycles.
In yet a further broad form of the invention there is provided an estimation and control system for a pump; said pump of the type having an impeller located within a pump cavity in a pump housing; said housing having a fluid inlet in fluid communication with said cavity; said housing having a fluid outlet in fluid communication with said pump cavity; said impeller urged to rotate about an impeller axis so as to cause fluid to be urged from said inlet through said pump cavity to said pump outlet; said impeller urged to rotate by impeller urging means; said impeller supported for rotational movement by impeller support means; said pump maintained at or near a predetermined operating point by control means acting on said impeller urging means; said control means receiving as input variables at least a first input variable derived from said urging means; said control means receiving at least a second input variable also derived from said urging means; said control means thereby calculating an estimate of said operating point to an approximation of predetermined accuracy relying on signals available from said urging means; said control means controlling said pump by comparing said predetermined operating point with said estimate of said operating point; and wherein said pump is arranged to operate according to a relatively flat HQ characteristic.
Preferable there is no inflexion point of said HQ characteristic at or near said predetermined operating point.
Preferably said pump is specified in a range of 100-2000 rev/min (gal/min)1/2ft−3/4.
In a particular preferred form said HQ characteristic is sufficiently flat that head will remain constant to a sufficient approximation over a predetermined operating range whereby, over said operating range whereby, over said operating range, said system can assume that pump speed will be proportional to flow rate.
Nregurg=N(t) for Qdiastole=0 L/min
Embodiments of the present invention will now be described, with reference to the accompanying drawings, wherein.
FIG. 40 shows a porous metal substrate on a housing surface.
A preferred embodiment of the invention is the centrifugal pump 1, as depicted in FIGS. 1 and 2, intended for implantation into a human, in which case the fluid referred to below is blood. The pump housing 2, can be fabricated in two parts, a front part 3 in the form of a housing body and a back part 4 in the form of a housing cover, with a smooth join therebetween, for example at 5 in FIG. 1. The pump 1 has an axial inlet 6 and a tangential outlet 7. The rotating part 100 is of very simple form, comprising only blades 8 and a blade support 9 to hold those blades fixed relative to each other. The blades may be curved as depicted in FIG. 2, or straight, in which case they can be either radial or back-swept, i.e. at an angle to the radius. This rotating part 100 will hereafter be called the impeller 100, but it also serves as a bearing component and as the rotor of a motor configuration as to be further described below whereby a torque is applied by electromagnetic means to the impeller 100. Note that the impeller has no shaft and that the fluid enters the impeller from the region of its axis RR. Some of the fluid passes in front of the support 9 and some behind it, so that the pump 1 can be considered of two-sided open type, as compared to conventional open centrifugal pumps, which are only open on the front side. Approximate dimensions found adequate for the pump 1 to perform as a ventricular assist device, when operating at speeds in the range 1,500 rpm to 4.000 rpm, are outer blade diameter 40 mm, outer housing average diameter 60 mm, and housing axial length 40 mm.
As the blades 8 move within the housing, some of the fluid passes through the gaps, much exaggerated in FIGS. 1 and 3, between the blade bearing faces 101 and the housing front face 10 and housing back face 11. In all open centrifugal pumps, the gaps are made small because this leakage flow lowers the pump hydrodynamic efficiency. In the pump disclosed in this embodiment, the gaps are made smaller than is conventional in order that the leakage flow can be utilised to create a hydrodynamic bearing. For the hydrodynamic forces to be sufficient, the blades may also be tapered as depicted in FIGS. 3A and 3B, so that the gap 104 is larger at the leading edge 102 of the blade 8 than at the trailing edge 103 thereby providing one example of a wedge-shaped restriction defined by at least one “deformed surface” as described elsewhere in this specification and a corresponding opposing surface. The fluid 105 which passes through the gap thus experiences a wedge shaped restriction which generates a thrust, as described in Reynolds theory of lubrication (see, for example, “Modern Fluid Dynamics, Vol. 1 Incompressible Flow”, by N. Curie and H. J. Davies, Van Nostrand, 1968). For blades considerably thinner than their length, the thrust is proportional to the square of the blade thickness at the bearing face, and thus in this embodiment thick blades are favoured, since if the proportion of the pump cavity filled by blades is constant, then the net thrust force will be inversely proportional to the number of blades. However, the blade bearing faces can be made to extend as tails from thin blades as depicted in FIG. 3C in order to increase the blade face area adjacent the walls.
