Source: https://patents.google.com/patent/WO2006133209A1/en
Timestamp: 2019-05-21 01:01:24
Document Index: 720319545

Matched Legal Cases: ['art 200', 'art 200', 'art 200', 'art 200', 'art 200', 'art 200', 'art 200']

WO2006133209A1 - Blood pump - Google Patents
WO2006133209A1
WO2006133209A1 PCT/US2006/021955 US2006021955W WO2006133209A1 WO 2006133209 A1 WO2006133209 A1 WO 2006133209A1 US 2006021955 W US2006021955 W US 2006021955W WO 2006133209 A1 WO2006133209 A1 WO 2006133209A1
PCT/US2006/021955
2005-06-06 Priority to US60/687,659 priority
2006-06-06 Application filed by The Cleveland Clinic Foundation, Foster-Miller, Inc. filed Critical The Cleveland Clinic Foundation
2006-12-14 Publication of WO2006133209A1 publication Critical patent/WO2006133209A1/en
This application claims the benefit of U.S. Provisional Application No. US 60/687,659, filed June 6, 2005. Government Rights
The invention described in this application was supported, at least in part, by United Stated Government Contract Nos. HHSN268200448188C and HL67487 with the National Heart, Lung and Blood Institute and the National Institutes of Health. Technical Field
Children who require mechanical circulatory support after failing routine medical management represent the most critically ill subset of an already challenging patient population. As in adult patients, pediatric patients can now benefit from some of the exciting advances that are occurring in the field of mechanical support for cardiorespiratory failure. The pediatric population has not, however, received the same attention in terms of product development, as has the adult population. For example, currently there are no pulsatile or implantable VADs available for infants and small children in the United States, while at many centers ECMO remains their only available form of mechanical circulatory support. In addition, unique features of circulatory failure in children limit the applicability of advances made in device development for adults. Accordingly, there is a need for focused research and development leading to devices that provide circulatory support for children with full consideration of the anatomic and physiologic requirements unique to pediatrics. One consideration in the design and development of circulatory support systems for children is related to patient size. It is desirable for the pediatric mechanical circulatory support device to provide support across a large range of patients sizes - from newborns to young adults and through adulthood. Paracorporeal VADs that are currently available for children in Europe rely on a number of pump sizes to cover the range of patients encountered in pediatric practice, which substantially increases both development and patient costs. Also, paracorporeal systems result in major skin penetrations, and expose the circulatory flow path to risk of mechanical damage. Beyond implications for the pump itself, size considerations exist for all aspects of device design for children including cannulas, energy sources and control mechanisms. hi addition to considerations of patient size, the design of circulatory support systems for children takes into account other physiologic considerations unique to pediatrics. Children, especially newborns, may be more prone to complications related to anticoagulation. Higher doses of anticoagulation medications required for ECMO may make intracranial hemorrhage more common resulting in poorer neurologic outcomes compared to VAD supported children.
Therefore, it is desirable that the pediatric circulatory support system operates with minimal or no anticoagulation. Children are vulnerable to infectious complications and, as a result, a large percentage of children who die during mechanical circulatory support are those who succumb to infection. A large percentage of children require the urgent institution of support to treat cardiac arrest after cardiac surgery or in the setting of acute myocarditis. Therefore, it is desirable that designs for the circulatory support system allow for rapid deployment, which has been shown to substantially improve outcomes for children requiring support for cardiac arrest. Newborns often manifest an exaggerated systemic inflammatory response after cardiopulmonary bypass, which frequently evolves into multi-system organ failure during prolonged ECMO or VAD support. Therefore, it is desirable that the circulatory support system has maximal bioconipatibility to help prevent activation of systemic inflammatory cascades by providing minimal trauma to blood elements and possibly by providing pulsatile perfusion. -A-
The present invention relates to a blood pump that includes a stator assembly including a fluid inlet and a fluid outlet. The pump includes a rotor assembly including an impeller rotatable about an axis to move fluid from the inlet to the outlet. The pump also includes a motor that imparts rotation of the impeller about the axis. The motor includes a motor stator fixed to the stator assembly, a motor rotor fixed to the rotor assembly, and a radial motor gap between the stator and the rotor. The pump is configured to direct blood flow through the fluid inlet to the fluid outlet and a wash flow through the motor gap. The present invention also relates to a blood pump that includes a stator assembly including a fluid inlet and a fluid outlet. The pump includes a rotor assembly including an impeller rotatable about an axis to move fluid from the inlet to the outlet. The pump also includes at least one permanent magnet radial bearing for supporting the rotor assembly for rotation about the axis. The radial bearing includes at least one permanent magnet fixed to the housing and at least one permanent magnet fixed to the rotor assembly. The stator magnets and rotor magnets are axially offset from each other to produce magnetic forces which balance with hydrodynamic forces created by the pumping action of the impeller. The present invention also relates to an implantable blood pumping apparatus that includes a pump including a housing with a fluid inlet and a fluid outlet. The pump is operable to move fluid from the inlet to the outlet. An outflow sheath directs the flow along the outside of the pump.
