Note : Les descriptions sont présentées dans la langue officielle dans laquelle elles ont été soumises.
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MAGNETIC SUSPENSION BLOOD PUMP
BACKGROUND OF THE INVENTION
Field of the Invention
The present invention generally relates to rotary blood pumps that can be
implanted into the chest of humans and can be used to assist a human heart in
pumping
blood, and, more specifically, to such blood pumps that use magnetic
suspensions.
Description of the Prior Art
The implantable blood pumps according to the latest technology that are now
being
developed to assist the heart are turbo pumps. They come in axial flow
configurations,
such as the Jarvik 2000; centrifugal configurations, such as one being
developed by the
Cleveland Clinic; and mixed flow types such as the "Streamliner" being
developed at the
University of Pittsburgh. All employ a high-speed rotary impeller rotating at
thousands of
rpm. Most, including the Jarvik 2000, use hard-contact journal bearings to
support the
rotor. Such use is not desirable because blood damage and thrombosis can be
caused by
the bearings. To try to circumvent contact-bearing problems, magnetic bearings
are now
being employed, as in the "Streamliner" pump. These are non-contacting
bearings and
result in minimal blood damage, since the bearing clearances can be kept large
to reduce
shear stress in the blood. However, the problem still exists of having to
thoroughly wash
out all the bearing clearances with fresh blood. This washout is essential to
eliminate the
formation of thrombus.
Magnetic bearings must be packaged in a small space in order to minimize the
size
of the pump, and this can be quite difficult. Most magnetic bearing pumps are
too large
and therefore unacceptable.
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Another requirement for implantable blood pumps is low power consumption.
Pumps that employ magnetic bearings are notorious for their power consumption,
which
can be as high as 20 watts for just the bearings, 5 watts for the bearings)
being more
typical. The power delivered to the blood in a left ventricular assist device
(LVAD) is about
3.0 watts, so one does not want to expend more than 1.0 additional watt for
the magnetic
bearings.
Most magnetic bearings use permanent magnets or electromagnets to generate
radial magnetic fields that directly suspend the rotor radially. However, the
bearing radial
"stiffness" obtained using the relatively low air gap fields produced by the
magnets is not
high. A large bearing is, therefore, needed to hold imposed loads with small
radial
deflection.
Radially passive magnetic bearings are inherently unstable axially, as stated
by
Ernshaw's Law. Active axial control is, therefore, required to stabilize a
rotor suspended
by such bearings. Particularly for an axial flow turbo pump that has
substantial axial forces
acting on the rotor, the power consumed by the active coils can be
unacceptably large. A
"virtually zero power" (VZP) control loop is sometimes used to reduce power
consumption.
This control is generally known as VZP control and was first used back in the
1970s by J.
Lyman, one of the founders of magnetic suspensions.
Implantable turbo blood pumps are typically run at constant rpm because it has
been difficult to close the loop around the patient and physiologically vary
pump flow rate
according to the needs of the patient. By providing a base rate of flow,
increased blood
demand due to activity level is made up by the natural heart. However, a sick
heart cannot
make up much demand, and activity level is limited. Whatever cardiac output
demand is
made up by the patient's heart undesirably loads the sick left ventricle. To
physiologically
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control pump output flow, extraneous sensors have sometimes been added to
measure
physiologic parameters of the patient. These have included the addition of
blood pressure
transducers to measure the pump outlet pressure or differential pressure. This
is highly
undesirable because the addition of extraneous sensors can cause thrombosis
and long-
term hemodynamic reliability concerns. A known LVAD uses an invasively placed
series
ultrasonic flowmeter to determine pump flow rate since the LVAD cannot
directly measure
its output blood pressure.
The natural heart produces pulsatile flow. Experiments have shown that this
unsteady flow minimizes the onset of thrombosis in the larger arteries of the
body
because the flow pattern constantly changes. In a pulsatile flow pump, areas
of stagnant
flow are minimized or eliminated not only in the patient's arteries at the
pump outlet, but
within the pump itself. Current turbo pumps are direct current (DC) or steady
flow devices
that do not produce pulsatile flow. Even as the heart of a sick patient
recovers and
contributes some degree of pulsatile flow to the body, the degree of
pulsatility is much
less than that of the natural heart since the LVAD blood pump is unloading the
sick heart.
For "Bridge To Recovery" long-term implants, pulsatile flow from the LVAD is
highly
desirable.
SUMMARY OF THE INVENTION
Accordingly, it is an object of this invention to provide alternate means to
wash out
the magnetic bearing gaps with fresh blood to eliminate thrombus formation at
the
bearings.
It is another object of the invention to allow bearing washout under minimal
flow
conditions through the pump.
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It is still another object of the present invention to provide non-contact
active
washout means for the bearings.
It is yet another object of the present invention to provide a magnetic
bearing
geometry that is easily washed out by the blood flow to prevent areas of
stasis.
A further object of this invention is to provide a small size bearing system
that is
simple in construction and packageable with the various turbo pump types for
use with
both adults and children.
A still further object of the present invention is to provide a control system
that
requires very low power when used with the disclosed high load capacity
bearings.
It is yet a further object of the present invention to determine pump
differential
pressure in a direct manner without the addition of extraneous sensors.
It is an additional object of this invention to provide an active coil and
magnet
geometry that requires low power approaching zero to sustain axial loads.
