Heart pump controller
17 claims: 17 independent, 0 dependent
- 1心臓ポンプ用のコントローラであって、前記コントローラは処理システムを含み、前記処理システムは、 a)少なくとも1つの吸込口および少なくとも1つの吐出口を含むキャビティ内での、インペラであって前記吸込口から前記吐出口へ流体を付勢するためのベーンを含む前記インペラ、の第1軸方向における移動を検出し、 b)前記キャビティ内での前記インペラの軸方向位置を制御するための磁気軸受であって少なくとも1つのコイルを含む前記磁気軸受によって前記第1軸方向と反対の第2軸方向に前記インペラを移動させ、 c)前記磁気軸受によって使用されるパワーを示すインジケータを検出し、 d)前記インジケータに従って前記インペラの軸方向位置を前記磁気軸受によって制御させ、それにより、前記吸込口と前記吐出口との間の流体流量を制御 し、 前記コントローラは、 e)正常均衡位置から離れる前記インペラの軸方向移動を判定し、 f)前記磁気軸受によって前記インペラを前記正常位置に向かって移動させ、 g)前記磁気軸受によって使用されるパワーを監視し、 h)前記磁気軸受によって使用される前記パワーに従って新しい均衡位置を確定し、 i)前記磁気軸受によって前記インペラを前記新しい均衡位置まで移動させるためのものである、 コントローラ。
- 2前記コントローラは、 a)前記インジケータを閾値と比較し、 b)前記比較の結果に応じて、前記第2軸方向への前記インペラの移動を前記磁気軸受によって停止させるためのものである、請求項1に記載のコントローラ。
- 3前記コントローラは、 a)前記インペラの軸方向位置を位置限界と比較し、 b)前記インペラの軸方向位置を前記位置限界内に維持するように、前記磁気軸受を制御するためのものである、請求項1 または2 に記載のコントローラ。
- 4前記コントローラは、 a)前記インペラの軸方向移動を検出することによって前記キャビティの少なくとも一部内の圧力変化を判定し、 b)前記圧力変化に応答して前記インペラの軸方向位置を制御するためのものである、請求項1乃至 3 のいずれか1項に記載のコントローラ。
- 5前記コントローラは、 a)前記キャビティの少なくとも一部内の流体圧力の変化によって引き起こされる前記インペラの移動を検出し、 b)前記インペラの軸方向位置を前記磁気軸受に制御させ、それにより、前記キャビティの少なくとも一部について前記吸込口から前記吐出口への流体流量を変更するためのものである、請求項1乃至 4 のいずれか1項に記載のコントローラ。
- 6前記心臓ポンプは、 a)第1吸込口および第1吐出口を有する第1キャビティ部と、 b)第2吸込口および第2吐出口を有する第2キャビティ部と、 c)前記インペラ上に設けられた第1および第2セットのベーンであって、各セットのベーンは、それぞれの吸込口からそれぞれの吐出口へ流体を付勢するためのものである、第1および第2セットのベーンと、を含む請求項1乃至 5 のいずれか1項に記載のコントローラ。
- 7前記コントローラは、 a)キャビティ部内の相対流体圧力の変化によって引き起こされるインペラの移動を検出し、 b)磁気軸受によってインペラの軸方向位置を制御させ、それにより、前記吸込口から前記吐出口への相対流体流量を変更するためのものである、請求項 6 に記載のコントローラ。
- 8前記正常均衡位置は、各吸込口から各吐出口への必要とされる流体流量を維持するために使用され、 前記新しい均衡位置は前記正常均衡位置からオフセットしており、かつ 前記新しい均衡位置は前記吸込口と前記吐出口との間の相対流体流量を調整するために使用される、請求項 1 に記載のコントローラ。
- 9前記インジケータは、前記磁気軸受によって使用される電流の指標を使用して確定され、かつ 前記コントローラは、前記インペラの軸方向移動をもたらすために、前記磁気軸受によって使用される電流の変化率を確定するためのものである、請求項1乃至 8 のいずれか1項に記載のコントローラ。
- 10前記コントローラは、 a)第1軸方向における前記インペラの移動を判定し、 b)前記第1軸方向と反対の第2軸方向に前記インペラを、 i)前記磁気軸受によって使用される前記パワーが所定量未満に減少するか、または、 ii)前記インペラの軸方向位置が位置限界に達するか、の少なくとも一方になるまで、移動させるように、前記軸受を制御するためのものである、請求項1乃至 9 のいずれか1項に記載のコントローラ。
- 11前記心臓ポンプは磁気ドライブを含み、 前記ドライブは、前記インペラ内の第1磁気材料と協働する磁界を使用時に発生し、前記インペラが回転することを可能にする少なくとも1つのドライブコイルを含み、かつ 前記コントローラは、前記ドライブを制御し、それにより、回転軸の回りの前記インペラの回転を引き起こすためのものである、請求項1乃至 10 のいずれか1項に記載のコントローラ。
- 12使用時に、少なくとも1つの軸受コイルは、前記インペラ内の第2磁気材料と協働する磁界を発生し、前記インペラの前記軸方向位置を制御可能にする、請求項1乃至 11 のいずれか1項に記載のコントローラ。
- 13前記心臓ポンプはハウジングを含み、 前記インペラは、 a)前記インペラの第1端部に設けられた第1磁気材料であって、前記インペラの回転を可能にするために前記ドライブと協働するための第1磁気材料と、 b)前記インペラの前記第1端部と相対向する第2端部に設けられた第2磁気材料であって、前記インペラの前記軸方向位置を制御可能にするために、前記磁気軸受と協働するための第2磁気材料と、を含み、 i)前記ドライブは前記ハウジングの第1端部に位置決めされ、前記ドライブおよび前記インペラは、前記ドライブと前記インペラとの間に第1吸引力をもたらすように構成され、かつ ii)前記磁気軸受は前記ハウジングの第2端部に位置決めされ、前記磁気軸受および前記インペラは、前記磁気軸受と前記インペラとの間に第2吸引力をもたらすように構成され、前記第1および第2吸引力は、前記インペラが、正常循環状態中に前記キャビティ内の軸方向中心位置に位置決めされるときに、均衡する、請求項 11 又は 12 に記載のコントローラ。
- 14前記処理システムは、 a)命令を記憶するためのメモリと、 b)プロセッサと、を含み、前記プロセッサは、前記命令を実行し、それにより、 i)前記第1軸方向における前記インペラの移動を判定し、 ii)前記第2軸方向への前記インペラの移動を前記磁気軸受にさせるための信号を生成し、 iii)前記磁気軸受によって使用されるパワーを示すインジケータを確定し、 iv)前記インジケータに従って前記インペラの軸方向位置を前記磁気軸受に制御させ、それにより、前記吸込口と前記吐出口との間の流体流量を制御するための信号を生成することを、前記プロセッサにさせる、請求項1乃至 13 のいずれか1項に記載のコントローラ。
- 15心臓ポンプであって、 a)少なくとも1つの吸込口および少なくとも1つの吐出口を含むキャビティと、 b)前記キャビティ内に設けられるインペラであって、前記吸込口から前記吐出口へ流体を付勢するためのベーンを含む、インペラと、 c)前記キャビティ内で前記インペラを回転させるためのドライブと、 d)前記キャビティ内で前記インペラの軸方向位置を制御するための少なくとも1つの軸受コイルを含む磁気軸受と、 e)コントローラと、を含み、前記コントローラは、 i)第1軸方向における前記インペラの移動を検出し、 ii)前記磁気軸受によって前記第1軸方向と反対の第2軸方向に前記インペラを移動させ、 iii)前記磁気軸受によって使用されるパワーを示すインジケータを検出し、 iv)前記インジケータに従って前記インペラの軸方向位置を前記磁気軸受にて制御させ、それにより、前記吸込口と前記吐出口との間の流体流量を制御 し、 前記コントローラは、 f)正常均衡位置から離れる前記インペラの軸方向移動を判定し、 g)前記磁気軸受によって前記インペラを前記正常位置に向かって移動させ、 h)前記磁気軸受によって使用されるパワーを監視し、 i)前記磁気軸受によって使用される前記パワーに従って新しい均衡位置を確定し、 j)前記磁気軸受によって前記インペラを前記新しい均衡位置まで移動させるためのものである 、心臓ポンプ。
- 16前記心臓ポンプは、 a)第1吸込口および第1吐出口を有する第1キャビティ部と、 b)第2吸込口および第2吐出口を有する第2キャビティ部と、 c)前記インペラ上に設けられた第1および第2セットのベーンであって、各セットのベーンは、それぞれの吸込口からそれぞれの吐出口へ流体を付勢するためのものである、第1および第2セットのベーンと、を含む、請求項 15 に記載の心臓ポンプ。
- 17心臓ポンプを制御する方法であって、コントローラ内に、 a)少なくとも1つの吸込口および少なくとも1つの吐出口を含むキャビティ内での、インペラであって前記吸込口から前記吐出口へ流体を付勢するためのベーンを含むインペラ、の第1軸方向における移動を検出すること、 b)前記キャビティ内での前記インペラの軸方向位置を制御するための磁気軸受であって少なくとも1つのコイルを含む磁気軸受によって前記第1軸方向と反対の第2軸方向に前記インペラを移動させること、 c)前記磁気軸受によって使用されるパワーを示すインジケータを検出すること、および、 d)前記インジケータに従って前記インペラの軸方向位置を前記磁気軸受によって制御させ、それにより、前記吸込口と前記吐出口との間の流体流量を制御すること、を含 み、 前記コントローラは、 e)正常均衡位置から離れる前記インペラの軸方向移動を判定し、 f)前記磁気軸受によって前記インペラを前記正常位置に向かって移動させ、 g)前記磁気軸受によって使用されるパワーを監視し、 h)前記磁気軸受によって使用される前記パワーに従って新しい均衡位置を確定し、 i)前記磁気軸受によって前記インペラを前記新しい均衡位置まで移動させるためのものである 方法。
Independent claims17
158 paragraphs, as filed
0001The present invention relates to a controller for a heart pump and a method of controlling the heart pump.
0002References herein to any conventional publication (or information drawn from that publication) or any known matter are conventional publications (or information drawn from that publication) or known matter. However, it cannot and should not be considered as an endorsement, permission, or any form of suggestion that this specification forms part of the common general knowledge in the relevant areas of effort.
0003The use of mechanical device therapy to treat heart failure is increasing with the aging of the population, and the number of donor organs for heart transplantation remains limited. The device can be used to bridge me between the patient and a heart transplant, to bridge me between the patient and recovery, or in fact as a permanent alternative. The latter support strategy requires a device with increased mechanical durability / life.
