JP2003501155A - Magnetic levitation supported blood pump - Google Patents

Abstract

SUMMARY OF THE INVENTION The present invention is a non-contact axial flow turbopump for propelling blood, comprising a pump housing (24) defining an axis of the pump and having inlet and outlet openings at opposite axial ends. And a rotor unit (17) for defining a rotor shaft and a tip of the opposed rotor shaft. The pump magnetically levitates and supports the rotor in the pump housing at both rotor shaft tips, avoids physical contact with the housing, and has the rotor shaft ends and the magnetic levitating support elements (13) (13 '). To define a fluid gap.

Description

Detailed Description of the Invention

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to rotary blood pumps that are implantable in the human chest and may be used to assist the human heart in pumping blood. , In particular to blood pumps of this type using magnetic levitation support.

Description of the Prior Art An implantable blood pump based on the current technology currently being developed to assist the heart is a turbo pump. They are axial flow type such as JARVIK 2000, centrifugal type developed by Cleveland Clinic, and mixed type such as "Streamliner" developed by University of Pittsburgh. is there. They all use high speed rotating impellers that rotate at speeds of thousands of rpm. Many, including JARVIK 2000, use journal bearings with solid contact to support the rotor. The use of such a bearing is not preferable because it may cause blood damage or thrombus. To avoid the problem of solid contact bearings Recently, magnetic bearings have been used, such as "streamliner" pumps. These magnetic bearings are non-contact bearings and can maintain large bearing gaps to reduce shear stress in blood, thus minimizing damage to blood. However, there is still the problem of having to flush the bearing gap with fresh blood. This washout is important to eliminate thrombus formation.

Magnetic bearings must be packaged in a small space to minimize the size of the pump, which is difficult to achieve. Many magnetic bearing pumps are too large to be widely accepted.

Another requirement for implantable blood pumps is low power consumption. Pumps using magnetic bearings are known to consume large amounts of power, requiring as much as 20 watts for bearings alone (typically around 5 watts). The power delivered to the blood in the left ventricular assist device (LVAD) is about 3
It is not desirable to dissipate more power than watts, and thus more than 1.0 watts for magnetic bearings.

Many magnetic bearings use permanent magnets or electromagnets to create a radial magnetic field to directly levitate the rotor in a radial direction. However, the radial "stiffness" of the bearing obtained with a relatively small air gap field, such as that produced by a magnet, is not high. Therefore, large bearings are required to support the added forces with only small radial displacements.

As is known as Ernshaw's law, radial passive magnetic bearings are basically axially unstable. Therefore, active axial control is required to stabilize a rotor levitationally supported by such bearings. In particular, in an axial turbo pump in which a considerable axial force is applied to the rotor, the power consumed by the active coil can be excessive. A “virtually zero power (VZP)” control loop may be utilized to reduce power consumption. This control, commonly known as the VZP control, was first used in the 1970s by J. Lyman, one of the founders of the magnetic levitation support technology.

Implantable turbo blood pumps usually operate at a constant rotational speed.
This is because it is difficult to close the loop around the patient and physiologically change the flow rate of the pump according to the patient's needs. By providing a basic flow rate, the increased demand for blood flow as a function of activity level is supplemented by the patient's own heart. However, a diseased heart cannot meet the great demand and thus limits the patient's activity level. Compensating the demand for cardiac output more or less with the patient's heart would place an undesirable load on the patient's left ventricle. In order to physiologically control the output flow of the pump, additional sensors are often used to measure physiological parameters of the patient. As such, there is a blood pressure transducer for measuring the outlet pressure of the pump and the differential pressure of the pump. The introduction of such an additional sensor is extremely undesirable because it may cause a thrombus or cause long-term hemodynamic reliability problems. Since some known LVADs cannot directly measure their output blood pressure, they measure the flow rate of the pump using a series of ultrasonic flow meters arranged to penetrate the body.

The natural heart produces a pulsating flow. Experiments have shown that such inconsistent flow minimizes thrombus development in large arteries of the human body, since the flow pattern is constantly changing. In a pulsatile pump, stagnation of flow is minimized or completely eliminated not only in the patient's artery at the pump outlet, but also within the pump itself. Current turbo pumps have a direct current (
DC), that is, a constant flow pump device. Even if the patient's heart recovers and is able to provide some degree of pulsation to the human body, the LVAD blood pump reduces the load on the patient's heart, so the degree of pulsation is , Much smaller than when using only the natural heart. “Bridge To Recovery”
For long term implants as such, it is highly desirable to be able to provide pulsatile flow from the LVAD.

SUMMARY OF THE INVENTION Accordingly, it is an object of the present invention to provide a novel means of flushing the magnetic bearing gap with fresh blood to eliminate thrombus formation in the bearing.

Another object of the invention is to allow the bearing to be flushed even when the blood flow through the pump is low.

Yet another object of the present invention is to provide a non-contact active flushing means for bearings.

Yet another object of the present invention is to provide a magnetic bearing structure that can be easily washed away by the bloodstream to prevent the formation of stasis areas.

A further object of the present invention is a compact bearing system of simple construction that can be packaged with various types of turbopumps for use by both adults and children. Is to provide.

A further object of the present invention is to provide a control system that consumes very little power even when used with high load bearings such as those disclosed herein.

A further object of the invention is to be able to directly measure the differential pressure of the pump without the addition of a separate sensor.

A further object of the present invention is to provide such an active coil and magnet structure which consumes little power to carry axial loads.

A further object of the present invention is to ensure the safety in pulsating flow formation so as to avoid an unfavorable state where blood flows backward in the pump.

A further object of the present invention is to provide reliable pulsating flow using a pump differential pressure determined directly by magnetic bearings.

A further object of the present invention is to reduce the length of an Archimedes screw axial flow impeller by providing a plurality of parallel flow blades with minimal overlap. It is in. For small size blood flow pumps as contemplated by the present invention, minimizing the axial length of the pump is desirable, especially in applications for women and children.

A further object of the present invention is to provide a stator having a compact short axial length which does not damage blood during pressure recovery in the impeller.

In order to achieve the above object and other objects as will be apparent from the following description, a blood pump according to the present invention defines a pump axis and has an inlet opening and an outlet opening at opposite axial ends. Includes pump housing. A rotor is provided that defines a rotor axis and opposing rotor axis ends. During the operation of the pump, the radial stability of the rotor is passively maintained so that the rotor axis is substantially aligned with the pump axis, and the rotor is magnetically levitated and supported in the pump housing. Magnetic levitation support means is provided at the end of the rotor axis. Control means are provided for maintaining axial stability of the rotor so that the rotor can absorb externally applied loads and avoid contact of the rotor within the pump housing. It The impeller means provided in the rotor functions to suck blood from the inlet opening and discharge it from the outlet opening as the rotor rotates. A drive means is provided for pumping blood by rotating the rotor and impeller means, forming a fluid gap between the rotor axis end and the magnetic levitation support means. Furthermore,
Blood flushing means are provided for continuously flowing blood into the fluid gap during rotation of the rotor to prevent thrombus formation in the fluid gap.