It is still an additional object of the invention to provide safety of
pulsatile flow by
eliminating the undesirable condition of reverse flow through the pump.
It is yet an additional object of this invention to provide pulsatile flow in
a reliable
manner using pump differential pressure determined directly by the magnetic
bearings.
It is also an object of this invention to shorten the length of an Archimedes
screw
type axial flow impeller by providing multiple parallel flow blades that
minimally overlap. In
mini-size blood pumps, for which this invention is intended, minimizing axial
length of the
pump is desirable particularly for applications in small women and children.
It is furthermore an object of this invention to provide a compact outlet
stator of
short axial length that does not damage blood while recovering impeller
pressure.
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In order to achieve the above objects, as well as others that will become
evident
hereafter, a blood pump in accordance with the present invention comprises a
pump
housing defining a pump axis, and inlet and outlet openings at opposite axial
ends of said
pump housing. A rotor is provided that defines a rotor axis and opposing rotor
axial ends.
Magnetic suspension means is provided within said pump housing at said rotor
axial ends
for magnetically suspending said rotor and passively maintaining the radial
stability of said
rotor so that said rotor axis remains substantially coextensive within said
pump axis
during operation. Control means is provided for maintaining axial stability of
said rotor so
that said rotor may absorb externally imposed axial loads and so that contact
of said rotor
within said pump housing is eliminated. Impeller means on said rotor operates
to draw
blood into said inlet opening and expel the blood through said outlet opening
with rotation
of said rotor. Drive means is provided for rotating said rotor and impeller
means to
thereby pump the blood, fluid gaps being formed between said rotor axial ends
and said
magnetic suspension means. Blood washout means is provided for continuously
moving
blood through said fluid gaps during rotation of said rotor to prevent
formation of thrombus
in said fluid gaps.
Preferably, the washout means provides positive or active flow of blood
through
fluid gaps where stagnation of blood might otherwise take place. In accordance
with
another feature of the invention, the drive means is arranged to drive the
impeller means
at a selected rotational speed, and means are provided for sensing the
pressure
differential within said pump housing and imparting a cyclic variation to said
selected
rotational speed of said drive means to provide pulsating movements of the
blood through
the pump and into the patient's circulatory system. Also, one structure for
washout relies
on generating differential pressures across the bearing gaps in a passive
manner using
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the flow itself. An alternate structure attaches Archimedes screw pumps to the
front and
rear of the rotor to actively pump blood through the bearing gaps. In
approximately 30% of
heart-assist patients the natural heart re-conditions sufficiently after a
year or two of LVAD
use so that the pump is no longer needed. Rather than explant the device, the
pump can
be left in place and operated at minimal flow and power consumption. The
active screw
pumps allow proper bearing gap washout when the pump is put to sleep and
minimally
used.
An important feature of the invention is to mount the impeller means on a
magnetically suspended rotor that is inherently stable in radial directions
and to provide
direct feedback signals useful for stabilizing the rotor in the axial
direction. This feature
substantially simplifies the design and construction of the pump, reduces its
cost of
manufacture and substantially enhances the reliability over extended periods
of use.
A compact high radial stiffness magnetic bearing uses axial fringing ring
magnetic
fields to passively support the pump rotor radially. The flux is focused or
concentrated
from a permanent magnet in the fringing rings to produce very high radial load
capacity in
a small size. This is different than typical radially passive magnetic
suspensions that
employ radial magnetic fields. Active axial control stabilizes the bearing
using a "Virtually
Zero Power" control feedback loop. Low power and small size make the bearing
applicable to axial flow and other configuration blood pumps particularly
suitable for
implantation. Differential pressure across the bearing fluid gaps, forcefully
positively or
actively washes the gaps with fresh blood to eliminate thrombus and flow
stagnation.
The rotor force on the magnetic bearings can be measured by the bearing
control system.
This allows the direct determination of differential pressure across the pump.
This
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parameter can be used to obtain a pulsatile output pressure and flow and to
exert
physiological control on the pump output so as to match the patient's
activity.
In the present invention, a very high radial stiffness bearing is obtained.
This is
accomplished by employing an axially directed fringing ring field that has a
radial load
capacity an order of magnitude higher than radially directed fields. This
allows one to use
a small diameter bearing that was not heretofore feasible. The high-load
capacity results
in low power consumption as well.
Until the present invention, it has not been possible to determine turbo pump
differential pressure in a direct manner. Turbo pump differential pressure can
be used in
part to exert physiological control on the pump flow rate as demanded by the
patient's
activity level and heart rate.
BRIEF DESCRIPTION OF THE DRAWINGS
The invention, in accordance with preferred and exemplary embodiments,
together
with further objects and advantages, is more particularly described in the
following
detailed description taken in conjunction with the accompanying drawings, in
which:
Fig. 1 is a longitudinal section through the center of the housing of an axial
flow
turbo pump embodying two magnetic bearings to suspend the rotor, the interior
of the
pump being partially broken away to show the internal fluid flow passages, the
motor
stator coils and rotor magnets being shown in dash outline.
Fig. 2 is a front elevational view at the inlet end of the pump shown in Fig.
1, as
viewed in the direction A.
Fig. 3 is a rear elevational view at the outlet end of the pump shown in Fig.
1, as
viewed in the direction B.
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Fig. 4 is a front elevational view at the inlet end of a centrifugal type
turbo pump in
accordance with the invention, as viewed in the direction A in Fig. 6,
employing the same
magnetic bearings as in the axial flow pump of Fig. 1.