The present invention further includes a blood pump including a housing including a fluid inlet and a fluid outlet. The pump also includes an impeller rotatable about an axis to move fluid from the inlet to the outlet. The pump further includes an inflow stator having vanes with a curvature reversed from the curvature, of vanes on the impeller. Brief Description of the Drawings
Fig. 2 A is a sectional view of a blood pump of the blood pumping system of Fig. 1; Figs. 2B and 2C are sectional views illustrating an alternative configuration of the blood pump of Fig. 2A;
Fig. 2D is a magnified view of a portion of the blood pump of Fig. 2 A; Fig. 2E is a sectional view illustrating an alternative configuration of the blood pump of Fig. 2 A; Figs. 2F-H are a magnified views of a portion of the blood pump of
Fig. 3 A is a sectional view of a portion of the blood pump of Fig. 2 A; Fig. 3B is a top view of a portion of the blood pump of Fig. 2 A; Fig. 3 C is a sectional view illustrating an alternative configuration of the blood pump of Fig. 2 A;
Figs. 5Aand 5B illustrate different implementations of the pump of Fig. 2A; Fig. 5C illustrates a guide wire feature of the pump of Fig. 2A; Fig. 5D illustrates another implementation of the pump of Fig. 2 A;
Figs. 7A-7C illustrate different implementations of the pump of Fig. 6; Fig. 8 A is a sectional view of the pump of Fig. 2A outfitted with an outflow sheath in accordance with a third embodiment of the present invention; Fig. 8B is a sectional view illustrating an alternative construction of the pump of Fig. 8 A;
Fig. 9 illustrates the pump of Fig. 8 A in an activated condition;
Figs. 1OA and 1OB are charts illustrating pressure vs. flow characteristics for test configurations of the pumps of Figs. 2A and 6, respectively; and
Figs. 1 IA-I IF are charts illustrating the effects of an inflow stator configuration of the present invention.
The present invention relates to a blood pump. In the embodiments illustrated herein, the blood pump is depicted as an implantable blood pump for use as a ventricular assist device (VAD). The pump of the present invention provides an implantable adult or pediatric ventricular assist device that may be used for short to long-term applications. Through flexible implant approaches, the pump is adaptable to patient size and to the special anatomic features that may be encountered when treating congenital heart disease. The pump may be implemented as a Right Ventricular Assist Device (RVAD), a Left Ventricular Assist Device (LVAD), or a Bi- Ventricular Assist Device (BVAD), with intravascular and intracorporeal extravascular implant options for each implementation. This flexibility provides the surgeon great freedom in matching the procedure with the range of patient size and anatomical variations found in congenital heart disease.
Fig. 1 illustrates an example configuration of a system 10 that includes a mixed flow pump 20 of the present invention. As used herein the term "mixed flow pump" is meant to describe a pump in which, as fluid flows through the impeller, the fluid has significant velocity imparted in both axial and radial directions.
Referring to Fig. 2A, the pump 20 includes a housing 22 with an inlet port 24, one or more radial outlet ports 26, and a wash flow port 28. The housing 22 has an open first end 30 that forms the inlet port 24 and a closed opposite end 32. The pump 20 includes an impeller 40 that is supported on a shaft 42 that is rotatable about an axis 44 of the pump. An inflow stator 46 is centered on the axis 44 and is positioned in the inlet port 24 adjacent the impeller 40. The impeller 40, shaft 42, and inflow stator 46 are constructed of non- ferrous materials, such as stainless steel, titanium, ceramics, polymeric materials, composite materials, or a combination of these materials. In one particular embodiment, the shaft 42 may be constructed of a Zirconia material.
The pump 20 includes a motor portion 50 that is adapted to impart rotation of the shaft 42 and impeller 40. The motor 50 may be any suitable electric motor, such as a multi-phase motor in which each phase is excited via pulse-width modulated voltage provided by the control unit 12. The motor 50 includes a stator 52 supported by the housing 22 and a rotor 54 supported on the shaft 42. The stator 52 comprises one or more poles or windings, such as copper wire windings, wound on a stator core. The rotor 54 comprises one or more permanent magnets, such as Neodymium Iron Boron (NdFeB) magnets, arranged in a cylindrical fashion on the shaft 42 and extending coaxially with the shaft. The control unit 12 is operative to supply motor control voltage to the motor stator 52 to excite the windings and induce rotation of the rotor 54. Referring to Figs. 3A and 3B, in one particular embodiment of the pump 20, the motor 50 has a four (4) pole, three (3) coil configuration. As shown in Fig. 3 A, the rotor 54 includes a back iron 76 having a cross-shaped cross section that defines recesses having perpendicularly oriented rectangular surfaces in which the permanent magnets 60 are received and supported. In the four pole configuration, the rotor 54 includes four permanent magnets 60 spaced equally about the shaft 42. As shown in Fig. 3A, the rotor 54 has an overall cylindrical configuration.