Suitably, the flushing means actively or actively flushes blood into the fluid gap where blood stagnation may occur. According to another feature of the invention, the drive means drives the impeller means at a selected rotational speed to detect a differential pressure within the pump housing and to provide a periodic variation in the selected rotational speed of the drive means. Means are provided for providing a pulsatile flow of blood that feeds and flows through the pump into the patient's circulatory system. Also, one structure for flushing relies on using the flow itself to passively create a differential pressure applied to the bearing gap. In another construction, Archimedean screws are attached to the front and back of the rotor to actively pump blood through the bearing gap. Approximately 30% of patients receiving cardiac support show full recovery of the natural heart after 1 or 2 years of LVAD use, eliminating the need for a pump.
Rather than removing the device, it is possible to leave the pump in place and operate with minimal flow and power consumption. Active screw pumps allow proper flushing of the bearing gaps even when the pump is in a sleep state and used minimally.

An important feature of the present invention is the provision of impeller means on a magnetically levitated rotor which is inherently stable in the radial direction, as well as direct feedback useful for axially stabilizing the rotor. Is to provide a signal. By this feature,
The pump design and construction is greatly simplified, manufacturing costs are reduced and reliability is significantly improved over a long period of use.

A small, high radial stiffness magnetic bearing uses an axial peripheral ring field to passively support the rotor of the pump in the radial direction. This magnetic field is focused or concentrated from the permanent magnets into the peripheral ring and, in a small size, provides a very large radial load capacity. This is unlike conventional radial passive magnetic levitation supports that use radial magnetic fields. Active axial control uses a "virtually zero power" control feedback loop to stabilize the bearing. The low power consumption and small size make it possible to apply this bearing to blood pumps of the axial flow type or other structures, especially suitable for implants. The differential pressure across the fluid gap of the bearing forces, positively or actively flushes the gap with fresh blood and prevents thrombus and flow stagnation. The rotor force on the magnetic bearing can be measured by the bearing control system. This allows a direct determination of the differential pressure across the pump. Using this parameter to obtain a pulsating flow,
Physiological control can be added to the pump output to suit the activity level of the patient.

The present invention provides a bearing with a very high radial stiffness. This is accomplished by using an axially oriented peripheral ring magnetic field that has a greater radial load capacity than the radially oriented magnetic field. This allows the use of small diameter bearings that were previously impossible. High load capacity also reduces power consumption.

Prior to the present invention, it was not possible to directly determine the differential pressure of a turbopump. The differential pressure of the turbopump can be used, as an example, to add physiological control to the pump flow rate to meet the needs of the patient's activity level and heart rate.

The invention will be explained more concretely in the following detailed description with reference to the accompanying drawings on the basis of preferred embodiments, together with further objects and advantages.

DETAILED DESCRIPTION OF THE INVENTION In the following description of the axial flow and centrifugal pumps according to the present invention, the structure of the magnetic bearing and other common elements such as the flow path and the magnetic circuit are the same. It is a thing. These common elements are designated by the same or the same reference numerals.

[0029]   FIG. 1 shows the pump axis A_PRepresents an axial flow pump P which defines
Axis A_rA rotor R, and a pump axis A as will be described in detail below.
_PMultiple spirally curved impeller blades mounted to rotate around
Have a card 4. The impeller blade 4 is attached to the rotor housing 17.
The housing is made of thin-walled titanium that is highly compatible with blood. fluid
All parts that are wetted by titanium consist of titanium or other suitable non-magnetic material. Rotor
At each axial end of R, a magnetic bearing pole piece 12 made of a ferromagnetic material such as iron,
12 'is provided. Annular permanent magnets 13, 1 magnetized in the axial direction
3'between the corresponding pole pieces 12, 12 'and the axial end walls 30, 30' of the rotor R
Are arranged coaxially. The rotor R includes the pole pieces 12, 12 'in the rotor housing.
It consists of a thin shell of titanium in the shape of a cylinder that seals inside 17.
Has been. The rotor housing 17 has no physical contact with the fixed part of the pump.
There is nothing to do.

[0030]   Magnetic bearings B and B'are provided at each axial end of the rotor R, one side of which is magnetic.
Fixed with stones 13 and 13 'and pole pieces 12 and 12' on the other side made of iron
It consists of yokes 28, 28 '. The path of the magnetic flux 50, 50 'is shown by the wavy outline.
And forms a closed loop in the plane shown. The magnetic flux is a fluid flow
Pass through the top 20, 20 'in the axial direction. York has an outlet fairing
It is sealed within the ring 6 and the inlet fairing 29.
Each fairing has a yoke 28, 28 'and an active bearing near the rotor R.
Having annular regions 6a, 29a for accommodating the coils 14, 14 ', each annular region
Are sealed by the walls 6b, 29b, so that the axial direction of the corresponding rotor
Cooperate with the end walls 30, 30 'to define a radial fluid gap 20, 20'.
The walls 6b, 29b have a thickness of about 0.010 inch and are the dimensions of the non-magnetic air gap.
Law G_aIt is set thin to minimize It is best shown in Figure 5.
As such, the fluid gap 20, 20 ′ will shear the blood depending on the maximum rpm of the rotor.
Approximately 4 mils to 30 mils so that the stress is below the hemolytic level (subhemolytic)
Range G_fhave. Fluid gap at both axial ends of rotor R
20, 20 'are generally chosen to be approximately equal, but as shown in FIG.
In addition, the non-magnetic air gap may be different. In the case of FIG. 15, the upstream side
Air gap G_aIs the air gap G on the downstream side_aGreater than ′, both fluid gap G _f , G_f′ Is the wall W₁~ W_FourMaintained to be approximately equal by changing the thickness of
Has been done. It is also clear that the wall thickness is the same and the fluid gap 20 is
While still doing the same, physically move the pole pieces 12, 12 'to one of these pole pieces or
Separate multiples from corresponding walls to equalize air gaps if desired, or
Different to act on the rotor during normal pump operation.
Similar results can be obtained by compensating for the average force that occurs.

In order to prevent flow separation or turbulence, the inlet fairing 29 is tapered outwardly in the radial direction at the portion indicated by reference numeral 29c to provide a streamlined shape. I am doing it. These are the entrance guide vanes 33
It is attached to the pump housing 24 via. These vanes are usually straight, but may be spirally curved. Four curved vanes 33a-33d are shown in FIG. A space 31 is provided between these inlet vanes in the circumferential direction. An end view of the entrance fairing 29 and its pointed tip 32 is shown in FIG. Similarly, the streamlined exit fairing 6 tapers radially inwardly at a portion 6c, with several evenly spaced rectifying vanes 5 extending the exit fairing 6 It is held in a central position with respect to the pump housing 17. A radial gap 8 is provided between the vane 5 and the rotor housing 17. FIG. 3 shows four straightening vanes 5 a-5 d, which are circumferentially spaced apart 31 ′, and the sharp end 7 of the outlet fairing 6.

Referring again to FIG. 1, the yoke 28 and the pole piece 12 are coaxial with each other in the axial direction at the outer peripheral portion thereof, as shown in the part showing the structure of the two magnetic bearings. And two thin peripheral rings 10, which are radially aligned with each other,
10 'and 11, 11' are provided. There may be one or more such peripheral rings. The radial stiffness and load capacity increase with the number of rings given a constant magnetic flux density in the air gap. The axial magnetic field between the corresponding opposite peripheral rings is considerably stronger than the magnetic field of the permanent magnets 13, 13 '. This is because the magnetic flux concentrates across the air gap of a narrow or radially thin peripheral ring having a smaller cross-sectional area than the magnets 13, 13 '. This produces a large passive radial load capacity or restoring force. However, the axial magnetic field produces unstable axial stiffness or force when the rotor is not centered axially. This occurs when there are unequal air gaps at each end. The load on the rotor while pumping blood creates an axial force, which must be counteracted so that the rotor does not come into contact with the fairings 6, 29.