Fig. 5 is a side elevational view of the pump shown in Fig. 4, showing the
impeller
housing and flow outlet, and further showing the rotary motor stationary
stator coils in
dash outline.
Fig. 6 is a longitudinal section through the center of the housing of the pump
shown
in Fig. 5, the interior of the pump also being partially broken away to show
the internal
fluid flow passages in the pump, the impeller and its shaft being supported by
two
magnetic bearings without contact.
Fig. 7 is a view similar to Fig. 1, but shows two non-contacting Archimedes
screws
(unsectioned for clarity) attached to the rotor for actively moving the blood
through the
fluid flow gaps formed between the rotor axial end surfaces and the magnetic
suspensions.
Fig. 8 is a block diagram of a circuit for calculating the transient or steady-
state
differential pressure across the rotor using the magnetic bearings as force
sensors.
Fig. 9 is a block diagram of an electronic feedback system used to cyclically
vary
pump rpm in order to cyclically control pump differential pressure.
Fig. 10 is a side elevational view of an axial flow three-blade impeller, with
only two
blades in view, emanating from a central hub.
Fig. 11 is an end view of the impeller shown in Fig. 10, showing all three
blades,
which are round at their outer diameters, as viewed along direction C in Fig.
10.
Fig. 12 is an end elevational view of an outlet end of the vane geometry that
forms
an outlet diffuser in an axial flow configuration, similar to the one shown in
Fig. 1.
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Fig. 13 is a side elevational view of the vanes shown in Fig. 12 with the
cylindrical
housing in cross section.
Fig. 14 is a partial cross sectional view of an alternate, very compact
fringing ring
magnetic bearing construction that employs a radially magnetized magnet rather
than an
axial one.
Fig. 15 is an enlarged section, partially broken away, of the magnetic bearing
and
rotor interface, showing the fluid flow gaps formed between the rotor and the
magnetic
bearing, and illustrating how the fluid flow gaps and air gaps may be modified
at the
upstream or inlet end with reference to the downstream or outlet end to
compensate for
average fluid forces acting on the rotor.
Figure 16 is a side elevational view of an axial end surface of the rotor
shown in
Fig. 1, illustrating a plurality of radial vanes formed on the surface to
provide positive or
active centrifugal forces on the blood within the fluid gaps to force the
blood to circulate
within said gaps and prevent stagnation therein.
Fig. 17 is a side elevational view of that portion of the rotor shown in Fig.
16 and
illustrating the radial vanes.
Fig. 18 illustrates the general linear relationship between axial instability
force and
axial displacement of the rotor from its position substantially equidistant
between the
magnetic bearings.
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DETAILED DESCRIPTION OF THE INVENTION
In the following descriptions of axial flow and centrifugal flow pumps in
accordance
with the invention, the magnetic bearing configuration is the same, as well as
other
common elements such as the fluid flow passages, magnetic circuits and the
like. These
common elements are assigned the same or primed identifying reference numerals
throughout for consistency.
Fig. 1 illustrates an axial flow pump P that defines a pump axis Ap and has a
rotor
R, having a rotor axis A~, and a plurality of helically curved impeller blades
4 mounted for
rotation about the pump axis AP, as to be more fully described. The impeller
blades 4 are
attached to a rotor housing 17, which is made of thin-walled titanium for
blood
compatibility. All parts wetted by the fluid are typically titanium or any
other suitable non-
magnetic material. Arranged at each axial end of the rotor R is a magnetic
bearing pole
piece 12, 12', which is iron or other high-saturation magnetic material. An
annular
permanent magnet 13, 13' is axially magnetized and co-axially positioned
between an
associated pole piece 12, 12' and an axial end wall 30, 30' of the rotor R.
The rotor R is
formed of a cylindrical thin-walled titanium shell that hermetically seals the
pole pieces
12, 12' in the rotor housing 17. The rotor housing 17 makes no physical
contact with the
stationary parts of the pump.
Magnetic bearings B, B' are provided at each axial end of the rotor R, a first
half of
which includes the magnets 13, 13' as well as the pole pieces 12, 12'. The
second half
of the magnetic circuit for the magnetic bearings consists of iron yokes 28,
28', which are
stationary. The paths of the magnetic flux 50, 50' are depicted in dash
outline and are
closed loops in the plane shown. The flux passes axially through fluid flow
gaps 20, 20'.
The yokes are sealed in the outlet fairing 6 and inlet fairing 29. Each
fairing includes an
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annular region 6a, 29a for housing the yokes 28, 28' and active bearing coils
14, 14'
proximate to the rotor R, each annular region being sealed by walls 6b, 29b to
define, with
associated rotor axial end walls 30, 30', radial fluid flow gaps 20, 20'. The
walls 6b, 29b
are about .010-inch thick and made thin to minimize the dimension Ga of the
non-
magnetic air gap. As best shown in Fig. 5, fluid flow gaps 20, 20' have a
dimension G, ,
this ranging from about 4 mils to 30 mils to keep the blood shear subhemolytic
depending
on maximum rotor rpm. While the fluid flow gaps 20, 20' at both axial lands of
the rotor R
should normally be approximately equal, the non-magnetic air gaps may be
selected to
have different values, as suggested in Fig. 15, in which the upstream air gap
Ga is greater
than the downstream air gap Ga', while both fluid flow gaps G,, G,' are
maintained to be
about equal by changing the thickness of the walls W, - W4. It should also be
clear that
the same result can be achieved by maintaining the wall thickness the same,
and the fluid
flow gaps 20 the same, while physically displacing the pole pieces 12, 12' and
selectively
moving one or more of such pole pieces away from their associated walls to
establish
desired air gaps that are equal or different to compensate for average forces
acting on the
rotor during normal operation of the pump.