0004Mechanical durability depends on the function of the device, especially the type of bearing mounted. First-generation pulsating devices require contact components, limiting the expected mechanical life of pulsating devices to less than three years. The reduction in size of second-generation non-pulsating rotating impeller devices has put the device at the forefront of VAD development.
0005However, early techniques for impeller support also imposed considerable limitations on device life due to the need for shafts, seals, and bearings (Patent Document 1). Subsequent improvements have resulted in devices that rely on pivot support immersed in blood (Patent Document 2). However, the expected usable life is still less than 5 years.
0006Since then, several techniques have been developed to improve device life, ranging from full magnetic levitation (Patent Document 3) to passive fluid levitation (Patent Document 4). These 3rd generation devices eliminate contact wear, reduce the number of moving parts, and increase life, perhaps up to over 10 years. These latest generation levitation techniques eliminate point-to-point contact, which can also improve the hemolytic performance of the pump.
0007Some commercial and research devices implement hydrodynamic or magnetic bearing technology. For example, Ventrassist (Ventracor, Sydney, New South Wales, Australia (Sydney, NSW, AU)) and HVAD (Heatware, Sydney, New South Wales, Australia (Sydney, NSW, AU)) are wholly or partially by fluid force. It incorporates an impeller that is surfaced and driven by an electromagnetic motor (Patent Documents 4 and 5 respectively). Duraheart (Terumo, Ann Arbor, MI, US, USA) uses an axial magnetic bearing with a permanent magnet coupling motor (Patent Document 3). Heartmate III (Thoratec, Woburn, MA, USA) uses a combined radial self-bearing motor (Patent Document 6), while Levacor (Worldheart, Ottawa, ON, Canada, Ontario, Canada). )) Uses an axial magnetic bearing and an electromagnetically coupled motor (Patent Document 7).
0008All of the above devices provide left ventricular assist (LVAD). However, a significant number of patients also require a device for right ventricular assistance (RVAD). The incidence of biventricular failure is not always apparent from the beginning in patients with heart failure, and right ventricular heart failure can occur in up to 40% of patients receiving LVAD assistance.
0009One of the most effective BiVAD techniques used in clinical practice uses an in vitro connection of two Thoratec PVA devices. Smaller 2nd and 3rd generation rotating systems have also been proposed, two separate, such as the combined Coraide and Dexaide (Patent Document 8) and two gyro pumps (Patent Document 2). Use a rotary pump.
0010However, all bichamber auxiliary systems currently available require the use of two devices with separate controllers, which can lead to left and right discharge control problems. The dual device approach also increases implantation size as well as treatment costs.
0011Single-rotation pumps have also been designed to enhance the function of both chambers of the incomplete heart, as described in Patent Documents 9-12. Each of these devices includes a double-sided impeller that rotates at a common speed, with each side impeller configured for left and right heart support, respectively. This introduces the inherent problem of the ability to independently control and thus balance the discharge rates from the left and right sides of the device. That is, an increase in impeller speed produces a corresponding increase in discharge from both cavities.
0012Patent Document 12 addresses this problem by introducing the ability to axially displace the rotating impeller in the cavity to simultaneously change the relative efficiencies on both sides of the device. However, the present application describes the control method used to achieve this axial displacement as active, thus using the feedback signal from the pressure sensor and actively locating the desired set axial location. We need a similar method to control and maintain.
0013The ability of the BiVAD system to maintain a balance between left and right discharges is essential for effective device operation. Hemodynamic parameters that can upset this balance are bronchial flow, relative changes in systemic and pulmonary vascular resistance, relative changes in left and right ventricular contractility, pulmonary or systemic congestion, and ventricular collapse. including. It makes us speculate that techniques for balancing these conditions, left and right VAD hydraulic outputs, are needed for long-term support.
0014Parameters such as motor power, velocity, differential pressure (suction port-discharge port), and discharge rate are required to operate and control the hydraulic output from each blood pump. Although determining motor power and speed is relatively easy, traditionally additional equipment such as pressure sensors and flow meters have been required to detect the remaining parameters. These components increase the likelihood of device failure. Therefore, the long-term reliability of the parts is limited. In addition, the addition of components to the device induces extra blood contact with other foreign substances, exacerbating the potential for blood damage.