Referring to Fig. 2 A, the pump 20 also includes radial bearings 100 that help support the shaft 42 and impeller 40 for rotation about the axis 44. In the illustrated embodiment, the radial bearings 100 include a front radial bearing 102 and a rear radial bearing 104 positioned adjacent opposite ends of the motor 50. The radial bearings 100 are permanent magnet bearings that utilize permanent magnets, such as NdFeB magnets. Each radial bearing 100 comprises a plurality of ring-shaped stator magnets 106 and a plurality of ring-shaped rotor magnets 108. hi the embodiment of Fig. 2 A, the front radial bearing 102 and rear radial bearing 104 each include ten stator magnets 106 and ten rotor magnets 108. The radial bearings 100 could have any desired number of stator and rotor magnets. The implementation of the permanent magnet radial bearings 100 helps eliminate the need for a seal, as is required with conventional mechanical radial bearings.
Mating or engaging surfaces of the front and rear axial bearings 142 and 144 may be coated or constructed with materials that produce low friction, such as Teflon®, diamond-like carbon coatings, ceramics, titanium, and diamond coated titanium. In one particular example, the axial bearing surfaces of the rotor assembly 120, i.e., the portions 150 and 154, are coated or otherwise formed with a chrome-cobalt material, and the axial bearing surfaces of the stator assembly 122, i.e., the portions 152 and 156, are coated or otherwise formed of a ceramic material, which has been shown to provide performance superior to that of conventional bearing surfaces, such as ceramic-on-ceramic bearing surfaces or diamond-like carbon-on-diamond-like carbon bearing surfaces, hi another example, the axial bearing surfaces of the rotor assembly 120, i.e., the portions 150 and 154, are coated or otherwise formed with a synthetic jewel material (e.g., synthetic ruby, sapphire, or diamond materials), and the axial bearing surfaces of the stator assembly 122, i.e., the portions 152 and 156, are coated or otherwise formed of a ceramic material.
The pump 20 is constructed such that parts that come into contact with blood are made of a biocompatible material. The motor magnets 60, back iron 76, and radial bearing rotor magnets 108 are encased or otherwise covered or coated on the shaft 42 by a biocompatible material 110. Examples of such materials are titanium and stainless steel. The motor stator 52, i.e., the stator core 64 and windings 62, and the radial bearing stator magnets 106 are also encased or otherwise covered or coated on the housing 22 by a biocompatible material 112. Further, the impeller 40 and inflow stator 46 are constructed, encased, or otherwise covered or coated with a biocompatible material. For example, the impeller 40 and inflow stator 46 maybe constructed of titanium or molded from a biocompatible polymeric material.
The inlet stator 46 may have a vane configuration with a curvature reversed from that of the vanes of the impeller 40. This helps produce a reverse pre-swirl in the inflow blood, i.e., a swirl in the blood in a direction opposite the rotation of the impeller 40. Testing has shown that a pre-swirl created in the inflow blood by the inlet stator 46 helps improve the performance characteristics of the pump 20. Figs. HA HF illustrate selected performance characteristics for a pump configured with the reversed curvature inlet stator 46 of the present invention versus a pump configured with a conventional non-curved or straight inlet stator. m the tests used to gather the data shown in Figs. 1 IA-I IF, the test pump was operated at a nominal speed of 60,000 RPM. To perform the tests, the pump was operated at this nominal speed pumping a fluid having a composition that simulates blood. An outlet conduit connected to the pump was clamped to restrict outlet flow from the pump. The pump was then operated at the nominal speed, the clamp was systematically opened to predefined positions, and data readings were taken at each position to gather the data points in Figs. 1 IA-I IF. Thus, in Figs. 1 IA-I IF, data point pairs for the reverse curved and straight inlet vane configurations correspond to these predefined clamp positions. For example, in
Figs. 1 IA-I IF, the data points on the far right ends of the curves correspond to the last of the predefined clamp positions. Going backward or to the left in Figs. 1 IA-I IF, the next-to-last data points correspond to the next-to-last predefined clamp position, and so on. For purposes of this description, a flow of three (3) liters per minute (LPM) at a 90 rnniHg pressure rise across the pump are used as nominal or baseline performance characteristics for purposes of comparing the different inlet stator configurations.