The inlet fairing 29 is radially outwardly tapered at 29c to provide
streamlined flow
without flow separation or turbulence. They attach to the pump housing 24 by
means of
inlet guide vanes 33. These vanes are generally straight or may be helically
curved. Four
curved vanes 33A-33D are shown in Fig. 2. Circumferential spaces 31 are formed
between the inlet vanes. The end view of the inlet fairing 29 is visible in
Fig. 2 as is its
pointed tip 32. Similarly, the streamlined outlet fairing 6 is radially
inwardly tapered, at 6c,
and several equally spaced flow straightener vanes 5 hold outlet fairing 6
centered in the
pump housing 17. A radial clearance 8 exists between the vanes 5 and the rotor
housing
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17. Fig. 3 shows four flow straightener vanes 5A - 5D with circum-ferentia)
spaces 31'
between them as well as the pointed end 7 of the outlet fairing 6.
Referring again to Fig. 1, and more particularly to the geometry of the two
magnetic bearings, the yokes 28 and pole pieces 12 terminate at their outer
diameters in
two axially concentric and radially aligned thin fringing rings 10, 10' and
11, 11',
respectively. Either one fringing ring or a plurality of such rings may be
used. The radial
stiffness and load capacity increase with the number of rings for a constant
magnetic flux
density in the air gaps. The axial magnetic field between associated opposing
fringing
rings is much stronger than the field in the permanent magnets 13, 13' due to
the flux
being concentrated across the narrow or radially thin fringing ring air gaps
whose cross
sectional area is less than that of the magnets 13, 13'. This creates a large
passive radial
load capacity or restoring force. However, the axial field creates an unstable
axial
stiffness or force when the rotor is not axially centered. This occurs when
there are
unequal air gaps at each end. A load on the rotor while it is pumping blood
will create an
axial force, which must be countered to keep the rotor from contacting the
fairings 6, 29.
A transient counterforce is obtained by using active bearing coils 14, 14'
which
surround the inside diameter of the yokes. When currents flow through the
coils
superimposed magnetic fields are placed in series with the magnet flux,
thereby modifying
it. A clockwise current in the upstream bearing coil 14, as viewed along
direction A in
Fig. 1, increases the flux, while a counter-clockwise current decreases the
flux. In this
way the upstream bearing B can be made to attract the rotor R to the right, as
viewed in
Fig. 1, with greater force by increasing the air gap flux. The downstream
bearing can, with
a field decreasing current, be made to attract to the left with lesser force.
This gives a net
positive force to the right. Reversing both currents results in a net force to
the left.
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By utilizing a closed loop control system to control coil current based on an
error
signal of the axial position of the rotor, acceptably small axial movements of
the rotor can
be maintained. However, with an external pressure force on the rotor, power
will be
dissipated in the coils in order to counteract such external forces.
This bearing geometry also avoids this problem by allowing steady-state or
slowly
varying external forces to be countered or neutralized by the permanent
magnets 13
instead of the coil by utilizing a different kind of control than position
sensing. If the rotor
is allowed to axially displace just the right amount, the bearing's axial
instability force can
hold the external load. The coil merely has to stabilize the axial position
using a Virtually
Zero Power (VZP) feedback loop. What this control loop does is drive the coil
current to
zero and the axial rotor velocity to zero in independent control loops. At the
stable axial
position under a load, the axial fringing forces are balanced with the load
and the rotor will
not move in either axial direction. In other words, rotor axial velocity will
be zero and no
DC current is required in either coil to hold a DC load.
A rotor axial velocity signal is needed to implement this VZP control scheme.
The
back emf voltage developed in at least one coil 14 can be used because it is
directly
proportional to rotor axial velocity and is stable over the long term with
regard to direction.
Thus, no added sensors are required to determine axial velocity.
To initially float a bearing that is axially touching, the axial force at the
contacting
end will generally be so great that the opposite coil will not have enough
force capability to
pull the rotor free. To solve this problem an initial DC current is
momentarily placed in
each coil to substantially buck out the flux from each magnet. This minimizes
the axial
instability force and allows the coils to be current controlled to center the
rotor. Once the
rotor is centered, the magnet bucking current can be reduced to zero.
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A position control loop instead of a VZP loop can be used to initially float
the
bearing. To do this an axial position signal is needed. This can be obtained
by measuring
coil inductance by superimposing a low level AC current. Inductance is
inversely
proportional to magnetic circuit axial air gap. Thus, the coils 14 can also be
used as a
position sensor. Once the bearing is floated, control can be switched to the
VZP mode.
Periodically, over the life of the pump, this position control mode may be
switched in
momentarily to guarantee long-term stability of the axial position. This is a
back-up and
verifier for VZP control.
The axial flow pump in Fig. 1 uses a constant diameter impeller blade 4
similar to
the Jarvik 2000 axial flow pump. This constant diameter impeller is ideal for
use with the
disclosed magnetic bearing because the rotor can axially displace any amount
in the
constant diameter housing to hold a load as commanded by the VZP loop.