The transient counter force is provided by the active bearing coils 14, 14 'which surround the inner diameter of the yoke. When an electric current flows through the coil, a magnetic field is formed in series so as to overlap the magnetic field of the magnet, and the magnetic field generated by the magnet is changed. When viewed in the direction of arrow A in FIG. 1, the clockwise current in the upstream bearing coil 14 increases the magnetic flux and the counterclockwise current decreases it. In this way, by increasing the magnetic flux in the air gap to increase the force, the upstream bearing B can pull the rotor R to the right in FIG. The downstream bearing can draw the rotor R to the left with a smaller force by passing a current in a direction that reduces the magnetic field. This results in a rightward positive force as a net result. By reversing both currents, a net leftward force can be generated.

By using a closed loop control system that controls the coil current based on the deviation signal of the axial position of the rotor, the axial movement of the rotor can be suppressed within an allowable range. However, when an external force is applied to the rotor, a certain amount of electric power is consumed in the coil to counter the external force.

In this bearing structure, by using a control method other than position detection, it is possible to counteract or neutralize an external force that is steady or gradually changing by the permanent magnet 13 instead of the coil. Therefore, such a problem can be avoided. If the rotor can be displaced axially with a suitable distance, the axial instability of the bearing can counteract external forces. The coil need only stabilize its axial position with a substantially zero power (VZP) feedback loop. What this control loop does is to drive the rotor axial velocity to zero and the coil current to zero in a separate control loop. When the axial position is stable under load, the peripheral axial force balances the load and the rotor does not displace in any axial direction. That is, the axial velocity of the rotor is zero and there is no need to supply any DC current to either coil to support the DC load.

To implement the VZP control method, an axial velocity signal of the rotor is required, which can take advantage of the back emf generated in at least one coil 14. This is because the back electromotive force generated in the coil 14 is directly proportional to the axial velocity of the rotor and is stable in the direction with respect to the long term. Therefore, no additional sensor is needed to determine the axial velocity.

In order to initially support the bearing in axial contact, the opposing coil will provide sufficient force to pull the rotor apart, since the axial force at the contact end is generally large. Can not occur. To solve this problem, the initial D
A C current is temporarily passed through each coil to generate a magnetic flux that opposes the magnetic flux from each magnet. This allows the axial instability to be minimized, the coils to be current controlled and the rotor to be centered. When the rotor is centered, the current to repel the magnet can be reduced to zero.

It is also possible to initially support the bearing in a floating manner by using a position control loop which replaces the VZP loop. This requires an axial position signal. This can be obtained by measuring the inductance of the coil by superimposing a low level AC current. Inductance is inversely proportional to the axial gap of the magnetic circuit. Therefore, the coil 14 can also be used as a position sensor. Once the bearing has floated, control can be switched to VZP mode. This position control mode may be selected periodically throughout the life of the pump, which
The long-term stability of the axial position can be maintained. This also serves as a backup of VZP control and a function of confirming it.

The axial pump in FIG. 1 uses an impeller blade 4 with a constant diameter similar to that in JARVIK 2000. This constant diameter impeller is ideal for use with the magnetic bearings disclosed herein. Because the rotor is VZP
This is because, as indicated by the loop, it can be axially displaced at any distance within the housing of constant diameter to support the load. Sufficient axial gaps 25, 26 and 20, 20 'are used. In contrast, a mixed flow turbopump uses an impeller that is radially tapered over its length. The impeller is located in close proximity within the tapered housing bore. Therefore, the amount of axial displacement without contact with the housing is much more limited, which limits the axial force capacity of the bearing in the VZP mode.

The most important objective in the design of blood pumps is to eliminate fluid stagnant areas to avoid thrombus formation. A stagnant part of the flow where the flow velocity becomes zero (
In particular, thrombus formation can be initiated in the surface) or in a portion where the flow velocity becomes almost zero due to backflow. This has been a problem for many years in magnetic levitation bearings. This is because the complete floatation of the rotor creates fluid gaps, gaps or passages that are difficult to flush with fresh blood.
This problem is solved according to the invention by forcing the blood flow through these areas. Referring to FIG. 1, blood is introduced into the pump along the direction of arrow A. Blood is introduced at a pressure close to the stagnation pressure, i.e. under high pressure, through a straight or tapered hole 1. Blood flows to a deflecting means or deflector 16 consisting of a sharp pin and is directed radially from there through a fluid gap 20, which is the gap separating the rotor and stator. This flow is also sucked out at the outer periphery of the gap indicated by reference numeral 15 by the suction generated at the inlet of the rotary impeller 4 located slightly downstream of the gap. This differential pressure causes a forced flow that flushes the gap 20.

The gap 20 'at the outlet or downstream end of the rotor is flushed with the aid of a small circumferential lip or scoop 9 having a radial height H approximately equal to the gap 20'. The scoop 9 sends a part of the high-speed flow downstream of the impeller blade 4 in the radial direction toward the gap 20 ′. In Scoop 9,
There is a high stagnation pressure compared to the outlet pressure of line 19. This differential pressure actively flushes the bearing gap 20 '. The exit cone with the sharp focus point 18 causes the flow through the radial gap to join the axial conduit 19 with almost no stagnation points due to the sharp edge 18.
To further reduce stagnation at the tip 18, the tip may be offset slightly with respect to the center of rotation or pump axis. This avoids zero speed. The pointed deflector 16 at the inlet can also achieve the same effect by being slightly offset from the pump axis.

In the disclosed structure, the bearing gap, ie, the region of backflow in the fluid gap 20, 20 ′, does not occur because of the unidirectional differential pressure in these left and right gaps. This causes a forced flow for unidirectional flushing. In this way, since there is no stagnation region, the probability of thrombus occurring is reduced. The rotating rotor R also functions like a low efficiency centrifugal pump and exerts a radially outward pumping action in the gaps 20, 20 '. However,
This effect is much smaller than the forced pressure flow so that no backflow occurs in the rear gap 20 '. In the front gap 20, radial flow is enhanced by centrifugal action. As shown in FIGS. 16 and 17, the efficiency with which the blood flows radially outward in the upstream gap 20 is determined by the appropriate shape of the outer surface of the axial end wall 30 in the radial direction. It can be improved by providing the fins F. These fins F project into the gap 20 at a fraction of their width. Its size is on the order of a few mils. Such a fin F is
Actively transports blood radially outward at the inlet gap 20. Those skilled in the art are familiar with the size, shape, orientation and / or location of fins that produce a particular good result.

An important feature of the present invention is the unique flush structure that creates the differential pressure and the magnetic bearing structure with the ability to generate active forces in both directions. Rotary motor structures using brushless motor technology are well known in the art and will not be described in detail here. Permanent magnets 21 are shown in dashed lines in the area 21 in the rotor.
It is located in. The magnetic field of the rotor magnet interacts with the rotating magnetic field generated by the plurality of stator coils 22 surrounding the pump housing, and generates rotating torque as the coils are commute. Further detailed description is omitted. However, the secondary load generated by the rotary motor should be minimized so as not to add unwanted load to the magnetic bearings.