Fig. 1 IA illustrates stage pressure rise versus flow characteristics for a pump fit with a curved inlet stator 46 at the line indicated at 400 versus a pump fit with a conventional or non-curved inlet stator at the line indicated at 402. The stage pressure rise is the inlet pressure measured immediately before the stator vane within the shroud diameter, subtracted from the outlet pressure measured in the outlet chamber representative of the aorta.
Fig. 1 IB illustrates adjusted stage pressure rise versus flow characteristics for a pump fit with a curved inlet stator 46 at the line indicated at 404 versus a pump fit with a conventional or non-curved inlet stator at the line indicated at 406.
For comparison, the non-adjusted values from Fig. 1 IA are included in Fig. 1 IB at 400 and 402. The adjusted stage pressure rise is the estimated pressure just outside the pump inlet subtracted from the outlet pressure measured in the outlet chamber representative of the aorta. The estimated pressure outside the pump inlet is calculated by subtracting reentrant flow losses due to pump insertion into a larger cavity from the measured inlet pressure.
Referring to Figs. 1 IA and 1 IB, it can be seen that, other conditions being equal, the pump outfitted with the reversed curved vane inlet stator is capable of achieving the 3 LPM flow at a pressure rise far in excess of the nominal value of 90 mmHg. In comparison, in the same conditions, the straight vane inlet stator falls to meet the 3 LPM flow.
Fig. HC illustrates adjusted motor current versus flow characteristics for a pump fit with a curved inlet stator 46 at the line indicated at 410 versus a pump fit with a conventional or non-curved inlet stator at the line indicated at 412. The adjusted motor current is the free running speed current subtracted from the recorded motor current. Fig. 1 ID illustrates estimated motor torque versus flow characteristics for a pump fit with a curved inlet stator 46 at the line indicated at 414 versus a pump fit with a conventional or non-curved inlet stator at the line indicated at 416. The adjusted motor torque is the adjusted motor power divided by pump speed. Adjusted motor power is the adjusted motor current multiplied by the supply voltage.
Fig. 1 IE illustrates stage efficiency versus flow characteristics for a pump fit with a curved inlet stator 46 at the line indicated at 420 versus a pump fit with a conventional or non-curved inlet stator at the line indicated at 422. The stage efficiency is the non-adjusted hydraulic power divided by the adjusted motor power.
Fig. 1 IF illustrates adjusted stage efficiency versus flow characteristics for a pump fit with a curved inlet stator 46 at the line indicated at 424 versus a pump fit with a conventional or non-curved inlet stator at the line indicated at 426. For comparison, the non-adjusted values from Fig. 1 IE are included in Fig. 1 IB at 420 and 422. The adjusted stage efficiency is the adjusted hydraulic power divided by the adjusted motor power. The adjusted hydraulic power is the adjusted stage differential pressure rise multiplied by flow. The adjusted stage differential pressure is determined by subtracting reentrant flow losses due to pump insertion into a larger cavity from measured inlet pressure. Non-adjusted stage efficiency takes into account only the adjusted motor power.
As shown in Figs. 1 IC-I IF, the reversed curved vane inlet stator had higher current and torque ratings for corresponding conditions and also proved to have better efficiency while pumping at 3 LPM. From the data of Figs. 1 IA-I IF, it will be appreciated that the reversed curve inlet vane configuration improves the overall performance of the pump in comparison with a conventional straight vane inlet vane configuration. Thus, at the same speed, a pump fitted with the reversed curve inlet vanes will have a higher output flow. Similarly, to achieve the same output, the pump fitted with the reversed curve inlet vanes will operate at a lower speed. Because, of this, blood shear and resulting thrombosis fomiation can be reduced. This may also help reduce pump power consumption and extend battery life.
Referring to Figs. 2B and 2C, the impeller 40 may include a shroud 48 that helps to further improve the pump performance. The shroud 48 has a generally cylindrical configuration and may be formed as a single piece of material with the impeller 40 or may be formed separately and subsequently attached to the impeller.
The shroud 48 adds damping which helps stabilize the dynamics of the impeller 40 and/or rotor assembly 120.