Sufficient
axial clearances 25, 26 and 20, 20' are used. Mixed flow turbo pumps on the
other hand,
have a radially tapered impeller along its length. The impeller fits closely
in the tapered
housing bore. Consequently, axial motion is much more limited to avoid
contacting the
housing, which limits the axial force capability of the bearing in the VZP
mode.
The most critical goal of any blood pump design is to avoid areas of fluid
stagnation for avoiding thrombus formation. Thrombus can initiate at flow
stagnation
points of zero velocity, particularly at surfaces or by regurgitant or reverse
flow causing
zones of near zero velocity. This has plagued magnetic bearing suspensions
because
totally suspending the rotor results in fluid spaces, gaps or passageways that
are difficult
to flush out with fresh blood. This problem is solved in the present invention
by forcing the
blood flow through these zones under pressure. Referring to Fig. 1, blood
enters the
pump along direction A. It enters a straight hole or a tapered hole 1 near
stagnation
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pressure and, therefore, under high pressure. It flows to pointed pin diverter
or deflector
16, where it is diverted radially into the fluid gap 20, which is a gap that
separates the
rotor and the stator. This flow is additionally sucked out at the gap outside
diameter at 15
by the suction created at the inlet of rotary impeller 4, which is slightly
downstream from
the gap. This differential pressure creates a forced flow that washes out the
gap 20.
Gap 20', at the outlet or downstream end of the rotor, is washed out with the
assistance of a small circumferential lip or scoop 9 of radial height H, about
equal in size
to the gap 20', that forces some of the high velocity flow downstream of the
impeller
blades 4 to go radially into the gap 20'. High stagnation pressure exists at
the scoop 9
relative to the exit pressure of conduit 19. This differential pressure
actively washes out
the bearing gap 20'. Pointed exit cone with sharp convergent point 18 blends
the radial
gap flow into the axial conduit 19 without a substantial stagnation point due
to the sharp
edge 18. To further reduce stagnation at tip or point 18, the tip can be
located slightly off
the center of rotation or pump axis. This eliminates stationary velocities.
The tip deflector
16 at the inlet can also be located off center in relation to the pump axis
with similar
advantage.
Areas of reverse flow in the bearing clearances or fluid flow gaps 20, 20'
have
been eliminated in the disclosed design because unidirectional pressure
differentials
exist across the left and right side gaps. This results in forced flow for
unidirectional
washout. Thus, no stagnant areas should exist, which should result in low
probability of
thrombus. The spinning rotor R also acts like an inefficient centrifugal pump
tending to
pump fluid radially outwardly at the gaps 20 and 20'. However, this effect is
much
weaker than the forced pressure flow, so that reverse flow is eliminated at
the rear gap
20'. Radial flow is augmented in the front gap 20 by this centrifugal action.
Referring to
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Figs. 16 and 17, the efficiency in forcing the blood radially outwardly at the
upstream gap
20 can be enhanced by providing suitably shaped radial fins F on the exterior
surface of
the axial end wall 30. These fins F project into the gap 20 a fraction of the
gap width, on
the order of several mils. Such fins F provide an active movement of the blood
radially
outwardly at the inlet gap 20. The specific sizes, shapes, orientations and/or
positions of
the fins for providing the specific desired results will be well known to
those skilled in the
arts.
An important feature of this invention is the magnetic bearing geometry with
bi-
directional active force capability and unique washout flow geometry that
develops
differential pressures. The rotary motor construction is known art using
brushless motor
technology, so it will not be described in detail. Permanent magnets 21 are
located in the
dotted zone 21 in the rotor. The rotor magnet's field interacts with the
rotating magnetic
field produced by the plurality of stator coils 22 surrounding the pump
housing to create
rotation torque when the coils are commutated. Further details are not given.
However, it
is noted that the side loads produced by the rotary motor should be minimized
so as not
to unduly load the magnetic bearings.
The disclosed bearing geometry uses two separate magnetic bearings B, B' that
are spaced apart axially. This axial separation provides moment or cocking
stability to the
suspended rotor R. The rotor's center of gravity 23 as shown is ideally
located at the
approximate midpoint between the bearings. This imposes equal radial loads on
the
bearings from rotor weight or shock loads. Any motor radial forces are also
equally
shared.
Having thus described a preferred embodiment for axial and mixed flow turbo
pumps, it will now be described how the same basic bearing design and washout
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differential pressures can be applied to centrifugal turbo pumps. In Fig. 5 a
typical
centrifugal pump includes a fluid flow exit opening 2T positioned tangentially
to volute
housing 49. The smaller diameter bearing housing 46 encloses the two magnetic
bearings
that support the rotor. The inside of the pump is shown in Fig. 5. Here only
the bottom of
the centrifugal impeller blade 39 is shown in section. The top blade is shown
by taking a
section between the blades. The spaced-apart magnetic bearings are of the same
basic
geometry previously disclosed except for one change, which is an alternate
embodiment.
The ring magnets 13 have been placed in the stator yoke instead of in the
rotor pole. In
this way they do not rotate, and are not subjected to centrifugal forces. The
rotating iron
pole piece 12 and the rotor R' can be more perfectly balanced and become
simplified in
construction. The magnetic bearing coils 14 now surround the magnet directly
and
equally well modify the magnet's flux as before.