The bearing structure disclosed in the present specification has two bearings that are spaced apart from each other in the axial direction.
It has two separate magnetic bearings B, B '. Such axial division provides the floated rotor R with stability against moment or cocking. The center of gravity 23 of the rotor is ideally located approximately at the midpoint of both bearings, as shown. This makes it possible to equalize the radial load applied to the bearing due to the weight of the rotor or the impact load. Also, the radial force of the motor is equally distributed.

The preferred embodiments of the axial flow and mixed flow type turbo pumps have been described above. Next, how the same bearing design and flush differential pressure can be applied to the centrifugal turbo pump will be described. In FIG. 5, a typical centrifugal pump comprises a fluid outlet opening 27 ′ arranged tangentially to the volute housing 49. A bearing housing 46 having a smaller diameter houses the two magnetic bearings that support the rotor. The interior of the pump is shown in FIG. Here, only the lower part of the centrifugal impeller blade 39 is shown in cross section. The upper blades are shown by taking a cross section between the blades. The magnetic bearings arranged at a distance from each other have the same basic structure as that described above, but have one difference as a modified embodiment. That is, the annular magnet 13 is arranged in the yoke of the stator instead of the pole of the rotor. In this way, the magnets do not rotate and are not subject to centrifugal forces. The rotating iron pole piece 12 and the rotor R'can be more perfectly balanced and the structure can be simplified. In this example, the magnetic bearing coil 14 directly surrounds the magnet, but the magnetic flux of the magnet can be suitably changed as described above.

In FIG. 6, the direction of fluid flow is indicated by arrow A. Blood strikes the inlet fairing 29 and is introduced into the line 1 as in the case of axial pumps.
Blood is deflected radially by the sharp pin 16 and flows into the bearing fluid gap 20. The static fluid pressure at edge 32 is higher than the pressure at the outlet of passage 20 due to its low flow velocity. This is achieved by making the cross-sectional area of the inlet passage 47 larger than the cross-sectional area of the annular passage 44, as can be calculated according to the Bernoulli definition. The gap 20 is forcibly flushed not only by such pressure difference, but also by the suction produced in the gap 20 by the centrifugal impeller 39 having a plurality of radial blades. The fairing 29 is connected to the pump housing 46 by evenly arranged thin vanes 37. The right side of the pump outlet, that is, the downstream magnetic bearing yoke 28
6 is centered and held with respect to the housing. Exit vane 36
May be straight or spirally curved for optimal fluid introduction towards the inner diameter of the impeller vanes. The fluid is pushed outward in the radial direction by the impeller vanes 39 and introduced into the spiral chamber 45 extending in the circumferential direction. The swirl chamber 45 communicates with the fluid outlet opening 27 '.

The right or outlet bearing axial gap 20 is flushed by using the scoop 9 to guide a portion of the inlet flow radially into the gap. The scoop 9 extends above the titanium rotor housing 17 with a height H, as described above for the axial pump. The scoop 9 has suction due to the impeller, but there is also equal suction at the base of the impeller blade 39 in the centrifugal housing 49. Therefore, this suction does not create a flow through the slot 20. In the scoop 9, there is a high fluid pressure close to the stagnation pressure. At the solid center of the impeller 39, at the trailing edge of the gap leading to the gaps 43 and 43 ', the gaps 43, 43' are washed away by the centrifugal force caused by viscous drag in the gap. This creates an additional negative pressure in line 19. Such a pressure differential relative to the stagnation pressure in scoop 9 causes a flow in gap 20. The flow diverter 18 achieves a smooth transition from radial flow to axial flow without creating a stagnation point. The tip portion 18 can be disposed slightly offset from the center to eliminate stagnation at the tip portion. Thus, the asymmetrical flow created by the off-centered tip in conduit 19 creates a swirl that flushes housing 49 adjacent conduit 19. In the absence of such swirl, stagnation of flow can occur within housing 49. The plurality of passages 41 in the impeller shaft 38 allow communication between the gap 20 and the conduit 19. The impeller shaft 38 has a radial clearance 42 with respect to the fixed bearing, which clearance is large enough to avoid shear stress below the hemolytic level in the blood. The gaps 20, 43, 43 'have similar sizes. The shaft 38 is preferably made of titanium and seals the pole pieces 12 while at the same time supporting them. All blood contacting surfaces preferably consist of titanium both in the design of the axial pump and in this design. Titanium increases biocompatibility,
Even if it has a biolite carbon or wear resistant titanium nitride coating or other coating so that it will not burn when it comes into contact with the bearing surface due to bearing failure or excessive shock load Good.

In the magnetic levitation bearing disclosed in the present specification, the levitationally supported rotor is
The housing can be touched down or touched in the event of a control system failure. This provides a design fail-safe that allows it to continue functioning until the failed component is replaced. When starting the bearing, the axial control system is first activated to avoid contact. The rotor is then brought to the desired rpm.

As described above, the magnetic bearings B and B ′ are inherently unstable in the axial direction. Therefore, a VZP coil control system is needed. This instability is preferably used to determine the axial force acting on the rotor. The differential pressure acting on the rotor is calculated directly by dividing this force by the effective area of the rotor. This is Figure 8
Is shown in the block diagram of FIG. The axial bearing position is monitored, such as by measuring coil inductance, at the portion indicated by 60 to determine the associated force. By dividing the force obtained by the area of the rotor at 64, the desired differential pressure DP is calculated. A lookup force table that is useful for non-linear data is not needed in practice. This is because the bearing force is sufficiently linear if a large peripheral ring blood gap is used. FIG.
Please refer to. To obtain a transient force, the contribution of the rotor's inertial mass to the force is subtracted to obtain a net external force.

As mentioned above, the inductance of the coil 14 in FIG. 1 can be used to monitor the axial position of the rotor. The coil electronics for determining rotor position can be simplified by using a separate small auxiliary position coil 54 shown in FIG. Ideally, the coil 54 is wound under the main bearing coil 14 as shown in FIG. Since only one position coil is required, FIG. 7 shows one position coil 54 arranged in the inlet bearing B, but another redundant position coil 54 'is arranged in the outlet bearing B'. Good. In comparison, it is not desirable to integrate a separate non-contact ultrasonic or magnetic axial position sensor with the pump. Once the differential pressure of the pump is known, this information can be used to provide the basis for physiological control of the pump or to provide accurate pulsatile flow, depending on the activity level of the patient.

In the state of being implanted in a person, the pump differential pressure can be varied with this frequency by periodically changing the pump rpm with a predetermined frequency. Since the gauge pressure at the inlet of the pump when drawing blood from the heart is relatively constant and low (several millihg), the differential pressure provides a good measure of the higher pump overpressure. The fluctuating pump rpm changes the pump flow instantaneously,
This change changes the outlet pressure to the human body, for example when pumping blood into the aorta. The differential pressure of the pump is monitored by the bearing. The error in the magnitude of the pump differential pressure is corrected by the feedback loop, which fine tunes the pump rpm until the desired pressure variation is achieved. This is shown schematically in the control system of FIG. In FIG. 9, a desired periodical differential pressure fluctuation is input, and a frequency is set in a portion indicated by reference numeral 66. A frequency discriminator 68 separates pump pressure differential fluctuations from pressure fluctuations caused by different frequency natural heart beats. The magnitude of the differential pressure generated across the pump is compared in comparator 70 with the set point differential pressure. The deviation signal at the output of the comparator is input to a proportional amplifier 70 having a selected gain K,
This controls the rpm 74 of the drive motor, which will reflect variations in the axial position of the rotor at 76 due to the associated pressure changes.