Referring to Figs. 4A and 4B, the radial bearings 100 operate on a repulsive force principle. Each pair of permanent magnet (PM) rings 106 and 108 has north and south poles aligned in the radial direction. In operation, the radial bearings 100 help overcome the side pull of the motor 50 and maintain the rotor assembly 120 suspended relative to the stator assembly 122. The radial bearings 100 also have an axial stiffness that, in combination with hydraulic forces, helps determine the position of the rotor assembly 120 relative to the stator assembly 122. To increase the bearing stiffness, the neighboring PM stator rings 106 and rotor rings 108 are placed in opposing polarity, i.e., north-to-north and south-to-south. The non- ferromagnetic construction of the pump components adjacent the radial bearings 100 helps maintain the magnetic flux paths of the magnets 106 and 108, which helps achieve a relatively low axial side pull during operation of the pump 20. The PM stator magnets 106 may extend 360° about the rotor assembly 120. Alternatively, one or more of the PM stator magnets 106 may extend less than 360° about the rotor assembly 120. This may help produce a net magnetic force that helps stabilize the submerged rotor assembly 120 during use. Figs. 4 A and 4B illustrate an unstable equilibrium condition and an axially offset condition, respectively, of the radial bearings 100. Referring to Fig. 4A, in the unstable equilibrium condition of the radial bearings 100, the magnetic poles of the rotor magnets 108 and stator magnets 106 are axially aligned with each other. This is the desired condition of the radial bearings 100 during operation of the pump 20 because, when the bearings are in this position, the rotor assembly 120 is in a position in which the axial bearings 140 are not loaded. The magnetic flux paths resulting from this arrangement are indicated generally by the arrows in the rotor magnets 108 and stator magnets 106. In this axially aligned position, the flux paths are aligned and the attractive/repulsive forces of the magnets 106 and 108 acting on the stator assembly 122 and rotor assembly 120 are radial in nature, as shown by the arrows identified at 170 in Fig. 4A.
Referring to Fig. 4B, in the axially offset condition of the radial bearings 100, the magnetic poles of the rotor magnets 108 and stator magnets 106 are offset from each other along the axis of rotation 44. This distance may be relatively small (e.g., .0002-.002 in.). This is the pre-loaded, axially offset condition prior to operation of the pump 20. The magnetic flux paths resulting from this arrangement are indicated generally by the arrows in the rotor magnets 108 and stator magnets 106. In this axially offset position, the flux paths are misaligned and the attractive/repulsive forces of the magnets 106 and 108 acting on the stator assembly 122 and rotor assembly 120 have radial components, as shown by the arrows identified at 172 in Fig. 4B, and axial components, as shown by the arrows identified at 174 in Fig. 4B.
According to the present invention, the pump 20 is constructed to produce a net axial force that urges the rotor assembly 120 to move axially relative to the stator assembly 122 to the axially offset condition of Fig. 4B. To achieve this, the rear stop 156 of the rear axial bearing 144 and the front stop 152 of the front axial bearing 142 are moved rearward from the positions that would maintain the radial bearings 100 at the unstable equilibrium point. As a result, when the pump 20 is at rest, the rotor assembly 120 moves rearward against the rear stop 156 under the net axial pull of the radial bearing magnets 106 and 108 to the axially offset condition of Fig. 4B. According to the present invention, the thrust of energy transfer to the fluid by the impeller 40 and the static pressure gradient front to back on the rotor assembly 120 produce hydrodynamic forces that counteract the net axial force of the radial bearing misalignment and help move the magnets 106 and 108 toward the unstable equilibrium condition of Fig. 4A. In operation of the pump 20, fluctuations in applied load, such as those resulting from the natural heart beat of the patient, result in a cyclical front-to-back oscillation of the rotor assembly 120 relative to the stator assembly 122. This helps cycle the loads on the axial bearings 140, which helps reduce friction and heat in the bearings and also helps produce a cyclical washing of the bearings. As a result, these cyclical loads help prevent thrombosis formation in the pump 20 by permitting cyclical washing at the front and rear stops 152 and 156.
Referring to Figs. 2F-2H, the stop point 156 includes a permanent magnet axial bearing 160 that exerts an axial force on the rotor assembly 120. The force exerted on the rotor assembly 120 by the bearing 160 opposes axial forces placed on the rotor assembly by the radial bearings 100 and helps eliminate occasional mechanical contact at the stop point 156. The stop point 156 also includes surface profiles, such as recesses 162. As shown in Figs. 2G and 2H, the surface profiles 162 have a generally concave curved configuration and are recessed into the surface of the stop point 156. The profiles 162 help generate hydrodynamic lifting forces that help minimize or eliminate contact forces when the rotor 120 comes very close to the stop point 156. These hydrodynamic forces help counteract the residuals from the summing of the other axial forces acting on the rotor 120. The pump 20 may be configured for a number of different implementations, including intravascular and intracorporeal extravascular implementations, as appropriate for patient size. Intravascular implementations may be used for larger patients, such as larger pediatric patients through adolescence and adulthood. Intracorporeal extravascular implementations may be used for smaller patients, such as neonatal and very young pediatric patients. The pump 20 illustrated in the embodiment of Figs. 1-3 is configured for intravascular implementations. Examples of these intravascular implementations are shown in Figs. 5A-5C.