In one particular form, the tails join adjacent blades no as to form a complete shroud with wedges or tapers incorporated therein. An example of a shroud design as well as other variations on the blade structure will be described later in this specification.
There are many profiles of bearing surface which will generate the wedge-shaped restriction. In the preferred embodiment the amount of material removed simply varies linearly or approximately linearly across the blade between the body and trailing edges. Alternative taper shapes can include a radiused leading edge or a step in the blade bearing face, though the corner in that step may represent a stagnation line posing a thrombosis risk.
For a given minimum gap, at the trailing blade edge, the hydrodynamic force is maximal if the gap at the leading edge of the blade end face is approximately double that at the trailing edge of the blade end face. Thus the taper, which equals the blade face leading edge gap minus the trailing edge gap, should be chosen to match a nominal minimum gap, once the impeller has shifted towards that edge. Dimensions which have been found to give adequate thrust forces are a taper of around 0.05 mm for a nominal minimum gap of around 0.05 mm, and an average circumferential blade bearing face thickness of around 6.0 mm for 4 blades. For the front face, the taper is measured within the plane perpendicular to the axis. The axial length of the housing between the front and back faces at any position should then be made about 0.2 mm greater than the axial length of the blade, when it is coaxial with the housing, so that the minimum gaps are both about 0.1 mm axially when the impeller 100 is centrally positioned within the housing 2. Then, for example, if the impeller shifts axially by 0.05 mm, the minimum gaps will be 0.05 mm at one face and 0.15 mm at the other face. The thrust increases with decreasing gap and would be much larger from the 0.05 mm gap than from the 0.15 mm gap, about 14 times larger for the above dimensions. Thus there is a net restoring force away from the smaller gap.
An indicative plan view of impeller 100 relative to housing 2 is shown in 2 having a concentric volute 13.
Careful design of the entire pump, employing computational fluid dynamics, is necessary to determine the optimal shapes of the blades 8, the volute 13, the support and the housing 2, in order to maximise hydrodynamic efficiency while keeping the bulk fluid hydrodynamic forces, shear and residence times low. All edges and the joins between the blades and the support should be smoothed.
The means of providing the driving torque on the impeller 100 of the preferred embodiment of the invention is to encapsulate permanent magnets 14 in the blades 8 of the impeller 100 and to drive them with a rotating magnetic field pattern from oscillating currents in windings 15 and 16, fixed relative to the housing 2. Magnets of high remanence such as sintered rare-earth magnets should be used to maximise motor efficiency. The magnets can be aligned axially but greater motor efficiency is achieved by tilting the magnetisation direction to an angle of around 15° to 30° outwards from the inlet axis, with 22.5° tilt suitable for a body of conical angle 45°. The magnetisation direction must alternate in polarity for adjacent blades. Thus there must be an even number of blades. Since low blade number is preferred for the bearing force, and since two blades would not have sufficient hearing stiffness to rotation about an axis through the blades and perpendicular to the pump housing (unless the blades are very curved), four blades are recommended. A higher number of blades, for example 6 or 8 will also work.
The cover winding 16 looks similar but the coils need not avoid the inlet tube and so they appear more triangular in shape. The body winding has a more complex three dimensional shape with bends at the ends of the body support section. Each winding consists of three coils. Each coil is made from a number of turns of an insulated conductor such as copper with the number of turns chosen to suit the desired voltage. The coil side mid-lines span an angle of about 50°-100° at the axis when the coils are in position. The coils for body and cover are aligned axially and the axially adjacent coils are connected in either parallel or series connection to form one phase of the three phase winding. Parallel connection offers one means of redundancy in that if one coil fails, the phase can still carry current through the other coil. In parallel connection each of the coil and body winding has a neutral point connection as depicted in FIG. 5A, whereas in series connection, only one of the windings has a neutral point.
An alternative three phase winding topology, depicted in FIG. 5S, uses four coils per phase for each of the body and cover windings, with each coil wrapping around the yoke, a topology called a “Gramm ring” winding.