A typical or normal set point pressure variation is 120/80. This means that during the pump cycle the blood pressure reaches a maximum of 120 mmhg (systole) and a minimum of 80 mmhg (diastole). By being able to achieve accurate systolic and diastolic pressures, backflow in the pump is avoided, eliminating the need for bioinvasive flow meters and other sensors. The proportional gain indicated by "K" provides suitable sensitivity in control to achieve stable and accurate aortic pressure.

In this way, physiological flow control of the pump is possible based on the differential pressure. For example, if a patient's exercise level is increased, the peripheral resistance to exercise is reduced, resulting in a decrease in his or her mean blood pressure (unless cardiac output is increased). Therefore, the average rpm of the pump is simply increased in the feedback loop until the desired average blood pressure, which is typically 100 mmhg, is maintained. This automatically increases the LVAD flow rate and maintains a constant mean aortic pressure.

The heart rate of the patient can also be used as a parameter to set the mean aortic pressure and the desired systolic / diastolic ratio. The patient's own heart rate can be electronically monitored, as is done in pacemaker devices. However, it is not preferable because it makes it necessary to introduce electrodes for monitoring the heart. According to the invention, the heart rate of the patient can be directly monitored by analyzing the frequency component of the force exerted on the magnetic bearing. If the periodic changes in pump rpm forming the pulsating flow are made at a frequency slightly different from the natural heart rate, the frequency of the heart's blood pressure can be extracted from the total differential pressure signal. In that way, the average pump output pressure can be made proportional to the heart rate of the patient's heart, which is a known physiological function. A feedback loop similar to that shown in FIG. 9 can be used.

In this way, it is possible to monitor the output pressure of the natural heart and monitor what kind of pumping action the natural heart is performing and recovering over a long period of time.
Gradually reducing the contribution of blood pumping by LVAD, if desired, can ultimately restore the patient to a condition where LVAD is no longer needed. The obtained heart and pump pressure data can be used to monitor functional status. This is a great advantage for devices that are implanted over a long period of time.

1 and 7 are still shown in cases where it is desired to significantly reduce the flow rate through the LVAD, such as when the patient may recover to the point where the LVAD is no longer needed. Such flushing of the bearing gap 20 must be maintained. Referring to FIG. 7, a spiral deep groove screw pump is attached to each end surface of the rotor R. The inlet screw 52 and the outlet screw 51 respectively pump fluid towards or away from the impeller. The outer circumference of the screw has a gap to prevent it from contacting the housing during normal operation. These screws are designed to contact the housing to act as mechanical backup bearings when subjected to radial transient shock loads that cannot be supported by magnetic bearings. The relatively wide axial land on the periphery of the screw provides a low contact pressure bearing surface for such purpose. Screw 51
, 52 is actively pumped through the bearing gaps 20, 20 'by 50, and does not have to depend on the pump flow rate to generate the differential pressure for flushing. This is also applicable to the centrifugal pump design shown in FIG.

The scoop 9 of FIG. 1 in the outlet bearing gap 20 ′ is omitted in FIG. This is because the screw pump 51 can generate a sufficient flow without a scoop. The ability to omit the scoop, if desired, provides design flexibility in preventing thrombus formation at this location. Screw 51
, 52 may be radially offset from the pump axis A _p to avoid the formation of blood stagnation regions therein.

FIG. 10 shows an improved axial flow Archimedean threaded impeller 80 for a blood pump that uses multiple short blades. The basic or conventional Archimedes screw pump, invented by Archimedes of ancient Greece for irrigation of low-pressure grain crops, consists of a single screw with several turns. To minimize back leakage in these pumps, such as in implantable blood pumps, where significant pressure must be generated, at least one completely 360 degree is used to minimize back leakage in these pumps. You need a screw that turns. Particularly in order to form the screw for one rotation with a shallow helix angle, the axial length of the screw becomes considerably large. Long impellers cannot fit in a compact axial flow pump as taught in the present invention. Therefore, the length L of the impeller must be reduced.

Testing of the prototype by the inventor has found that it is possible to produce Archimedes screws with comparable hydraulic efficiency whose overall length can be considerably shortened by using a plurality of short blades arranged in the circumferential direction. It was With more blades, the length can be reduced accordingly. These blades have a helix angle similar to a single long screw and perform equivalently. Because of the use of multiple blades that partially overlap each other, the back leak in the impeller is equivalent to that of a single longer 360 degree turn of the screw. This has been confirmed experimentally.

Due to the finite lateral width of the blades, using more blades than necessary reduces the open flow area between the blades. This reduces the flow rate and increases the shear stress of the blood at a given flow rate to undesired levels. This causes hemolysis. For this reason, it is ideal to limit the axial length of the blades by using the minimum number of blades needed to achieve the desired result.
An impeller using two blades has its axial length halved, and an impeller having three blades has an axial length of one-third, and if four blades are used, its length is one-fourth. However, when the number of blades exceeds four, the profit obtained becomes smaller.

The blade overlap D in FIG. 10 may range from zero to ideally several degrees in the circumferential direction. This eliminates a straight fluid leakage path that may exist between the leading edge of one blade and the trailing edge of an adjacent blade. The presence of such a back leak path will significantly reduce the efficiency of the impeller. That is, when the gap D is negative, an excessive leak occurs at high pressure and the efficiency of the pump decreases. In the preferred prototype, a three-blade impeller with an outer diameter of 0.80 inches is used, each blade having a length of 0.90 inches and an overlap D of 10 degrees. FIG. 11 shows these blades in end view as A4, B4 and C4. The three blades shown each extend over a range of 120 degrees + 10 degrees overlap = 130 degrees.

The Archimedes screw pump pumps fluid by the pushing action of an advancing screw within the surrounding housing, but produces no lift. Aircraft and ship propellers do not function this way, relying on the lift forces generated by the blades and without the housing surrounding them. Overlapping blades are not taught in the design of these propellers. This is particularly noticeable in a two-blade propeller in which the propeller blades are positioned 180 degrees apart and a large gap is provided between the blades. Lifting propellers are not a design problem for these applications because, unlike blood pumps, they do not generate high pressure to propel ships and aircraft.

12 and 13 show a novel outlet stator or diffuser designed to redirect the generally tangential flow from the impeller to the axial flow to be discharged at the pump outlet. There is. The vanes are curved almost 90 degrees. The fluid is introduced almost tangentially. The cross-section of the flow path between the vanes at the inlet to the diffuser is kept about the same as at the outlet of the impeller. this is,
Match the blood flow rate to minimize damage to blood cells at the diffuser inlet. The rotational kinetic energy of blood flowing out of the impeller is converted into static pressure as the flow travels through the inter-vane passages where the flow passage area gradually increases. At the diffuser outlet where the fluid is all axial, the flow velocity has a minimum value.

It is important to convert tangential flow to axial flow without fluid separation at the back of the vane. Otherwise, turbulent blood damage and hemostasis may occur. This is accomplished by interposing a shorter auxiliary vane 55 between the main curved vanes 56. These auxiliary vanes 55 are not located at the vane entrance, and are set back by a distance Y. Vane 5 in front of diffuser entrance
The absence of 5 provides the additional circumferential space needed to ensure the desired flow path cross-sectional area for flow velocity matching. This is important in such small diameter small pumps, where the main vanes, which have a finite thickness T, occupy considerable space. It is intended that the space between the vanes, which has been narrowed by the addition of the secondary vanes 55, is provided at the portion where the curvature of the main vanes where the separation of the fluid is likely to start becomes severe. The narrower gap W acts to force the flow against the convex wall downstream of the vanes, minimizing fluid separation.