Referring to Fig. 5 A, the pump 20 is shown in an intravascular implementation as a right ventricular assist device (RVAD). In the RVAD implementation, the pump 20 is inserted into the heart 200 through an incision 202 in the pulmonary artery 206 at the intersection of the pulmonary trunk 204 and the pulmonary artery. The pump 20 is positioned with the inlet 24 extending through the pulmonary semilunar valve 212 into the right ventricle 210 and the outlet 26 positioned in the pulmonary trunk 204. In operation, the pump 20 operates as described above to assist the right ventricle 210 in pumping blood to the pulmonary artery 206. Referring to Fig. 5B, the pump 20 is shown in an intravascular implementation as a left ventricular assist device (LVAD). In the LVAD implementation, the pump 20 is inserted into the heart 200 through an incision 220 in the aorta 222. The pump 20 is positioned with the inlet 24 extending through the aortic semilunar valve 226 into the left ventricle 224 and the outlet 26 positioned in the aorta 222. In operation, the pump 20 operates as described above to assist the left ventricle 224 in pumping blood to the aorta 222.
Referring to Figs. 5A and 5B, the guide wire 230 and pump 20 are inserted into the heart 200 through the incisions 202 and 220. The guide wire 230 may be advanced forward of the pump 20 and guided to the desired location in the organ, i.e., the right ventricle 210 or left ventricle 224. The pump 20 can then be delivered to the desired location using the stiffened guide wire 230 to maneuver and guide placement of the pump. The position of the pump 20 can then be adjusted by sliding the sheath 232 of the power cable 14 over the guide wire 230. Referring to Fig. 5D, two pumps 20 are shown in an intravascular implementation as bi-ventricular assist devices (BVAD). Essentially, the BVAD implementation incorporates two pumps 20 arranged in the RVAD and an LVAD implementations described above in Figs. 5A and 5B. In Fig. 5D, the guide wire 230 of Figs. 5A-5C is not shown for purposes of illustrating the pumps 20 with out this feature. The guide wire 230 of Figs. 5A-5C is suited for use in the
BVAD implementation of Fig. 5D. Thus, in the BVAD implementation, an RVAD pump 2OR is inserted through an incision 202 in the pulmonary artery 206 and is oriented with the inlet 24 positioned in the right ventricle 210 and the outlet 26 positioned in the pulmonary trunk 204. An LVAD pump 2OL is inserted through an incision 220 in the aorta 222 and is oriented with the inlet 24 positioned in the left ventricle 224 and the outlet 26 positioned in the aorta 222. In operation, the RVAD pump 2OR assists the right ventricle 210 in pumping blood to the pulmonary artery 206 and the LVAD pump 2OL assists the left ventricle 224 in pumping blood to the aorta 222.
A second embodiment of the present invention is illustrated in Fig. 6. The second embodiment of the invention is similar to the first embodiment of the invention illustrated in Figs. 1-5D. Accordingly, numerals similar to those of Figs. 1-5D will be utilized in Fig. 6 to identify similar components, the suffix letter "a" being associated with the numerals of Fig. 6 to avoid confusion. According to the second embodiment, the pump 20a is configured for intracorporeal extravascular RVAD, LVAD, or BVAD implementations. To accomplish this, the pump 20a of the second embodiment includes an attached catheter or cannula that facilitates insertion in the heart and a catherter or graft to facilitate connection to the vasculature. The catheter or cannula is axially deforniable, radially non- collapsible, and impermeable under the physiological and biological conditions associated with the blood pump usages described herein.
The configuration of the pump head housing 250 of the second embodiment helps facilitate extravascular implementations of the pump 20a. More particularly, the pump head housing 250 helps facilitate discharging blood along the outside diameter of the motor/bearing housing 22a into the outlet cannula 254. The configuration of Fig. 6 permits wash flow in the motor gap 34a through the wash flow ports 28a under the influence of arterial pressure. As an additional feature of the embodiment of Fig. 6, the primary flow, being contained within the outlet cannula 254 next to the motor 50a and motor housing 22a, may also have some enhanced cooling effects on the motor. Since the primary flow of the pump 20a is outside the pump rather than through the motor gap 34a, the motor gap can be kept at a minimum size, which helps reduce the overall diameter and size of the pump. Figs. 7A-7C illustrate intracorporeal extravascular implementations of the pump 20a of Fig. 6. Referring to Fig. 7A, the pump 20a is shown in an intracorporeal extravascular RVAD implementation, hi this RVAD implementation, the pump 20a is implanted in the patient next to the heart 200a. The outlet cannula 254 is connected via incision 202a to the pulmonary artery 206a at the intersection of the pulmonary trunk 204a and the pulmonary artery. The inlet cannula 252 is connected via incision 282 to the right atrium 280 or, alternatively, the right ventricle 210a. In operation, the pump 20a operates as described above to assist the right ventricle 21 Oa by pumping blood from the right atrium 280 through the inlet cannula 252 to the pulmonary artery 206a via the outlet cannula 254.