0015Previous attempts to regulate discharge from each device and balance left / right discharge requirements often rely on the use of pressure sensors to detect left atrial pressure (LAP). It was. A feedback mechanism is then used to reduce the LVAD rate or increase the RVAD rate in the presence of reduced LAP. Another technique involves surgically introducing a shunt between the left and right atrium to safely protect it from the accumulation of possibly catastrophic fluid in either atrium. Alternatively, Patent Document 13 includes, via a variable occlusion valve, a conduit connecting the right atrium to the left atrium whose flow rate is altered. However, this technique introduces the complexity associated with active feedback control, such as additional blood contact conduits and the need for sensors. Moreover, while this solution can help balance the fluid distribution, it does not provide a way to control changes in device discharge.
0016As already mentioned, the ability to alter the left and right discharges of BiVAD, especially during the postoperative period when the neurohumoral self-regulatory mechanism is at least partially ablated by anesthetics and serious illness. is important.
0017Many control algorithms exist for active flotation of magnetic bearing systems. Most focus on maintaining the centering of the rotor position, but there are alternative techniques that focus on minimizing power consumption. The latter controller uses the passively generated force from the permanent magnets in the magnetic circuit to resist external forces that would otherwise require power from the electromagnetic coil. This results in the movement of the impeller from the centered position until the equilibrium between the external force and the permanent magnet force is reached. Therefore, the power consumption of the active electromagnetic coil is returned to the minimum state. Numerous rotating blood pump designs implement this form of zero power control.
0018Masuzawa et al. (2004) implemented zero power control in a reluctance type radial magnetic motor bearing to completely levitate the rotor of a centrifugal blood pump. The system includes permanent magnets that place magnetic materials concentrically around the rotor to provide additional bias flux to the magnetic circuit. These magnets are used by zero power controllers to reduce power consumption when compared to center position controllers. During operation, the application of radial hydraulic pressure to the rotor causes translation of this rotor in the direction perpendicular to the axis of rotation and opposite the applied force. However, no significant effect on pump discharge is observed with respect to this movement. The reason is that the alternating radial gaps have the least effect on hydraulic efficiency. (Non-Patent Document 1).
0019Patent Document 14 axially levitates the rotor of a centrifugal blood pump by means of a Lorentz type magnetic bearing system. The system also places magnetic materials concentrically around the rotor. However, the magnetic force acts perpendicular to the polar plane in the direction parallel to the axis of rotation. Additional permanent magnets not included in the magnetic bearing circuit are configured to provide offsetting forces when they encounter axial hydraulic pressure. This offsetting force is effective when it allows the zero power controller to translate the impeller in the same direction as the applied force. This movement may be adapted to change the discharge rate of the device, but movement in the same direction as the applied force is undesirable and increases the discharge rate, for example when a reduction is guaranteed. It will be.
0020Patent Document 15 uses Lorentz type axial magnetic bearings arranged concentrically around the rotor. Radial levitation is achieved using the configuration of repulsive permanent magnets (Patent Documents 16 and 17) composed of a Halbach array (Patent Document 15). These magnets are used to achieve zero power control that rearranges the axial position of the impeller in response to hydraulic axial forces. Since this configuration uses repulsive magnets to achieve this, the low stiffness of the repulsive magnets may not provide sufficient force offset for a given displacement. Axial rearrangement is the opposite of the direction of the applied force, but the shrouded configuration and the placement of the impeller vane under the impeller reduce the pump discharge in response to the force generated at the moment of ventricular deviation. Not realized. Therefore, the zero power controller is a bearing. bower) can be minimized, but does not provide discharge control based on changing preload conditions. In addition, the aforementioned axial gap (0.005 inch) between the bottom impeller shroud and the casing is too small, and the impeller blade height is too high, even if the shroud is half open, the largest shaft. It is not possible to generate a clear change in hydraulic performance with directional translation enabled.
0021Patent Document 18 describes a pump including a housing, a stator supported within the housing, and a rotor assembly. The rotor assembly includes a rotor supported in the housing for rotation with respect to the stator around the shaft. The rotor assembly also includes a first impeller operably coupled to the first axial end of the rotor for rotation by the rotor around the shaft. The rotor assembly further includes a second impeller operably coupled to the second axial end of the rotor facing the first axial end for rotation by the rotor around the axis. The rotor assembly is movable along the axis with respect to the housing to adjust the hydraulic performance characteristics of the pump. This axial movement depends on the difference in inflow pressure between the left and right atrium, thus resulting in a fully passive movement of the rotor in response to the force generated. In this case, the force generated by such a pressure difference moves the rotor in the direction of the applied force, thus requiring the use of an open impeller and a gap below the impeller, which in impeller efficiency. Effective changes, thus reducing discharge due to axial translation. A further embodiment describes a solenoid-type mechanism for axially displacing the impeller. The results obtained by this mechanism allow the selection of only the complete right or complete left operating point of the rotating impeller assembly, but not the axial position between the two points.
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<p num="0023"><nplcit num="1"><text>"Zero Power Control for Magnetically Suspended Artificial Heart" by Masuzawa, T., H. Omura, and Y. Okada Proceedings of the Joint Lecture on Automatic Control (Jido Seigjo Rengo Koenkai Koen Ronbunsyu), 2004, 47: p.322</text></nplcit></p>
<p num="0024"> It is an object of the present invention to substantially overcome, or at least improve, one or more of the shortcomings of existing configuration configurations.</p>
<p num="0025"> In the first broad form, the invention seeks to provide a controller for a cardiac pump, the controller comprising a processing system, the processing system. a) First axial movement of the impeller, which is the impeller and contains the vanes to urge the fluid from the inlet to the outlet, within the cavity containing at least one inlet and at least one outlet. Judge, b) A magnetic bearing for controlling the axial position of the impeller in the cavity, which moves the impeller in the second axial direction opposite to the first axial direction by a magnetic bearing containing at least one coil. c) Determine the indicator of the power used by the magnetic bearings d) The purpose is to control the axial position of the impeller according to the indicator by a magnetic bearing, thereby controlling the fluid flow rate between the suction port and the discharge port.</p><p num="0026"> Typically, the controller a) Compare the indicator to the threshold b) The purpose is to stop the movement of the impeller in the second axial direction by a magnetic bearing according to the result of comparison.</p><p num="0027"> Typically, the controller is for minimizing the power used by the magnetic bearings. Typically, the controller a) Compare the axial position of the impeller with the position limit and b) The purpose is to control the magnetic bearing so that the axial position of the impeller is maintained within the position limit.</p><p num="0028"> Typically, the controller a) Determine pressure changes in at least part of the cavity b) This is to control the axial position of the impeller in response to pressure changes.</p><p num="0029"> Typically, the controller is for determining the pressure change by detecting the axial movement of the impeller. Typically, the axial position determines the separation distance between the vane and the cavity surface, and the separation distance is used to control the fluid flow rate from the inlet to the outlet.</p><p num="0030"> Typically, the controller a) Detect impeller movement caused by changes in fluid pressure within at least part of the cavity b) The purpose is to have the magnetic bearing control the axial position of the impeller, thereby changing the fluid flow rate from the suction port to the discharge port for at least a part of the cavity.</p><p num="0031"> Typically, the controller a) The separation distance between the vane and the cavity surface is reduced by magnetic bearings, thereby increasing the fluid flow rate from the suction port to the discharge port. b) The separation distance between the vane and the cavity surface is increased by magnetic bearings, thereby reducing the fluid flow rate from the suction port to the discharge port. For at least one of them.</p><p num="0032"> Typically, the controller a) Determine the axial movement of the impeller away from the normal equilibrium position b) Move the impeller toward the normal position by magnetic bearings c) Monitor the power used by magnetic bearings and d) Establish a new equilibrium position according to the power used by the magnetic bearings e) The magnetic bearing is used to move the impeller to a new equilibrium position.