The fluid flow direction is labeled by arrow A in Fig. 6. Blood impinges on
inlet
fairing 29 and enters conduit 1 as in the axial flow pump. It diverges
radially at point pin
16 into the bearing fluid flow clearance 20. The static fluid pressure is
higher at the edge
32 than at the exit of passage 20 because the fluid velocity is smaller here.
This is
guaranteed by choosing the cross sectional area of inlet passage 47 to be
greater than
that of annular passage 44 as calculable by Bernoulli's Theorem. Gap 20 is
forcefully
washed out not only by this pressure differential but also by the remaining
suction at gap
20 created by the centrifugal impeller 39, which has a plurality of radial
blades. Fairing 29
is attached to the pump housing 46 using equally spaced thin vanes 37. The
right or
downstream side magnetic bearing yoke 28 at the outlet of the pump is held
centered in
the housing using outlet vanes 36 that may be straight or helically curved to
optimally
introduce fluid into the impeller vanes at their inside radius. Fluid is
propelled radially
m
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outwardly by impeller vanes 39 and then enters the circumferential volute
space 45, which
communicates with fluid exit opening 2T.
The bearing axial gap 20 on the right or outlet side is washed out by
directing a
portion of the inlet flow radially into the gap by use of scoop 9, which
extends above
titanium rotor housing 17 by an amount H, previously discussed in connection
with the
axial flow pump. Although suction exists from the impeller at scoop 9, an
equal suction
exists at the root of impeller blade 39 inside centrifugal housing 49. Thus,
this suction
does not cause flow to take place through slot 20. The near stagnation fluid
pressure at
scoop 9 is high. At the trailing end of the gap that terminates in gaps 43 and
43', at the
solid central portion of the impeller 39, the centrifugal force that is
imposed by gap
viscous drag will wash out gaps 43 and 43'. This creates an additional
negative pressure
at conduit 19. This differential pressure relative to the scoop 9 stagnation
pressure
forces flow down gap 20. Flow diverter 18 ensures smooth transition of radial
flow to
axial without stagnation points. The tip 18 can be located slightly off-center
to eliminate
stagnation at the tip. The asymmetric flow thus formed from an off-center tip
in conduit 19
forms a vortex to wash out housing 49 adjacent conduit 19 where flow
stagnation could
otherwise occur. A plurality of channels 41 in the impeller shaft 38 permit
communication
between gap 20 and conduit 19. The impeller shaft 38 has a radial clearance 42
with the
stationary bearing that is sufficiently large to prevent subhemolytic shear
stress in the
blood. Gaps 20, 43 and 43'are similarly sized. The shaft 38 is preferably made
of
titanium and is hermetically sealed to and supports pole pieces 12 . All
surfaces that
contact the blood are preferably titanium in this and in the axial flow
design. The titanium
may have a biolite carbon or wear-resistant coating of titanium nitride or
other coating to
is
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enhance bio-compatibility and gaulling resistance of bearing surfaces in the
event of
contact due to bearing failure or too large a shock load.
In the magnetic suspensions disclosed, the suspended rotor may touch down or
contact the housing if a control system failure should occur. This makes the
design
failsafe so it can continue to function until replaced. To start up the
bearing initially, the
axial control system is energized first to eliminate contact. Then the rotor
is spun up to
the desired rpm.
The magnetic bearings B, B' are, as noted, inherently axially unstable. That
is why
a VZP coil control system is needed. This instability force is used to good
advantage to
determine the axial force on the rotor. The differential pressure on the rotor
is then
calculated in a direct manner by dividing this force by the effective cross-
sectional area of
the rotor. This is shown in block diagram form in Fig. 8. The axial bearing
position is
monitored at 60, such as by measuring the coil inductance to determine the
associate
force. The desired differential pressure DP is calculated by dividing the
derived force by
the rotor area, at 64. In practice, a look up force table, which is useful for
non-linear data,
is not needed because the bearing force is sufficiently linear when large
fringing ring
blood gaps are used. See Fig. 18. To obtain transient force, the force
contribution of the
mass-inertia of the rotor is subtracted to provide the net external force.
As suggested, the inductance of the coils 14 in Fig.1 can be used to monitor
the
rotor axial position. The coil electronics for determining rotor position can
be simplified by
using a separate small auxiliary position coil 54 (Fig. 7). Ideally, the coil
54 is wound
beneath the main bearing coil 14, as shown in Fig. 7. While only a single
position coil is
needed, Fig. 7 shows a position coil 54 located in the inlet bearing B, while
a redundant
position coil 54' may be located in the outlet bearing B'. Less desirable is
an added on or
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separate non-contacting ultrasonic or magnetic axial position sensor that
would have to
be integrated with the pump. Once pump differential pressure is known, this
information
can be used to provide an accurate pulsatile flow as well as to provide a
basis for
physiologic control of the pump, in response to patient activity level.
When used in a person, cyclically varying the pump rpm at a given frequency
will
cause the pump differential pressure to vary at the same frequency. Since the
gage inlet
pressure to the pump, when extracting blood from the heart, is relatively
constant and low,
being only several mmhg, differential pressure is a good measure of the much
higher
pump discharge pressure. Varying pump rpm instantaneously varies pump flow
rate,
which changes outlet pressure to the body (when discharging into the aorta,
for example).