Referring to Fig. 7B, the pump 20a is shown in an intracorporeal extravascular LVAD implementation, hi this LVAD implementation, the pump 20a is implanted in the patient next to the heart 200a. The outlet cannula 254 is connected via incision 220a to the aorta 222a. The inlet cannula 252 is connected via incision 286 to the apex 284 of the left ventricle 224a or, alternatively, the left atrium. In operation, the pump 20a operates as described above to assist the left ventricle 224a by pumping blood from the left ventricle through the inlet cannula 252 to the aorta 222a via the outlet cannula 254. Referring to Fig. 7C, two pumps 20a are shown in an intracoiporeal extravascular implementation as bi-ventricular assist devices (BVAD). Essentially, the BVAD implementation incorporates two pumps 20a arranged in the RVAD and an LVAD implementations described above in Figs. 7A and 7B. An RVAD pump 20Ra is implanted in the patient next to the heart 200a. The outlet cannula 254R is connected via incision 202a to the pulmonary artery 206a and the inlet cannula 252R is connected via incision 282 to the right atrium 280 or, alternatively, the right ventricle. An LVAD pump 20La is implanted in the patient next to the heart 200a. The outlet cannula 254L is connected via incision 220a to the aorta 222a and the inlet cannula 252L is connected via incision 286 to the apex 284 of the left ventricle 224a or, alternatively, the left atrium. In operation, the RVAD pump 20Ra assists the right ventricle 210a by pumping blood from the right atrium 280 through the inlet cannula 252R to the pulmonary artery 206a via the outlet cannula 254R. hi operation, the LVAD pump 20La assists the left ventricle 224a by pumping blood from the left ventricle through the inlet cannula 252L to the aorta 222a via the outlet cannula 254L.
The sheath 300 allows for reducing the overall size of the pump 20b. For reference, referring back to the embodiment of Figs. 5A-5D, those skilled in the art will appreciate that, for intravascular implementations of a pump that is not fit with a sheath 300, the pump extends through the heart valve and is positioned with the inlet and outlet positioned on opposite sides of the valve. For example, in an LVAD implementation, the pump extends through the heart valve with the inlet positioned in the left ventricle and the outlet positioned in the aorta. As another example, in an RVAD implementation, the pump extends through the heart valve with the inlet positioned in the right ventricle and the outlet positioned in the pulmonary trunk. As shown in Figs. 5 A-5D, to achieve these extents, the pump has a configuration in which the inlet is extended to reach into the heart chamber while the outlet is positioned on the opposite side of the heart valve. Those skilled in the art, however, will appreciate that this may result in an unwanted pressure drop on the inlet side of the pump.
Referring to Fig. 9, according to the present invention, the sheath 300 functions to extend the outlet of the pump 20b, which eliminates the need to extend the inlet. Fig. 9 illustrates an implementation of the pump 20b of Fig. 8 A. Those skilled in the art, however, will appreciate that the pump of Fig. 8B may also be used in the implementation of Fig. 9. In the LVAD implementation shown in Fig. 9, the inlet 24b and outlet 26b of the pump 20b are positioned in the heart chamber, i.e., the left ventricle 224b. The sheath 300, however, extends through the heart valve 226b into the aorta 222b and thereby effectively places the outlet in the aorta. It will be appreciated that, using this technique, the need for an inlet extension, and any resulting pressure drop, can be eliminated.
The pump 20 also incorporates features that help provide high thrombus resistance without anticoagulation. One such feature is that all surfaces are continuously washed with flowing blood. There are no dead end spaces or crevice- like geometries. The back and forth oscillation of the rotor helps ensure that the blood contacting surfaces inside the pump, including the front and rear stop points 152 an 156, are washed. Also, most surfaces are slightly heated, which helps inhibit platelet aggregation. Further, the Teflon® and diamond-like carbon coatings applied to various pump surfaces may also help prevent coagulation.
Another coating that may be used to help prevent coagulation is a synthetic cell membrane material.