</p><p num="0033"> The equilibrium position is typically used to maintain the required fluid flow rate from the inlet to the outlet. Typically, a heart pump a) A first cavity having a first suction port and a first discharge port, b) A second cavity with a second suction port and a second discharge port, c) First and second sets of vanes provided on the impeller, each set of vanes for urging fluid from their respective inlets to their respective outlets. Includes a second set of vanes.</p><p num="0034"> Typically, the axial position determines the separation distance between each set of vanes and the respective cavity surface, and the separation distance is used to control the fluid flow rate from the inlet to the outlet.</p><p num="0035"> Typically, the controller a) Detects impeller movement caused by changes in relative fluid pressure in the cavity and b) The purpose is to control the axial position of the impeller by a magnetic bearing, thereby changing the relative fluid flow rate from the suction port to the discharge port.</p><p num="0036"> Typically, the controller a) Determine the axial movement of the impeller away from the normal equilibrium position b) Move the impeller toward the normal position by magnetic bearings c) Monitor the power used by magnetic bearings and d) Establish a new equilibrium position according to the power used by the magnetic bearings e) The magnetic bearing is used to move the impeller to a new equilibrium position.</p><p num="0037"> Typically, the normal equilibrium position is used to maintain the required fluid flow rate from each inlet to each outlet. Typically, the new equilibrium position is offset from the normal equilibrium position.</p><p num="0038"> Typically, the new equilibrium position is used to regulate the relative fluid flow rate between the inlet and outlet. Typically, the indicator is determined using an indicator of the current used by the magnetic bearing.</p><p num="0039"> Typically, the controller is for determining the rate of change of the current used by the magnetic bearings to result in the axial movement of the impeller. Typically, the controller a) Judge the movement of the impeller in the 1st axis direction b) Place the impeller in the 2nd axis direction opposite to the 1st axis direction. i) The power used by magnetic bearings is reduced to less than a given amount or ii) The purpose is to control the bearing so that it is moved until the axial position of the impeller reaches the position limit or at least one of them.</p><p num="0040"> Typically, the heart pump includes a magnetic drive, the controller is for controlling the drive, thereby causing the impeller to rotate around the axis of rotation. Typically, the drive includes at least one drive coil that generates a magnetic field in use that cooperates with the magnetic material in the impeller and allows the impeller to rotate.</p><p num="0041"> Typically, the drive is positioned at the first end of the housing and the drive and impeller are configured to provide suction between the drive and the impeller. Typically, during use, at least one coil creates a magnetic field that collaborates with the magnetic material in the impeller, allowing the axial position of the impeller to be controlled.</p><p num="0042"> Typically, the magnetic bearing is positioned at the second end of the housing and the magnetic bearing and impeller are configured to provide attractive force between the magnetic bearing and the impeller. Typically, the impeller is a) With the first magnetic material to work with the drive to allow the impeller to rotate, b) Includes a second magnetic material for working with magnetic bearings to allow control of the axial position of the impeller.</p><p num="0043"> Typically a) The first magnetic material is provided on the first end of the impeller, b) The second magnetic material is provided at the second end of the impeller opposite the first end.</p><p num="0044"> Typically a) The drive is positioned at the first end of the housing and the drive and impeller are configured to provide a first suction force between the drive and the impeller. b) The magnetic bearing is positioned at the second end of the housing, the magnetic bearing and impeller are configured to provide a second attractive force between the magnetic bearing and the impeller, and the first and second attractive forces are the impeller. Are approximately balanced when positioned approximately axially centered within the cavity during normal circulation.</p><p num="0045"> In a second broad form, the invention aims to provide a method of controlling a cardiac pump, the method being in a controller. a) First axial movement of the impeller, which is the impeller and contains the vane to urge the fluid from the suction port to the discharge port, within the cavity containing at least one suction port and at least one discharge port. Judging, b) Moving the impeller in the second axial direction opposite to the first axial direction by a magnetic bearing that controls the axial position of the impeller in the cavity and contains at least one coil. c) Determine the indicator of the power used by the magnetic bearing, and d) Including controlling the axial position of the impeller according to the indicator by a magnetic bearing, thereby controlling the fluid flow rate between the suction port and the discharge port.</p><p num="0046"> In a third broad form, the present invention aims to provide a controller for a cardiac pump, the controller comprising a processing system, the processing system. a) Movement of the impeller from an equilibrium position within the cavity containing at least one suction port and at least one discharge port, including a vane for urging fluid from the suction port to the discharge port. Judge, b) A new balanced position, suction, based on the index of power used by the magnetic bearing, which is the magnetic bearing and contains at least one coil to control the axial position of the impeller in the cavity. It is used to move the impeller to a new equilibrium position by magnetic bearings, which is used to control the fluid flow rate from the port to the discharge port.</p><p num="0047"> In a fourth broad embodiment, the present invention aims to provide a controller for a heart pump, the controller comprising a processing system for controlling the axial position of the impeller within the cavity, the cavity being the first. Includes a first cavity with a suction port and a first discharge port and a second cavity with a second suction port and a second discharge port, the impeller contains the vanes of the first and second sets, and the vanes of each set. Is for urging fluid from each inlet to each outlet, and the controller will move the impeller into the first cavity when the relative pressure in the first cavity increases relative to the second cavity. Positioned, thereby controlling the axial position to increase the relative fluid flow rate from the first outlet to the second outlet.</p><p num="0048"> In a fifth broad embodiment, the present invention aims to provide a controller for a heart pump, the controller including a processing system for controlling the axial position of the impeller within the cavity, the cavity being a suction port and Including the outlet, the impeller includes a vane to urge the fluid from the inlet to the outlet, and the controller moves the impeller away from the inlet as the pressure in the cavity increases. , Control the axial position to reduce the discharge port fluid pressure.</p><p num="0049"> A cardiac pump comprising a controller according to any one of the broad forms of the invention. In a sixth broad form, the present invention aims to provide a heart pump, which is a heart pump. a) With a cavity containing at least one inlet and at least one outlet, b) An impeller provided in the cavity, including a vane for urging the fluid from the suction port to the discharge port. c) A drive to rotate the impeller in the cavity, d) With a magnetic bearing containing at least one bearing coil to control the axial position of the impeller in the cavity, e) Controller and Including, the controller i) Determine the movement of the impeller in the 1st axis direction ii) The impeller is moved in the 2nd axial direction opposite to the 1st axial direction by the magnetic bearing. iii) Determine the indicator indicating the power used by the magnetic bearing iv) The purpose is to control the axial position of the impeller according to the indicator with a magnetic bearing, thereby controlling the fluid flow rate between the suction port and the discharge port.</p><p num="0050"> Typically, a heart pump a) A first cavity having a first suction port and a first discharge port, b) A second cavity with a second suction port and a second discharge port, c) First and second sets of vanes provided on the impeller, each set of vanes for urging fluid from their respective inlets to their respective outlets. Includes a second set of vanes.</p><p num="0051"> Typically, the axial position determines the separation distance between each set of vanes and the respective cavity surface, and the separation distance is used to control the fluid flow rate from the inlet to the outlet.</p><p num="0052"> Typically, the drive a) The first magnetic material provided in the impeller and b) Includes at least one drive coil, which generates a magnetic field in use that cooperates with the magnetic material in the impeller and allows the impeller to rotate.</p><p num="0053"> Typically, the first magnetic material comprises a number of circumferentially spaced permanent magnets mounted within the impeller, with adjacent magnets having opposite polarities. Typically, the drive is positioned at the first end of the cavity and the drive and impeller are configured to provide suction between the drive and the impeller.