Yet another three phase winding topology, depicted in FIG. 5C, uses two coils per phase for each of the body and cover windings, and connects the coil sides by azimuthal end-windings as is standard motor winding practice. The coils are shown tilted to approximately follow the blade curvature, which can increase motor efficiency, especially for the phase energising strategy to be described below in which only one phase is energised at a time. The winding construction can be simplified by laying the coils around pins protruding from a temporary former, the pins shown as dots in 2 rings of 6 pins each in FIG. 5C. The coils are labeled alphabetically in the order in which they would be layed, coils a and d for phase A, b and e for phase B, and c and f for phase C. Instead of or as well as pins, the coil locations could be defined by thin fins, running between the pins in FIG. 5C, along the boundary between the coils. The coil connections depicted in FIG. 5C are those appropriate for the winding nearest the motor terminals for the case of series connection, with the optional lead from the neutral point on the other winding included.
FIG. 6 depicts an alternative embodiment of the invention as an axial pump. The pump housing is made of two parts, a front part 19 and a back part 20, joined for example at 21. The pump has an axial inlet 22 and axial outlet 23. The impeller comprises only blades 24 mounted on a support cylinder 25 of reducing radius at each end. An important feature of this embodiment is that the blade bearing surfaces are tapered to generate hydrodynamic thrust forces which suspend the impeller. These forces could be used for radial suspension alone from the straight section 26 of the housing, with some alternative means used for axial suspension, such as stable axial magnetic forces or a conventional tapered-land type hydrodynamic thrust bearing. FIG. 6 proposes a design which uses the tapered blade bearing surfaces to also provide an axial hydrodynamic bearing. The housing is made with a reducing radius at its ends to form a front face 27 and a back face from which the axial thrusts can suspend the motor axially. Magnets are embedded in the blades with blades having alternating polarity and four blades being recommended. Iron in the outer radius of the support cylinder 25 can be used to increase the magnet flux density. Alternatively, the magnets could be housed in the support cylinder and iron could be used in the blades. A slotless helical winding 29 is recommended, with outward bending end-windings 30 at one end to enable insertion of the impeller and inward bending windings 31 at the other end to enable insertion of the winding into a cylindrical magnetic yoke 32. The winding can be encapsulated in the back housing part 20.
FIG. 10 is a side section, indicative view of the impeller 204 defining the orientations of central axis FF, top taper face DD and bottom taper face BB, which tapers are illustrated in F in side section view.
In order to provide both radial and axial direction control at least one set of surfaces must be angled with respect to the longitudinal axis of the impeller (preferably at approximately 4 5° thereto) thereby to generate or resolve opposed radial forces and an axial force which can be balanced by a corresponding axial force generated by at least one other tapered or deformed surface located elsewhere on the impeller.
In the forms thus far described top surfaces of the blades 8, 207 are angled at approximately 450 with respect to the longitudinal axis of the impeller 100, 204 and arranged for rotation with respect to the internal walls of a similarly angled conical pump housing. The top surfaces of the blades are deformed so as to create the necessary restriction in the gap between the top surfaces of the 7 blades and the internal walls of the conical pump housing thereby to generate a thrust which can be resolved to both radial and axial components.
A further modification of this arrangement is illustrated in FIG. 19 wherein impeller 304 includes secter-shaped blades 305 having curved leading and trailing −3 8 portions 306, 307 respectively thereby defining channels 308 having fluted exit portions 309.
It is to be understood that, whilst the example of FIG. 21 shows the surfaces of the shroud 411 angled at approximately 450 to the vertical, other inclinations are possible extending to an inclination of 00 to the vertical which is to say the impeller 410 can take the form of a cylinder with surface rippling or other deformations which impart the necessary hydrodynamic lift, in use.
Characteristics and advantages which flow from the arrangement described above and with reference to the embodiments includes.
In this instance the estimation and control system 10 operates on and receives sensor feedback from pump assembly adapted for implantation in human body 12 and arranged to operate in parallel across at least a part of heart 13 so as to at least assist if not fully take over the pumping function of heart 13.
Curve fitting of this plot produced the equation Q=20.29+4.73 ln(Pin)−55√(n) where Q is flow rate in L/min, Pin is electrical input power to the motor in Watts and n is motor speed in rpm. The maximum error for this prediction was 4% for the combined data. Pressure head across the pump was described by the relationship ΔP=−13.68−6.59 ln(Pin)+2.18e-5 (n)2 with equivalent accuracy. Two different rotor designs have been tested in this manner to date both yielding similar accuracy curve fits of the form Q=a+b·ln(Pin)+c·√(n) and of the form ΔP=a+b·ln(Pin)+c·(n)2.