The construction of the pumps 20 , 2OA and 2OB disclosed herein have small package sizes in comparison with other implantable VADs. This allows for implementation of the pump 20 in the various intravascular and intercorporeal extravascular LVAD, RVAD, and BVAD scenarios described above. The small package size of the pump 20 is made possible by a variety of factors. One such factor is that the primary flow of the pump 20 being placed outside the pump. Another factor is that the pump 20, operating at high RPM (up to 60,000 RPM or more), is able to produce a relatively high output from a relatively small displacement volume. Example configurations illustrating small package size characteristics of the pumps 20 and 2OA are set forth in Table 1 : Table 1
As shown in Figs. 1OA and 1OB, even with the small package sizes shown in Table 1, the intravascular pump (see Fig. 2A) and the intracorporeal pump (see Fig. 6) are easily capable of operating at or around the nominal performance ratings for flow (3 LPM) and pressure (90 mmHg).
1. A blood pump comprising: a stator assembly comprising a fluid inlet and a fluid outlet; a rotor assembly comprising an impeller rotatable about an axis to move fluid from the inlet to the outlet; and a motor for imparting rotation of the impeller about the axis, the motor comprising a motor stator fixed to the stator assembly, a motor rotor fixed to the rotor assembly, and a radial motor gap between the stator and the rotor; the pump being configured to direct a primary flow from the fluid inlet to the fluid outlet over an outside diameter of the motor and being configured to direct a wash flow through the motor gap.
2. The blood pump recited in claim 1 , wherein the rotor assembly further comprises at least one permanent magnet radial bearing for supporting the rotor assembly for rotation about the axis.
10. The blood pump recited in claim 3, wherein: the radial bearing stator comprises a plurality of ring shaped stator magnets arranged next to each other in opposing polarity; and the radial bearing rotor comprises a plurality of ring shaped rotor magnets arranged next to each other in opposing polarity; the pump being configured such that, during operation, the stator magnets and rotor magnets are positioned with like polarities opposing each other.
12. The blood pump recited in claim 1 , further comprising front and rear axial bearings comprising a surface on the rotor and a mating surface on the stator assembly.
21. The blood pump recited in claim 1 , further comprising means for measuring a pump internal temperature as a pump control input.
27. The blood pump recited in claim 1 , wherein the rotor has a 2-pole magnetic geometry.
29. The blood pump recited in claim 1 , wherein the outer surface defining the radial motor gap includes at a least a portion having a non-cylindrical profile.
34. The blood pump recited in claim 1 , further comprising an inflow cannula that is axially deformable, radially non-collapsible, and impermeable.
35. The blood pump recited in claim 1 , further comprising a biocompatible power cable configured to help support and advance the blood pump along a blood vessel.
42. A blood pump comprising: a stator assembly comprising a fluid inlet and a fluid outlet; a rotor assembly comprising an impeller rotatable about an axis to move fluid from the inlet to the outlet; and at least one permanent magnet radial bearing for supporting the rotor assembly for rotation about the axis, the radial bearing comprising at least one permanent magnet radial bearing stator fixed to the housing and at least one permanent magnet radial bearing rotor fixed to the rotor assembly, the stator magnets and rotor magnets being axially offset from each other to produce magnetic forces which balance with hydrodynamic forces created by the pumping action of the impeller.
43. An implantable blood pumping apparatus comprising: a rotary dynamic pump having a housing with a fluid inlet and a fluid outlet, the pump being operable to move fluid from the inlet to the outlet; and an outflow sheath for directing the pumped flow along the outside of the pump.
44. A blood pump comprising: a housing comprising a fluid inlet and a fluid outlet; an impeller rotatable about an axis to move fluid from the inlet to the outlet; and an inflow stator having vanes with a curvature reversed from the curvature of vanes on the impeller.
PCT/US2006/021955 2005-06-06 2006-06-06 Blood pump WO2006133209A1 (en)
US60/687,659 2005-06-06
EP06784606.3A EP1898971B1 (en) 2005-06-06 2006-06-06 Blood pump
AU2006255059A AU2006255059A1 (en) 2005-06-06 2006-06-06 Blood pump
CA002611313A CA2611313A1 (en) 2005-06-06 2006-06-06 Blood pump
WO2006133209A1 true WO2006133209A1 (en) 2006-12-14
US (5) US8177703B2 (en)
MACRIS M P ET AL: "IN VIVO EVALUATION OF AN INTRAVENTRICULAR ELECTRIC AXIAL FLOW PUMP FOR LEFT VENTRICULAR ASSISTANCE", ASAIO JOURNAL, LIPPINCOTT WILLIAMS & WILKINS / ASAIO, HAGERSTOWN, MD, US, vol. 40, no. 3, 1 July 1994 (1994-07-01), pages 719 - 722, XP000498262, ISSN: 1058-2916 *
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