</p><p num="0054"> Typically, during use, at least one bearing coil creates a magnetic field that collaborates with a second magnetic material in the impeller, allowing the impeller's axial position to be controlled. Typically, the bearing magnetic material is a ferromagnetic core.</p><p num="0055"> Typically, the bearing is a permanent magnet and includes a permanent magnet to provide attractive force between the permanent magnet and the bearing magnetic material. Typically, the coil is for generating a magnetic field that is either complementary or vice versa to the magnetic field generated by the permanent magnet, thereby bearing and bearing. Controls the net magnetic field to and from the magnetic material.</p><p num="0056"> Typically, the magnetic bearing is positioned at the second end of the cavity and the magnetic bearing and impeller are configured to provide an attractive force between the magnetic bearing and the impeller. Typically, the impeller is a) With the first magnetic material to work with the drive to allow the impeller to rotate, b) Includes a second magnetic material for working with magnetic bearings to allow control of the axial position of the impeller.</p><p num="0057"> Typically a) The first magnetic material is provided on the first end of the impeller, b) The second magnetic material is provided at the second end of the impeller facing the first end.</p><p num="0058"> Typically a) The drive is positioned at the first end of the housing and the drive and impeller are configured to provide a first suction force between the drive and the impeller. b) The magnetic bearing is positioned at the second end of the housing, the magnetic bearing and impeller are configured to provide a second attractive force between the magnetic bearing and the impeller, and the first and second attractive forces are the impeller. Are approximately balanced when positioned approximately axially centered within the cavity during normal circulation.</p><p num="0059"> It is understood that the broad form of the invention may be used individually or in combination.</p>
0060<figref num="1A">It is the schematic sectional drawing of the 1st Example of a heart pump.</figref><figref num="1B">It is the schematic sectional drawing of the 2nd Example of a heart pump.</figref><figref num="2">It is a flowchart of one Example of the method of controlling the position of an impeller in a heart pump.</figref><figref num="3">It is the schematic of one Example of a controller.</figref><figref num="4">It is a flowchart of the 2nd Example of the method for controlling the position of an impeller in a heart pump.</figref><figref num="5">FIG. 6 is a schematic exploded perspective view of an embodiment of a drive and magnetic bearing system for a heart pump.</figref><figref num="6">FIG. 5 is a schematic side view of the drive and magnetic bearing system of FIG.</figref><figref num="7">It is a schematic plan view of the drive system of FIG.</figref><figref num="8">It is a schematic plan view of the magnetic bearing system of FIG.</figref><figref num="9">FIG. 5 is a schematic plan view of the rotor of the drive system of FIG.</figref><figref num="10">It is a schematic plan view of the rotor of the magnetic bearing system of FIG.</figref><figref num="11A">It is a schematic side view which shows the operation principle of the axial position control by a magnetic bearing.</figref><figref num="11B">It is a schematic side view which shows the operation principle of the tilt control by a magnetic bearing.</figref><figref num="12A">It is a schematic side view which shows the method of rearranging the impeller position by zero power control.</figref><figref num="12B">It is a schematic side view which shows the method of rearranging the impeller position by zero power control.</figref><figref num="12C">It is a schematic side view which shows the method of rearranging the impeller position by zero power control.</figref><figref num="13A">It is a schematic side view which shows the method which the single-sided impeller position is rearranged in order to affect the device discharge amount.</figref><figref num="13B">It is a schematic side view which shows the method which the single-sided impeller position is rearranged in order to affect the device discharge amount.</figref><figref num="13C">It is a schematic side view which shows the method which the single-sided impeller position is rearranged in order to affect the device discharge amount.</figref><figref num="14A">FIG. 5 is a schematic side view showing a method of rearranging the double-sided impeller positions in order to affect the device discharge rate.</figref><figref num="14B">FIG. 5 is a schematic side view showing a method of rearranging the double-sided impeller positions in order to affect the device discharge rate.</figref><figref num="14C">FIG. 5 is a schematic side view showing a method of rearranging the double-sided impeller positions in order to affect the device discharge rate.</figref><figref num="15">It is a graph which shows the relative hydraulic performance of the left discharge amount and the right discharge amount in an exemplary BiVAD.</figref><figref num="16A">It is a schematic diagram of one Example of BiVAD.</figref><figref num="16B">It is a schematic diagram of one Example of BiVAD.</figref><figref num="16C">It is a schematic diagram of one Example of BiVAD.</figref><figref num="16D">It is a schematic diagram of one Example of BiVAD.</figref><figref num="16E">It is a schematic diagram of one Example of BiVAD.</figref><figref num="16F">It is a schematic diagram of one Example of BiVAD.</figref><figref num="16G">It is a schematic diagram of one Example of BiVAD.</figref><figref num="16H">It is a schematic diagram of one Example of BiVAD.</figref><figref num="17A">It is a schematic diagram which shows the axial force generation obtained on the impeller of BiVAD under various general conditions.</figref><figref num="17B">FIG. 5 is a schematic diagram showing the resulting axial force generation on the BiVAD impeller under various general conditions.</figref><figref num="17C">FIG. 5 is a schematic diagram showing the resulting axial force generation on the BiVAD impeller under various general conditions.</figref><figref num="17D">FIG. 5 is a schematic diagram showing the resulting axial force generation on the BiVAD impeller under various general conditions.</figref><figref num="18A">It is a schematic diagram showing the adaptation of BiVAD to the change of relative vascular resistance.</figref><figref num="18B">It is a schematic diagram showing the adaptation of BiVAD to the change of relative vascular resistance.</figref><figref num="18C">It is a schematic diagram showing the adaptation of BiVAD to the change of relative vascular resistance.</figref><figref num="19A">It is a schematic diagram which shows the indication of BiVAD for the heart chamber deviation.</figref><figref num="19B">It is a schematic diagram which shows the indication of BiVAD for the heart chamber deviation.</figref><figref num="19C">It is a schematic diagram which shows the indication of BiVAD for the heart chamber deviation.</figref><figref num="20A">FIG. 6 is a schematic diagram showing the resulting axial force generation on the VAD impeller under various general conditions.</figref><figref num="20B">FIG. 6 is a schematic diagram showing the resulting axial force generation on the VAD impeller under various general conditions.</figref><figref num="20C">FIG. 6 is a schematic diagram showing the resulting axial force generation on the VAD impeller under various general conditions.</figref><figref num="20D">FIG. 6 is a schematic diagram showing the resulting axial force generation on the VAD impeller under various general conditions.</figref><figref num="21A">It is a schematic diagram which shows the adaptation of VAD to the change of relative vascular resistance.</figref><figref num="21B">It is a schematic diagram which shows the adaptation of VAD to the change of relative vascular resistance.</figref><figref num="21C">It is a schematic diagram which shows the adaptation of VAD to the change of relative vascular resistance.</figref><figref num="22A">It is a schematic diagram which shows the adaptation of VAD to the heart chamber deviation.</figref><figref num="22B">It is a schematic diagram which shows the adaptation of VAD to the heart chamber deviation.</figref><figref num="22C">It is a schematic diagram which shows the adaptation of VAD to the heart chamber deviation.</figref><figref num="23A">FIG. 3 is a graph of exemplary left and right atrial pressure (LAP, RAP) for an exemplary cardiac pump within a circulatory loop that simulates several circulatory states.</figref><figref num="23B">FIG. 6 is a graph of exemplary systemic and pulmonary resistance (SVR, PVR) for an exemplary cardiac pump within a circulatory loop that simulates several circulatory states.</figref><figref num="23C">FIG. 6 is a graph of exemplary impeller axial positions for an exemplary cardiac pump within a circulatory loop that simulates several circulatory states.</figref><figref num="23D">FIG. 6 is a graph of the use of exemplary magnetic bearing power (MB power) for an exemplary cardiac pump in a circulating loop that simulates several circulating conditions.</figref><figref num="23E">FIG. 6 is a graph of exemplary aortic and pulmonary pressures (AoP, PAP) for an exemplary cardiac pump in a circulatory loop that simulates several circulatory states.</figref>
0061Examples of the present invention are described herein with reference to the accompanying drawings. A first embodiment of a cardiac pump is described below with reference to FIG. 1A. In this embodiment, the cardiac pump 100A defines a cavity 120 and includes a housing 110 that includes an impeller 130A. The impeller 130A effectively divides the cavity 120 into first and second cavities 121, 122A. The housing 110 includes first and second suction ports 141 and 142 and corresponding first and second outlets 151 and 152 for fluid communication with the first and second cavities 121 and 122A, respectively.