Pump differential pressure is monitored by the bearings. Any error in pump
differential
pressure amplitude is corrected by a feedback loop, which fine tunes the pump
rpm until
the desired pressure variation is obtained. This is shown schematically in the
control
system of Fig. 9. In Fig. 9, the desired cyclic differential pressure
variation is input as well
as by setting a cyclic frequency, at 66. A frequency discriminator 68
separates the
pump's differential pressure variation from that produced by the beating of
the natural
heart at a different frequency. The differential pressure amplitude across the
pump is
compared to the set point differential pressure at comparator 70. An error
signal at the
output of the comparator is input to a proportional amplifier 72 having a
selected gain K to
control the drive motor rpm 74, which, in turn, reflects variations in rotor
axial position, at
76, due to the associated variations in pressures.
A typical or normal set point pressure variation is 120/80. This means that
during
the pump cycle, blood pressure reaches a maximum of 120 mmhg (systolic
pressure) and
a minimum of 80 mmhg (diastolic pressure). By being able to achieve accurate
systolic
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and diastolic pressures, flow reversal through the pump is avoided and an
invasive
flowmeter or other sensor is not needed. The proportional gain labeled "K"
provides the
desired sensitivity of control to give a stable and accurate aortic pressure.
Physiologic control of pump flow rate is now possible based on differential
pressure.
If, for example, the patient's exercise level increases, his or her average
blood pressure
will decrease, due to a decrease in peripheral resistance with exercise, if
cardiac output is
not increased. Therefore, pump average rpm is merely increased in a feedback
loop until
the desired average blood pressure is maintained, which is typically 100 mmhg.
This will
automatically increase LVAD flow rate in order to maintain a constant average
aortic
pressure.
The patient's own heart rate can also be used as a parameter to set the
average
aortic pressure as well as the desired systolic/diastolic ratio. The patient's
own heart rate
may be monitored electrically as is done in pacemaker devices. However, this
is
undesirable because of the need to introduce heart monitoring electrodes. With
the
present invention the patient's heart rate can be directly monitored instead
by analyzing
the frequency content of the force on the magnetic bearings. If the cyclic
variation in
pump rpm (that produces pulsatile flow) is done at a frequency somewhat
different from
that of the natural heart, the heart's blood-pressure frequency can be
extracted from the
total differential pressure signal. Average pump output pressure can then be
made
proportional to the patient's heart rate which is a known physiological
function. The same
type of feedback loop can be used as in Fig. 9.
Thus, one can monitor the output pressure of the natural heart itself and how
well
the natural heart is pumping and recovering over time. One can wean a patient
with a
recovering heart off the LVAD by reducing LVAD flow contribution if desired.
Having heart
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and pump individual pressure data serves as a performance monitor. This is a
major
advantage for a long-term implantable device.
In the event that it is desired that the flow rate through the LVAD be
substantially
reduced - for example, in patients that can be weaned therefrom - we need to
still
guarantee full washout of the bearing gaps 20 shown in Figs. 1 and 7.
Referring to Fig. 7,
helical deep groove screw pumps are attached to each end face of the rotor R.
The inlet
screw item 52 and outlet screw 51 both pump fluid, toward the impeller and
away from it
respectively. The screw outside diameter has a clearance so that it does not
contact the
housing during normal operation. In the event of radial shock transient loads
that cannot
be sustained by the magnetic bearings, the screws are designed to contact the
housing
as mechanical back-up bearings. Wide axial lands on the screw thread outer
diameter
provide low contact stress bearing areas for this purpose. The flow that is
actively pumped
through the bearing gaps 20, 20' by the screws 51, 52 is sufficient so that
one need not
rely on pump-flow rate to generate differential pressures for washout. This
can be applied
to the centrifugal pump configuration in Fig. 6 as well.
The scoop 9 in Fig. 1 at the outlet bearing gap 20' has been eliminated in
Fig. 7
since screw pump 51 will produce sufficient flow without it. The option to
eliminate the
scoop, if desired, provides flexibility in design to eliminate thrombus
formation at this
location. The pointed tips of the screws 51 and 52 may be radially offset from
the pump
axis AP to avoid zones of blood stagnation at the tips.
Fig. 10 depicts an improved axial flow Archimedes screw impeller 80 for the
blood
pump that uses multiple short blades. A basic or conventional Archimedes screw
pump
(originally invented by the ancient Greek Archimedes for low-pressure crop
irrigation use)
consists of a single screw thread of several turns. At least one complete 360-
degree turn
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is needed to minimize back leakage in these pumps when substantial pressures
have to
be generated as in implantable blood pumps. To obtain one complete turn at
shallow
helix angles in particular, the screw thread axial length becomes quite long.
A long length
impeller cannot fit in a mini-size axial flow pump as taught in the present
invention.
Therefore, the impeller length L has to be reduced.
It has been verified with prototype testing by the inventor that an equivalent
hydraulic efficiency Archimedes screw can be made that is substantially
shorter in overall
length by using multiple short blades around the circumference. The more
blades that are
used, the shorter is the length. These blades have the same helix angles as a
single long
screw to give similar performance. Because multiple blades are used that
partially
overlap, back leakage flow through the impeller is identical to that of a
single 360 degree
one-turn screw of longer length. This has been experimentally verified.
Due to the finite transverse thickness of the blades, the addition of too many
blades results in a reduction of open flow area between the blades. This
decreases flow
rate and undesirably increases blood shear stress at a given flow rate. This
can cause
hemolysis. For this reason, the minimum number of blades needed to achieve the
desired result with an acceptably short axial length is ideal. A two-blade
impeller
decreases axial length to half, a three blade to a third and a four blade to a
fourth, with
diminishing returns beyond four blades.