It is also desirable to have a small number of vanes so as to minimize the surface area in contact with the blood to minimize thrombus formation. By appropriately selecting the vane thickness T, the desired average passage width W can be achieved for a given number of vanes. Eight relatively thin vanes are used in the preferred embodiment.

As shown in FIGS. 12 and 13, the curvature radius “r” is as large as possible in the vane.
Can also reduce fluid separation.
Using a large radius of curvature leads to an increase in the axial length of the diffuser. In a small pump, it is not preferable that it is long in the axial direction. If the shear stress of the blood can be sufficiently low, then a small passage width W can be used and a short axial vane with a small radius of curvature "r" can be employed without causing fluid separation. In the preferred embodiment, if W is in the range of 0.10 to 0.40 inches and the curvature "r" is in the practical range of 0.30 to 1.25 inches, fluid separation is minimal. Be converted.

In FIG. 7, the impeller 4 is shown as having a greater axial length than that shown in FIG. The advantage of this structure is that the box 2 indicated by the dashed line
The point is that the armature magnet of the motor located in 1 can be located completely below the impeller blades. Since the impeller and the magnet rotate together, eddy current is not generated in the blade due to the rotation of the rotor. This is not the case in Figure 1 where very short impellers are used. In this construction, the fixed exit stator blades are placed in the magnetic field above the rotating magnets, so that eddy currents are generated in these blades and energy is consumed. This leads to increased power consumption of the motor.

FIG. 14 shows another embodiment of the peripheral ring magnet bearing. This example
It has a further advantage. An axial sectional view of the bearing B is shown in FIG. A radially magnetized magnet 13 "is used and is arranged between the circular peripheral rings 11, 11 '. The magnet ring consists of a plurality of radially magnetized pie-shaped sections. Well, or it may be radially magnetized as a single magnet, the axial instability of the bearing is due only to the magnetic flux in the air gap of the peripheral ring. It does not generate any instability from being contained, thus one advantage of this structure is that it reduces axial instability as compared to using the axially magnetized magnet shown in FIG. In addition, the inner diameter of the bearing consists of an almost empty space, and it can be configured as a small pump that can be used for children.

Although the present invention has been described with reference to particular embodiments, various modifications and changes are described without departing from the scope and concept of the invention as described herein and as defined in the appended claims. Please understand that is possible.

[Brief description of drawings]

1 is a longitudinal cross-sectional view along the center of the housing of an axial turbopump using two magnetic bearings to suspend the rotor, the interior of the pump being partially broken to show the internal fluid passages. And the motor stator coils and rotor magnets are shown in dotted lines.

[Fig. 2]   2 is a front elevational view at the inlet end of the pump shown in FIG. 1, looking in direction A; FIG.

[Figure 3]   2 is a rear elevational view of the pump shown in FIG. 1 looking in direction B at the outlet end. FIG.

4 is a front elevational view at the inlet end of a centrifugal turbopump according to the present invention using the same magnetic bearing as the axial flow pump of FIG. 1, viewed from the direction A of FIG.

5 is a side elevational view of the pump shown in FIG. 4 showing the impeller housing and fluid outlet, further showing the fixed stator coil of the rotary motor in dotted lines.

6 is a longitudinal cross-sectional view along the center of the housing of the pump shown in FIG. 5, with the interior of the pump shown partially broken away to show the fluid passages within the pump. The impeller and its shaft are held by two magnetic bearings without contact.

FIG. 7 is a view similar to FIG. 1, but with two non-contacts attached to the rotor to actively flow blood into the fluid gap formed between the magnetic levitation support and the axial end surface of the rotor. A type Archimedean screw (not cut for clarity) is shown.

FIG. 8 is a block diagram of a circuit for calculating a transient or steady differential pressure applied to a rotor using a magnetic bearing as a force sensor.

FIG. 9 is a block diagram of an electronic feedback system used to periodically vary the pump rpm to control the pump differential pressure.

FIG. 10 is a side elevational view of an axial flow 3-blade impeller emerging from the central hub.
Only two blades are shown.

11 is an end view of the impeller of FIG. 10 as viewed from the direction C of FIG. 10, in which all three blades having a circular outer diameter are shown.

12 is an end elevation view of the outlet end of a vane structure forming an outlet diffuser in an axial flow structure similar to that shown in FIG.

FIG. 13 is a side elevational view of the vane shown in FIG. 12, with the cylindrical housing shown as a cross-sectional view.

FIG. 14 is a partial cross-sectional view of a very compact peripheral ring magnetic bearing structure using magnets that are magnetized radially rather than axially as an alternative embodiment.

FIG. 15 is an enlarged cross-sectional view showing a part of the boundary between the rotor and the magnetic bearing in a partially broken manner,
FIG. 6 shows the fluid gap formed between the rotor and the magnetic bearings and how the fluid and air gaps at the upstream or inlet ends relative to the downstream or outlet ends to compensate the average fluid force acting on the rotor. It is changed.

16 is an elevational view of the axial end surface of the rotor shown in FIG. 1, positively or actively imparting a centrifugal force to the blood in the fluid gap to circulate the blood in the fluid gap and create a stagnation. FIG. 4 shows a plurality of radial vanes formed on the rotor surface to prevent it.

FIG. 17 is a side elevational view of a portion of the rotor shown in FIG. 16, showing radial vanes.

FIG. 18 shows a generally linear relationship between axial displacement of the rotor and axial instability forces from positions that are approximately equidistant between the magnetic bearings.

─────────────────────────────────────────────────── ─── Continued front page    (81) Designated countries EP (AT, BE, CH, CY, DE, DK, ES, FI, FR, GB, GR, IE, I T, LU, MC, NL, PT, SE), OA (BF, BJ , CF, CG, CI, CM, GA, GN, GW, ML, MR, NE, SN, TD, TG), AP (GH, GM, K E, LS, MW, MZ, SD, SL, SZ, TZ, UG , ZW), EA (AM, AZ, BY, KG, KZ, MD, RU, TJ, TM), AE, AG, AL, AM, AT, AU, AZ, BA, BB, BG, BR, BY, CA, C H, CN, CR, CU, CZ, DE, DK, DM, DZ , EE, ES, FI, GB, GD, GE, GH, GM, HR, HU, ID, IL, IN, IS, JP, KE, K G, KP, KR, KZ, LC, LK, LR, LS, LT , LU, LV, MA, MD, MG, MK, MN, MW, MX, MZ, NO, NZ, PL, PT, RO, RU, S D, SE, SG, SI, SK, SL, TJ, TM, TR , TT, TZ, UA, UG, US, UZ, VN, YU, ZA, ZW

Claims (44)