0062The impeller 130A includes first and second sets of vanes 131, 132 such that the rotation of the impeller 130A around the axis 160 urges the fluid from the inlets 141, 142 to the corresponding outlets 151, 152. .. In use, rotation of the impeller 130A is achieved using a drive such as the magnetic drive 170. The magnetic drive 170 typically includes at least one coil positioned adjacent to the first cavity 121 at the first end of the housing 110. In use, the coil creates a magnetic field that collaborates with the magnetic material in the impeller 130A, allowing the impeller to rotate. This tends to provide a suction force between the drive 170 and the impeller 130A that urges the impeller 130A axially towards the first cavity 121.
0063In use, the relative physical separation distance between the vanes 131, 132 of the set and the corresponding cavity surfaces 123, 124A is the relative efficiency of the vanes 131, 132, and therefore the suction ports 141, 142 and the corresponding discharge ports 151, Control the relative flow rate to and from 152. The position of the axial impeller 130A is typically controlled using a magnetic bearing 180. The magnetic bearing 180 typically includes at least one coil positioned at the second end of the housing 110 adjacent to the second cavity 122A. In use, the coil creates a magnetic field that interacts with the ferrous material in the impeller 130A, as well as in cooperation with the magnetic material in the magnetic bearing stator, allowing control of the axial position of the impeller 130A. This tends to provide an attractive force between the magnetic bearing 180 and the impeller 130A that urges the impeller 130A axially towards the second cavity 122A.
0064The drive 170 and bearing 180 are typically coupled to the controller 190 to control the operation of the heart pump 100. The controller is also typically coupled to a sensor, the embodiment of which is described in more detail below, which allows the position of the impeller to be determined.
0065A second embodiment of a cardiac pump is described herein with reference to FIG. 1B. In this embodiment, similar codes are used to specify similar features, which are therefore not described in detail.
0066In this embodiment, the cardiac pump 100B includes a second modified cavity 122B having a surface 124B extending over the housing 110, and a suction port is provided in the embodiment of FIG. 1A. Therefore, in this embodiment, the heart pump 110 does not include a second inlet or a second outlet. Further, the impeller 130B contains only a single set of vanes 131 positioned within the cavity 121 and extends through the impeller 130B to allow blood to flow from the second cavity 122B to the first cavity 121. Includes the existing aperture 135, thereby preventing stagnation between the impeller 130B and the second cavity surface 124B.
0067Upon use, cardiac pumps 100A, 100B may be coupled to the subject to supplement the pumping action of one or both of the left and right ventricles of the heart. For example, the cardiac pump 100A in Figure 1A can be coupled to both the lung and systemic circulatory systems, allowing the pump to operate as a BiVAD (Bi-Ventricular assist Device), which allows the pump to operate. , Supplements the pumping action of both the left and right ventricles of the heart. In this example, the left ventricle and right atrium connect to the first and second suction ports 141, 142, respectively, while the first and second outlets 151, 152 supply the discharge flow to the aorta and pulmonary artery.
0068In use, the cardiac pump 100A provides the pumping action provided by vanes 131, 132 of each set, with the impeller 130A positioned approximately axially centered within the cavity 120, commonly referred to as the nominal equilibrium point, in the left ventricle. And are arranged to be equal to the pumping action required by each of the right ventricles. This can be achieved by choosing suitable dimensions such as the length, height, and shape of each vane, and generally achieves a flow rate of about 5 L / min at each outlet 151, 152.
0069When the circulatory system is functioning properly, the pressure in the first cavity 121 is about 100 mmHg as opposed to 20 mmHg, which will be greater than the pressure in the second cavity 122A. In this example, blood flow between the first cavity 121 and the second cavity 122A is substantially blocked due to the presence of the impeller 130A. Depending on the geometry of the impeller, this normal pressure difference can provide the impeller 130A with a force acting towards, for example, the second cavity 122A.
0070In one embodiment, the cardiac pump exerts on the impeller 130A, including the forces resulting from the pressure difference and the attractive forces caused by the magnetic coupling between the impeller 130A and the drive 170 and between the impeller 130A and the bearing 180. These forces naturally balance to be about the same when the Impeller 130A is brought to the equilibrium point.
0071As described in more detail below, the position of the equilibrium point within the cavity 120 is controlled and is usually positioned in the axial center of the cavity 120 when the circulatory system is functioning properly. This can be achieved by selecting suitable magnetic properties for the impeller 130A, drive 170, and bearing 180. In these situations, the additional force required by the magnetic bearing 180 to maintain the impeller 130A at the equilibrium point is minimal, which is sometimes referred to as the "zero power configuration". In this regard, the term zero is not necessarily understood to be zero, but rather means that the power required is less than in the absence of such an equilibrium.
0072In this example, if there is a change in the relative pressure of the first and second cavities 121, 122A caused by an increase in pressure in the pulmonary circulation and / or a decrease in pressure in the systemic circulation, this is corrected. It will result in a pressure difference. The modified pressure difference provides a net force on the impeller 130A, causing the impeller 130A to move away from the equilibrium point towards the first cavity 121, thereby between the vanes 131 of the first set and the surface 123 of the first cavity. Reduce the separation distance between. This increases the efficiency of the pumping effect in the systemic system and decreases the efficiency of the pumping effect in the lung system. This will increase the flow into the systemic circulation while decreasing the flow into the pulmonary circulation, exacerbating the flow balance problem in the lungs and / or the systemic circulation.
0073Similarly, a decrease in pressure within the pulmonary circulatory system and / or an increase in pressure within the systemic circulatory system also provides a net force, causing the impeller 130 to move towards the second cavity 122A. This reduces the separation distance between the second set of vanes 132 and the second cavity surface 124A, thereby increasing the efficiency of the pumping effect in the pulmonary system and reducing the efficiency of the pumping effect in the systemic system. Therefore, it increases the flow balance problem in the lungs or the systemic circulatory system.
0074To address this situation, a control process is typically implemented by controller 190, which allows control of the position of impeller 130A within the cavity 120, as described in more detail below.
0075In the example of the heart pump 100B of Figure 1B, this can be coupled to the pulmonary or systemic circulatory system, allowing the pump to act as a VAD (ventricular auxiliary device), the pump to the left of the heart. Supplements the pumping action of the ventricle or right ventricle.