The reasons for low error in prediction given change in viscosity are postulated as follows: Firstly that the “flat” H-Q curves for this pump give small variation in pressure head for given flow rates. Secondly the nature of the hydro-dynamic hearing. Although the pump has relatively high disc friction forces, which tend to be most sensitive to viscosity changes, the rotor in this case conserves energy by repositioning in free space according to the fluid viscosity. Thirdly, the size, where surface roughness is relatively smaller than for smaller higher speed pumps. Fourthly, allowing speed to vary around a set point due to choosing a comparatively long time constant.
The measured motor efficiency is between 45% and 48% curves, for speeds between 2000 rpm and 2500 rpm and motor output power between 3 and 7 W. For example, at 2250 rpm and 3 W motor output (roughly rated conditions), the copper loss was 1.7 W, the eddy loss in the titanium was 1.0 W. and the iron loss in mild steel yokes was 0.7 W, giving a motor efficiency of 47%.
With reference to FIG. 32 the example 1 can be applied to the preferred embodiment of FIGS. 7 to 15 comprising pump assembly 200 incorporating an estimating and control system of the type described with reference to FIGS. 28 to 31.
In the forms thus far described top surfaces of the blades 8, 207 are angled at approximately 45° with respect to the longitudinal axis of the impeller 100, 204 and arranged for rotation with respect to the internal walls of a similarly angled conical pump housing. The top surfaces are deformed so as to create the-necessary restriction in the gap between the top surfaces of the blades and the internal walls of the conical pump housing thereby to generate a thrust which can be resolved to both radial and axial components.
By increasing the smallest radius from the centreline to the blades (i.e. to the nose of the blades) at the top and not at the bottom of the impeller, an axial thrust force can be imposed on the impeller toward the bottom. This arrangement can be carefully designed so as to bias the load to the bottom bearing and relieve the top bearing which is more highly loaded (in that it must resist both axial and radial loads).
The point at which the aortic valve just remains closed is the point of total assist given the name OCA (optimal cardiac assistance) This is the point at which minimum head pressure across the pump begins to rise with increasing pump speed. In other words during systole the left ventricle peak pressure begins to decrease as average pump speed is increased.
Therefore pumping at the point of optimal cardiac assistance and avoiding over pumping, the control algorithm should maintain minimum pump speed such that the minimum head pressure across the pump does not increase. Therefore the new desired set point Nnew to hold the optimal cardiac assistance point can be defined by the old speed value Nold reduced by a factor proportional to the increase in minimum systolic head pressure (ΔHsys) beyond the minimum possible head pressure (ΔHmin) Kp is the proportional constant. This is described by equation 1.
Nnew=Nold−[Kp*(ΔHsys−ΔHmin)] equation 1
Nregurg=N(t) for Qdiastole=0 L/min equation 2
In addition appropriate coatings and/or structure materials will be used on the respective rotor 500 and at least inner walls or the housing to minimise damage and/or damaging effects arising from a touchdown.
Further particular preferred coating arrangements are as follows.
Application of coatings to the Ti-6Al-4V substrate of a blood pump.
Also coatings applied using plasma immersion ion implantion of nitrogen, titanium nitride and carbon, or combinations of these treatments, for enhancement of hardness, improved elastic recovery under impact conditions, low friction and high wear resistance.
Overall the desirable characteristics to be achieved by the coatings are.
An alternative approach is to allow tissue overgrowing the bearing substrate to act as the EHD component. The substrate may be a porous metal substrate such that it allows a pseudoneointimal cell lining to grow into the pores on the substrate surface. It is commonly reported that the pseudoneointimal lining thickness is stable and around one cell deep. The advantage of using a bio-EHD” component on the surface is that damage to the EHD component may regenerate within a few hours of damage occurring. Possible drawbacks may be a tendency for a bio-EHD component to sustain damage under relatively low shear stresses commonly seen in a rotary blood pump and for sections of bio-EHD to be stripped away forming potentially dangerous embolii. This may be countered by additional surface treatments which promote stability of the pseudoneointima. Once again, the fundamental shape of the bearing, that is the “deformed surfaces” should remain substantially the same as for the hydrodynamic bearings of the blood pump embodiments previously described. FIG. 40 shows a porous metal substrate 700 and a pseudoneointimal layer 702 over a housing surface 701.
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