[Claims] 1. A blood pump for propelling blood therethrough, comprising: a pump housing defining a pump axis and having an inlet opening and an outlet opening at opposite axial ends; a rotor axis and an opposing rotor axis end; A rotor defining the rotor and radial stability of the rotor provided at both ends of the rotor axis in the pump housing to maintain the rotor axis substantially in line with the pump axis during operation of the pump. Magnetic levitation support means for substantially maintaining and magnetically levitationally supporting the rotor so as not to physically contact the housing, and defining a fluid gap between the rotor axial ends, and the rotor externally. Absorbs the load applied to and maintains an axial separation between the fluid gap and the rotor and the pump housing Control means for maintaining the stability of the rotor in the axial direction, and with the rotation of the rotor, the rotor is provided to function so as to suck blood from the inlet opening and expel it from the outlet opening. Impeller means, drive means for pumping blood by rotating the rotor and the impeller means, while the rotor is rotating, continuously flowing blood into the fluid gap, thrombus in the fluid gap A blood pump for preventing the formation of blood, and a blood pump. 2. The rotor has a cylindrical rotor housing constituted by a cylindrical wall coaxial with the rotor axis, and a non-magnetic circular lateral wall provided at each rotor axis end. The blood pump according to claim 1, wherein: 3. The cylindrical wall and the lateral wall are sealably coupled together to seal within the rotor housing.
Blood pump according to. 4. The magnetic levitation support means comprises magnetic field forming means for forming an axially oriented magnetic field across the fluid gap at each axial end of the rotor. The described blood pump. 5. The rotor is spaced apart from the rotor by a predetermined axial distance so that the magnetic levitation support means forms the fluid gap along the pump axis at a central portion in the pump housing. A support member provided at each axial end of the rotor, wherein the fluid gap extends at each axial end of the rotor between the pump axis and the radially outermost end of the rotor. The blood pump according to claim 4, wherein the blood pump is a blood pump. 6. The support members each have an inlet and outlet fairing configured to define a streamlined profile to reduce turbulence and fluid separation at the inlet and outlet openings, respectively. The blood pump according to claim 5, which is characterized in that. 7. The axial direction in which each support member contains an active bearing coil for generating at least one component of the axial magnetic field and which is provided in each axial end of the rotor and traverses the gap. Blood pump according to claim 5, characterized in that it comprises a permanent magnet for generating the second component of the magnetic field. 8. Associated with each active bearing coil is a magnetizable yoke through which the axial magnetic field flows, the axial magnetic field aligned across the fluid gap to the outer circumference of each axial end of the rotor. The blood pump according to claim 7, wherein the respective yokes are arranged so as to perform. 9. A pole piece cooperating with a corresponding permanent magnet is provided at each axial end of the rotor to direct the second component of the axial magnetic field near the outer circumference of each axial end of the rotor. The blood pump according to claim 8, wherein the blood pump is provided. 10. The pole piece and the yoke include magnetizable portions axially spaced across an associated fluid gap, the magnetizable portions having a cross-sectional area in a radial plane of the permanent magnet. Magnetic flux density across the fluid gap between the associated radially oppositely disposed peripheral rings, which are formed as a plurality of radially concentric, generally concentric peripheral rings that are smaller than the cross-sectional area of The blood pump according to claim 9, wherein the blood pump is increased. 11. The fluid gap is substantially the same size at both axial ends of the rotor so that the blood flow rate and the flushing flow rate at both axial ends of the rotor are approximately the same. The blood pump according to claim 1. 12. The permanent magnet provided at each axial end of the rotor defines a magnetic axis extending substantially in line with the pump housing axis and the rotor axis. Blood pump. 13. The blood pump according to claim 7, wherein each permanent magnet provided at each axial end of the rotor is magnetized in a radial direction. 14. Each of the support members includes an active bearing coil for generating one component of the axial magnetic field and a permanent magnet for generating a second component of the axial magnetic field. 6. A blood pump according to claim 5, wherein magnetizable yokes are disposed at each axial end of the rotor to provide a return path into the rotor for the axial magnetic field. 15. The flushing means for flushing the fluid gap at the inlet opening is provided generally along the pump axis and opens at one axial end in the direction of the inlet opening and at the other axial end. An inlet axial bore communicating with the associated fluid gap such that at least some of the blood entering the inlet opening is directed to flow through the inlet axial bore and the associated fluid gap. The blood pump according to claim 1, wherein 16. An outlet axis wherein the flushing means for flushing the fluid gap at the outlet opening has one axis end communicating with the fluid gap at the outlet end and opening at the other axis end in the direction of the outlet opening. A directional hole and at least some of the blood that is made to flow through the pump housing by the impeller means is diverted toward the associated fluid gap, thereby allowing blood to flow through the fluid gap at the outlet opening. 2. Blood pump according to claim 1, characterized in that it has diverting means for flowing through and discharging it through the other axial end to the outlet opening. 17. The blood flushing means has at least one protrusion provided on the rotor that protrudes into the fluid gap, the protrusion enhancing the centrifugal flow of blood associated with rotation of the rotor. The blood pump according to claim 1, wherein the blood pump is adapted to be operated. 18. The blood pump of claim 17, wherein the at least one protrusion is in the form of at least one generally radial fin. 19. A plurality of radial fins are provided on the rotor that project into the fluid gap at least at the inlet opening, the radial fins being angularly spaced from each other about the rotor axis. The blood pump according to claim 18, wherein: 20. The rotor has an annular flow relative to blood flowing from the inlet opening to the outlet opening in cooperation with a cylindrical wall defining an axis extending coincident with the rotor axis and the pump housing. A generally cylindrical rotor housing having an outer cylindrical surface defining a passage, the impeller means comprising a plurality of axially oriented spirals provided over at least an axial length of the outer cylindrical surface. Blades, which are generally equally angularly spaced from each other about the rotor axis and have a pitch and length such that they at least partially overlap circumferentially when viewed along the rotor axis. The back leak is reduced thereby, and the back leak is reduced.
Blood pump according to. 21. Three helical blades are provided, each helical blade extending a total of 130 ° about the rotor axis and overlapping 10 ° with adjacent helical blades. The blood pump according to claim 20. 22. An active bearing coil forming a part of said magnetic levitation support means, wherein said control means is arranged to detect a variation in axial position of said rotor; and a differential acting on said rotor based on said variation. A blood pump according to claim 1 including means for establishing pressure. 23. The control means includes a "substantially zero power" (VZP) control feedback loop for axially stabilizing the rotor based on sensed axial velocity of the rotor and coil current. The blood pump according to claim 22, wherein: 24. The blood pump according to claim 1, wherein the pump housing and impeller means are arranged to pump blood axially between the inlet opening and the outlet opening. 25. The blood pump according to claim 1, wherein the pump housing and impeller means are arranged to centrifugally expel blood in a radial direction. 26. The method according to claim 1, wherein the control means includes means for applying a differential pressure fluctuation of a desired cycle to the rotor so as to generate a pulsating flow of blood in the pump. Blood pump. 27. The control means comprises means for setting the frequency of the pulsating flow and a feedback loop for comparing the selected set frequency and differential pressure with a variation value detected as a function of the axial position of the rotor. 27. The blood pump of claim 26, comprising: 28. The pump housing and impeller means are arranged to produce an axial flow of blood, and the tangential flow of blood produced by rotation of the impeller means is directed to the axial flow at the outlet end. The blood pump according to claim 1, further comprising a flow conversion unit that converts the flow to prevent turbulence and separation of fluid at the outlet end. 29. The flow transforming means includes a plurality of curved vanes located at the outlet end, adjacent vanes having different axial lengths, thereby leading to the flow transforming means. 