0076In this embodiment, the cardiac pump 100B has a pumping action equal to the pumping action required by each ventricle, with the impeller 130B positioned approximately axially centered within the cavity 120, commonly referred to as the nominal equilibrium point. Arranged like this. This can be achieved by choosing suitable dimensions such as the length, height, and shape of each vane, and generally achieves a flow rate of about 5 L / min at the outlet.
0077In this embodiment, the fluid pressure in the cavities 121, 122B will result in a pressure difference across the impeller 130B depending on the configuration of the impeller 130B, as described in more detail below. The pump 100B also has any force exerted on the impeller 130B, including the force resulting from the pressure difference and the attractive force caused by the magnetic coupling between the impeller 130B and the drive 170 and between the impeller 130B and the bearing 180. When the Impeller 130B is brought to the equilibrium point, it naturally balances to be about the same. Again, a "zero power configuration" is implemented because the equilibrium points within the cavity 120 are typically positioned in the axial center of the cavity 120 when the circulatory system is functioning properly.
0078In this embodiment, if there is a pressure change in the cavities 121, 122B, this will result in a corrected pressure difference on the impeller, causing the impeller 130B to move away from the equilibrium point. The direction in which this occurs depends on the pressure difference created before and after the impeller 130B, which also exacerbates the flow problem.
0079An embodiment of the control process is described herein with reference to FIG. This control process is equally applicable to the cardiac pumps 100A, 100B of the first and second embodiments, and therefore generally refers to the cardiac pump 100 with the impeller 130 and the respective cardiac pumps 100A, 100B. Is specified only when this affects the process.
0080In this embodiment, at step 200, the process comprises determining the movement of the impeller 130 in the first axial direction according to a signal from sensor 195. The sensor is designed to detect the movement of the impeller 130 in any suitable way. For example, this can be achieved by the use of pressure sensors capable of detecting changes in relative pressure within the first and second cavities 121, 122 or within the systemic and pulmonary circulatory system, followed by the impeller. Brings 130 movements. Alternatively, this can be achieved by detecting the distance between the impeller 130 and a sensor 195, such as a suitable position sensor, as described in more detail below.
0081The first direction is the first cavity 121 or the first, depending on the nature of the variation from the normal pressure difference, and in particular whether this is caused by an increase and / or decrease in one or both of the systemic and pulmonary systems. It will be appreciated that 2 can be directed to cavity 122.
0082In step 210, the magnetic bearing 180 is used to move the impeller 130 in the second axial direction opposite to the first axial direction. So, for example, this can include increasing or decreasing the force applied by the magnetic bearing 180, thereby changing the normal pressure difference between the two cavities 121 and 122. Allows the Impeller 130 to move against the forces caused by it.
0083At step 220, an indicator indicating the power used by the magnetic bearing 180 is established. This may be achieved by any suitable method, such as monitoring the draw current by the magnetic bearing 180. This can include, for example, monitoring the current used to keep the impeller 130 in a fixed axial position, or, instead, a magnetic bearing as the impeller 130 moves in the second axial direction. It may include monitoring changes in the current drawn by 180.
0084At step 230, the indicator is used to control the axial position of the impeller 130. This is typically done to minimize the current drawn by the magnetic bearing 180 without the requirement of constraints, thereby providing a zero power configuration even in the presence of normal pressure differential changes. maintain. This situation results in a new equilibrium position offset from the axial center position. In particular, in the configuration described above, the new equilibrium position will be offset from the axial center position in the second direction due to the attractive force present between the impeller 130 and both the drive 170 and the magnetic bearing 180.
0085Therefore, if the impeller 130 is urged towards the second cavity 122 due to an increase in pressure in the first cavity 121, the new equilibrium position is offset from the axial center position towards the first cavity 121. Will be. At this position, the attractive force between the drive 170 and the impeller 130 increases, while the attractive force between the magnetic bearing 180 and the impeller 130 decreases. This provides a greater net force towards the first cavity 121 and balances the increased force towards the second cavity 122 caused by the change in pressure difference.
0086In an embodiment of the cardiac pump 100A with the new equilibrium position offset towards the first cavity 121, the reduced separation distance between the vanes 131 of the first set and the surface 123 of the first cavity is within the first cavity 121. The pumping action of the first discharge port 151 is increased, thereby increasing the discharge amount from the first discharge port 151. Similarly, there is a reduction in pumping action in the second cavity 122A, resulting in a reduction in discharge from the second discharge port 152. Therefore, this works to restore the normal flow balance of the circulatory system. It will be understood that once the normal pressure difference is restored, this will provide a net force towards the first cavity 121. Applying the process described above would return the impeller 130A to a normal equilibrium position with an axial center within the cavity 120.
0087Therefore, in the case of the cardiac pump 100A, the process described above works to maintain zero power control by adjusting the position of the impeller 130A in the cavity 120 when changes in normal pressure difference occur between the circulatory systems, impeller. Allows the 130A to be brought to a new equilibrium position. Further, in one embodiment, the inherent attractive force between the impeller 130A and the drive 170 and the magnetic bearing 180 provides a new equilibrium position offset from the axial center of the cavity 120, thereby providing suction ports 141, 142 and Controlling the relative flow rate between the corresponding outlets 151, 152, the new equilibrium position can then be used to compensate for changes in vascular resistance in the circulatory system.
0088Therefore, it will be understood that this control process allows the cardiac pump 100A to perform relative flow control while maintaining a zero power configuration. In the embodiment of the cardiac pump 100B, the increased pressure in the cavities 121, 122B provides a pressure difference that moves the impeller 130B towards the first cavity 121 before and after the impeller 130B. In this example, the new equilibrium position would therefore be offset towards the second cavity 122B, and the increase in the separation distance between the vanes 131 of the first set and the surface 123 of the first cavity would be the first cavity 121. It results in a reduction in pumping action within. This reduces the flow pressure, thereby offsetting the increase in flow pressure between the first cavity 121 and the second cavity 122B.
0089Therefore, for the cardiac pump 100B, the process described above adjusts the position of the impeller 130B in the cavity 120 when the normal pressure difference before and after the impeller is caused by pressure changes in each circulatory system to which the pump 100B is mounted. By doing so it works to maintain zero power control. Unlike the BiVAD application for pump 100A, compensating for pressure differences is more important than compensating for relative flow rates between circulatory systems. The reason is that the relative flow rate will be affected by the ventricles to which the pump 100B is not attached. Therefore, in the VAD embodiment for pump 100B, the inherent attractive force between the impeller 130B and the drive 170 and the magnetic bearing 180 results in a new equilibrium position offset from the axial center of the cavity 120, thereby eventually. The pressure change of the is also offset, and the new equilibrium position can then be used to compensate for the change in vascular resistance in the circulatory system.
0090Therefore, it will be understood that this control process allows the cardiac pump 100B to perform pressure control while maintaining a zero power configuration. In this embodiment, if flow rate control is required, this can be achieved by adjusting the rotational speed of the impeller 130B. It is understood that this is less complicated than trying to change the rotational speed of the impeller 130A in the cardiac pump 100A, which can result in undesired flow differences due to the different configurations of vanes 131, 132 in the first and second sets. Will be done.
0091The process described above may be accomplished in any suitable manner, but in one embodiment this is accomplished by controller 190. Therefore, the controller 190 typically controls the drive 170 to cause the impeller 130 to rotate. In addition to this, the controller 190 monitors the power drawn by the magnetic bearing 180 and uses it to control the axial position of the magnetic bearing 180 and thus the impeller 130, as described in more detail below. Will be done.