Blade overlap D in Fig. 10 may be zero up to ideally several degrees of
circumference. This eliminates a straight through line of sight fluid leakage
path that
would otherwise exist between the trailing edge of one blade and the leading
edge of an
adjacent blade. Existence of this back leakage path would substantially lower
impeller
efficiency. In other words, if clearance D is negative, excess leakage will
occur at high
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pressures resulting in poor pump efficiency. In a preferred prototype a three-
blade
impeller of 0.80 in O.D. uses blades of 0.90 in. in length with 10 degrees of
overlap D.
Fig. 11 shows these blades as A4, B4 and C4 in end view. For the three blades
shown,
each extends 120 degrees + 10 degrees overlap = 130 degrees.
An Archimedes screw pumps fluid by the pushing action of the advancing screw
helix on the fluid in the surrounding housing and produces no lift. An
airplane or ship
propeller does not operate this way because it relies on the lift produced by
the blades
and there is no enclosed housing used. Overlapping the blades is not taught in
the
design of these propellers. This is particularly evident in two-blade
propellers where the
propellers are located 180 degrees apart and there is a large clearance
between the
blades. A lift generating propeller is a satisfactory design for these
applications because
high pressures are not generated in ship or aircraft propulsion, unlike in
blood pumps.
Figs. 12 and 13 illustrate a novel exit stator or diffuser that is designed to
redirect
the substantially tangential flow off the impeller into axial flow for
discharge at the pump
outlet. The vanes are curved almost 90 degrees. Flow enters nearly
tangentially. The
flow cross sectional area between vanes at the entrance into the diffuser is
kept
approximately the same at the impeller exit. This matches blood flow velocity
for minimal
blood cell trauma at the diffuser entrance. The rotational kinetic energy of
blood exiting
the impeller is converted into static pressure as flow traverses in the ever-
increasing flow
area between vanes. Minimum flow velocity exists at the diffuser exit where
flow is all
axial.
It is essential to accomplish tangential-to-axial flow redirection without
flow
separation on the back surface of the vanes. Otherwise, blood damage and
thromboemboli from turbulence can occur. This is accomplished by interposing
shorter
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auxiliary vanes 55 between primary curved vanes 56. These auxiliary vanes 55,
which do
not exist at the vane entrance, are set back a distance Y. The absence of the
vanes 55 at
the inlet or front of the diffuser provides the extra circumferential space
needed to obtain
the desired flow cross-sectional area for flow velocity matching. This is
essential because
the finite thickness T of the primary vanes requires a substantial space in
such a small
diameter mini-size pump. The resulting narrower spacing between vanes where
secondary vanes 55 are added is intended to be at the onset of the more highly
curved
portion of the primary vanes where flow separation tends to initiate. The
narrower gap W
forces the flow to hug the wall on the downstream convex surface of the vanes,
thereby
minimizing flow separation.
It is also desirable to minimize the number of vanes in order to reduce the
surface
area in contact with blood for minimizing thrombosis. Proper choice of vane
thickness T
can creates the desired average channel width W for a given number of vanes.
In a
preferred embodiment eight relatively thin vanes are used.
Flow separation is also reduced by using as large a radius of curvature "~" as
possible for the vanes as shown in Figs. 12 and 13. However, a larger radius
creates a
longer axial length for the diffuser. In a mini-pump a long axial length is
undesirable. A
shorter axial vane with small radius of curvature "r" can be employed without
flow
separation by using a smaller channel width W, provided blood shear stress
remains
acceptably low. In a preferred embodiment, minimal flow separation exists when
W
ranges from the 0.10 to 0.40 in and the radius of curvature "r" falls within
practical size
limits of 0.30 to 1.25 in.
Referring to Fig. 7, the impeller 4 is shown to have a longer axial length
than in Fig.
1. An advantage of this configuration is that the motor armature magnets,
which are
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located in dotted box 21, are completely located beneath the impeller blades.
Since the
impeller and magnets rotate together, no eddy currents are generated in the
blades due
to rotor rotation. This is not the case in Fig. 1 where a very short impeller
is shown. In this
configuration eddy currents will be dissipated in the stationary outlet stator
blades
because they are located in the field above the rotating magnets. This will
increase motor
power consumption.
Referring to Fig. 14, an alternate embodiment of the fringing ring magnetic
bearing
will be discussed because it offers additional advantages. A longitudinal
section through
the bearing B is shown in Fig. 14. A radially magnetized magnet 13" is
employed and
located between the circular fringing rings 11 and 11'. The magnet ring may be
composed of a plurality of radially magnetized pie-shaped sections or it can
be radially
magnetized in one piece. The bearing's axial instability force is only due to
the air gap flux
of the fringing rings. The magnet produces no instability force since its
field is contained
in the iron. So the one advantage of this configuration is its reduced axial
instability force
compared to the axially magnetized magnets of the type shown in Fig. 1. Also,
the inside
diameter of the bearing is now substantially open space, allowing more compact
incorporation in smaller pumps for use with children.
While this invention has been described in detail with particular reference to
preferred embodiments thereof, it will be understood that variations and
modifications will
be effected within the spirit and scope of the invention as described herein
and as defined
in the appended claims.
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