29. The blood pump of claim 28, wherein the additional circumferential space required to obtain the desired flow path cross-sectional area for flow velocity matching at the inlet is provided. 30. The drive means is disposed on the pump housing in a substantially axially centered position within the rotor to define a center of gravity of the rotor, and a stator provided on the pump housing so as to be substantially axially aligned with the center of gravity. A blood pump according to claim 1, further comprising a coil, whereby equal radial loads are applied at both magnetically levitated axial ends. 31. The blood pump of claim 1, wherein the blood flushing means includes means for actively moving blood through the fluid gap during rotation of the rotor. 32. The blood flushing means diverges at least one Archimedes screw for diverting blood at the pump inlet opening and forcing the diverted blood to flow through a fluid gap at the inlet opening. 32. The blood pump of claim 31, comprising: 33. The blood flushing means includes at least one Archimedes screw for actively withdrawing blood from the fluid gap at the pump outlet opening and directing the withdrawn blood to the pump outlet opening. The blood pump according to claim 31. 34. The blood pump according to claim 15, further comprising means for preventing stagnation of blood in the inlet axial hole. 35. The stagnation preventing means has an element projecting at least partially into the inlet axial bore from the rotor and is adapted to move blood therein as the rotor rotates. The blood pump according to claim 34, wherein: 36. The blood pump according to claim 35, wherein the element is provided eccentrically with respect to the rotor axis. 37. The blood pump according to claim 16, further comprising means for preventing stagnation of blood in the outlet axial hole. 38. The stagnation prevention means has an element that at least partially projects from the rotor into the outlet axial hole, and is adapted to move blood therein as the rotor rotates. The blood pump according to claim 37, wherein the blood pump is a blood pump. 39. A blood pump according to claim 38, wherein the element is provided eccentrically with respect to the rotor axis. 40. A blood pump for propelling blood therethrough, a pump housing defining a pump axis and having an inlet opening and an outlet opening; a rotor defining a rotor axis and opposing rotor axis ends; Provided at both ends of the rotor axis within the pump housing and generally maintaining radial stability of the rotor to maintain the rotor axis substantially in line with the pump axis during operation of the pump, Support means for rotatably supporting the rotor in the housing and defining a fluid gap between the rotor axis ends; and the rotor for absorbing an externally applied load, the fluid gap and the rotor. The axial stability of the rotor so as to maintain an axial separation between the pump housing and the pump housing. Control means for holding, impeller means provided in the rotor for functioning to suck blood from the inlet opening and discharge it from the outlet opening with rotation of the rotor, and rotate the rotor and the impeller means Drive means for pumping blood by causing the blood to flow continuously and actively into the fluid gap during rotation of the rotor to prevent thrombus formation in the fluid gap. And a blood pump. 41. A blood pump for propelling blood therethrough, a pump housing defining a pump axis and having an inlet opening and an outlet opening; a rotor defining a rotor axis and opposing rotor axis ends; Provided at both ends of the rotor axis within the pump housing and generally maintaining radial stability of the rotor to maintain the rotor axis substantially in line with the pump axis during operation of the pump, Support means for rotatably supporting the rotor within the housing and defining a fluid gap between the rotor axis ends for providing stability and radial stiffness of the rotor during operation of the pump. ,
Permanent magnet means for establishing a strong axial magnetic field across the fluid gap, and axial separation between the fluid gap and the rotor and the pump housing for the rotor to absorb externally applied loads. Control means for maintaining the stability of the rotor in the axial direction so that the rotor can be maintained, and with the rotation of the rotor, blood is sucked from the inlet opening and is discharged to the rotor from the outlet opening. Impeller means provided, drive means for pumping blood by rotating the rotor and the impeller means, blood is continuously flowed into the fluid gap during rotation of the rotor, and in the fluid gap. Blood pump for preventing the formation of blood clots. 42. A blood pump for propelling blood therethrough, comprising: a pump housing defining a pump axis and having an inlet opening and an outlet opening; a rotor defining a rotor axis and opposing rotor axis ends; Provided at both ends of the rotor axis within the pump housing and generally maintaining radial stability of the rotor to maintain the rotor axis substantially in line with the pump axis during operation of the pump, Support means for rotatably supporting the rotor in the housing and defining a fluid gap between the rotor axis ends; and the rotor for absorbing an externally applied load, the fluid gap and the rotor. The axial stability of the rotor so as to maintain an axial separation between the pump housing and the pump housing. Control means for holding, impeller means provided in the rotor for functioning to suck blood from the inlet opening and discharge it from the outlet opening with rotation of the rotor, and rotate the rotor and the impeller means A pumping means for pumping blood, and a flushing means for continuously flowing blood into the fluid gap during rotation of the rotor to prevent formation of thrombus in the fluid gap. The rotor means cooperates with the cylindrical wall that defines an axis that extends in line with the rotor axis and the pump housing to form an annular flow path for blood flowing from the inlet opening to the outlet opening. A generally cylindrical rotor housing having an outer cylindrical surface forming the impeller means, the impeller means having at least an axial length of the outer cylindrical surface. A plurality of axially-directed spiral blades, the blades being substantially equally angularly spaced from each other about the rotor axis and at least partially when viewed along the rotor axis. A blood pump having a pitch and a length that substantially overlap each other in the circumferential direction, thereby reducing back leak. 43. A blood pump for propelling blood therethrough, a pump housing defining a pump axis and having an inlet opening and an outlet opening; a rotor defining a rotor axis and opposing rotor axis ends; Provided at both ends of the rotor axis within the pump housing and generally maintaining radial stability of the rotor to maintain the rotor axis substantially in line with the pump axis during operation of the pump, Support means for magnetically rotatably supporting the rotor within the housing and defining a fluid gap between the rotor axial ends; and the rotor for absorbing an externally applied load, the fluid gap And an axial direction of the rotor so as to maintain an axial separation between the rotor and the pump housing. Control means for maintaining qualitative characteristics; impeller means provided on the rotor for functioning to suck blood from the inlet opening and expel it from the outlet opening with rotation of the rotor, the rotor and the impeller means Drive means for pumping blood by rotating the, and blood wash means for continuously flowing blood into the fluid gap during rotation of the rotor to prevent the formation of thrombus in the fluid gap. The pump housing and the impeller means are arranged to generate an axial flow of blood, and further converting the tangential flow of blood produced by rotation of the impeller means into an axial flow at the outlet end. And has a flow conversion means for preventing fluid separation and turbulence at the outlet end. Blood pump to be. 44. A blood pump for propelling blood therethrough, a pump housing defining a pump axis and having an inlet opening and an outlet opening; a rotor defining a rotor axis and opposing rotor axis ends; Provided at both ends of the rotor axis within the pump housing and generally maintaining radial stability of the rotor so as to maintain the rotor axis substantially in line with the pump axis during operation of the pump, Support means for rotatably supporting the rotor in the housing and defining a fluid gap between the rotor axis ends, the rotor absorbing the load applied externally, the fluid gap and the rotor The axial stability of the rotor so as to maintain an axial separation between the pump housing and the pump housing. Control means for holding, impeller means provided in the rotor for functioning to suck blood from the inlet opening and expel it from the outlet opening as the rotor rotates, and rotate the rotor and the impeller means The driving means for pumping blood by causing the blood to flow through the fluid gap during rotation of the rotor to prevent thrombus formation in the fluid gap. In order to detect the axial velocity of the rotor, the control means physiologically controls the pump according to the heart rate of the patient, which is generally proportional to the average pressure difference between the systolic blood pressure and the diastolic blood pressure. A blood pump comprising means arranged and means for establishing a differential pressure acting on the rotor based on the speed fluctuations.