0092It is understood from this that any suitable form of controller may be used, and exemplary controllers are described herein in more detail with respect to FIG. In this embodiment, the controller 190 includes an input / output device (I / O device) 302 such as a processor 300, a memory 301, an input button, a keypad, a display, or a selective external interface 303, wherein the controller 190 It makes it possible to receive a signal from the sensor 195 and provide a control signal to the drive 170 and the magnetic bearing 180. Therefore, the controller 190 may be in the form of a well-programmed processing system such as a computer, laptop, palmtop, PDA, or alternative, dedicated hardware, programmable logic controller, field programmable gate array ( It will be understood that it may be FPGA), or something similar.
0093In one embodiment, the controller 190 is formed from custom microelectronics, allowing the controller 190 to be physically implanted with a heart pump within the subject. Alternatively, controller 190 may be used to control the heart pump 100 via a wireless connection or similar.
0094In use, the processor 300 typically executes the instructions stored in memory 301, allowing the processor 300 to carry out the control processes described herein. In particular, the processor 300 receives a signal from the sensor 195 to allow the impeller position to be determined. Processor 300 then determines if the behavior of the bearing 180 and / or drive 170 needs to be modified, and if so, the appropriate signal applied to the bearing and / or drive as needed. Generate.
0095An embodiment of the process for using the controller 190 to control the position of the cardiac pump impeller is described herein with reference to FIG. For embodiments, this is described for the heart pump 100A of FIG. 1A, but it will be understood that similar operation occurs for the heart pump 100B.
0096In this embodiment, at step 400, controller 190 monitors the axial position of impeller 130A using sensor 195, and at step 410, changes in the pressure difference between the pulmonary and systemic circulation. As a result, it is determined whether or not the impeller has moved. This typically monitors the signal from sensor 195 to determine if the distance between impeller 130A and sensor 195 has changed, thereby moving impeller 130A from its current equilibrium position. Including to inform. The current equilibrium position may correspond to the normal equilibrium position if the circulatory system previously functioned according to normal hemodynamics, but this is not required.
0097At step 420, the controller determines the direction of axial movement by, for example, determining whether the separation distance between the sensor 195 and the impeller has increased or decreased, and at step 430 this has been determined. In use, the magnetic bearing allows the impeller 130A to cause reverse movement, thereby moving the impeller 130A towards a balanced position.
0098At step 440, the controller establishes an indicator based on the rate of change of current used by the magnetic bearing 180 as the impeller 130A moves back towards the equilibrium position. The indicator is compared to a predetermined value in step 450 to determine in step 460 whether the power consumption by the indicator, and thus the magnetic bearing 180, is acceptable. Its value therefore typically represents the minimum power use by a magnetic bearing 180 that meets the zero power requirement, and this value is typically predetermined and prior to the initial operation of the cardiac pump 100, the controller 190. It is stored in the memory 301 of.
0099In general, the power consumed by the magnetic bearing 180 decreases as the impeller 130A approaches the new equilibrium position associated with the changed pressure difference between the circulatory systems, thus indicating whether the impeller 130A has reached the new equilibrium position. It will be understood that this can be used for.
0100If the indicator is determined to be unsuccessful in step 460, the process moves to step 470 to determine the axial position of the impeller 130A. The axial position is then compared to the position limit in step 480, and in step 490 it is determined whether the impeller 130A is within the operating position limit. This is done to ensure that the impeller 130A does not get too close to the first or second cavity surfaces 123, 124A, which can affect impeller performance. Again, the position limit is typically a predetermined value stored in memory 301.
0101If the impeller 130A is still within the position limit but has not yet reached the new equilibrium point, the process returns to step 430 and continues to induce the movement of the impeller 130A.
0102Otherwise, if it is determined in step 460 that a new equilibrium point has been reached, or if the position limit has been reached in step 490, the movement of Impeller 130A will be halted at its current position. The current position therefore represents an axial position as close as possible to the equilibrium position in the zero power configuration.
0103The process can then be returned to step 400, allowing the process to be repeated if further relative pressure changes occur between the systemic and pulmonary systems. As mentioned above, the configuration is such that the new equilibrium position offsets fluctuations that deviate from the normal left / right discharge balance, thereby causing the circulatory system to have the normal discharge required by normal hemodynamics. Return to equilibrium.
0104When applied to the cardiac pump 100B, the process described above offsets any pressure changes in the circulatory system and returns the system to normal hemodynamic pressure. An embodiment of the drive 170 and the magnetic bearing 180 is shown in FIGS. 5-10.
0105In this embodiment, the magnetic bearing 180 includes three equidistant 120 ° U-shaped bearing stator cores 181. A multi-winding copper wire 182 is wound around the radial outer foot of each stator core 181 to generate magnetic flux towards the iron core 185 attached to the impeller 130 (not shown in this example). .. Custom NdFeB permanent magnets 183 may be attached to the radial inner foot of each core to provide bias flux in the flux path. The permanent magnet may also be located on the impeller 130 or within the section bridging the medial and lateral poles 184, and is preferably of high magnetic strength (grade N52). ) Should be. The axial bearing is magnetically coupled to an iron core 185 made of a ferromagnetic material and mounted on the impeller 130.
0106The drive 170 typically includes a slotted axial flux motor configured with up to 12 poles, but is preferably composed of a stator 171 with 6 poles 173 around which a centralized copper coil 172 is wound. .. The motor stator 171 alternates between the north pole 174a and the south pole 174b and is electromagnetically coupled to a circumferentially spaced permanent magnet 174 with opposite polarities attached to a ferromagnetic core 175 mounted on the impeller 130. To do.
0107The bearing stator 181, the motor stator 171, and the ferromagnetic cores 175, 185 are each preferably made of a material exhibiting high electrical resistivity and high magnetic permeability, such as iron cobalt or iron silicon alloy, respectively. The proposed material is VACO FLUX 48 (Vacuumschmelze GMBH & CO.KG, Germany). The materials are laminated to reduce eddy current loss.
0108Axial attraction f generated by magnetic drive and bearing systems 170, 180<sub>Z</sub>And motor torque τ<sub>Z</sub>Is derived from equations (1) and (2). The parameters used in each equation are listed in Table 1.
0109<tables num="1"><img id="000002" he="70" wi="158" file="JP5719829B2_D0001.tif" img-format="tif" img-content="drawing" /></tables>
0110<maths num="1"><img id="000003" he="15" wi="159" file="JP5719829B2_D0001.tif" img-format="tif" img-content="drawing" /></maths>
0111<maths num="2"><img id="000004" he="15" wi="159" file="JP5719829B2_D0001.tif" img-format="tif" img-content="drawing" /></maths> Impeller stator magnetic flux (B)<sub>s</sub>) And permanent magnet magnetic flux (B)<sub>r</sub>) Is assumed to follow the cosine waveform magnetic flux density as described in equations (3) and (4).
0112B<sub>s</sub>(θ, t) = B<sub>S</sub>cos (ωt-Mθ) (3) B<sub>r</sub>(θ, t) = B<sub>R</sub>cos (ωt-Mθ-φ) (4) The number of turns of the coils 172 and 182 and the geometric parameters of the permanent magnets 174 and 183 generate this magnetic flux.
0113The operating principle of axial position control by magnetic bearings is shown in Figure 11A. Motor stator and magnetic bearing magnetic flux B<sub>S</sub>The axial displacement of the impeller 130 can be controlled by changing the sizes of 176 and 186. The motor torque is also controlled by changing the phase difference φ.
0114Permanent magnets 174, 183 in the drive 170 and bearing 180 generate static bias fluxes 177, 187, respectively, to reduce the power requirements of the system. The attractive forces generated by these magnets 174, 183 are in equilibrium when the impeller 130 is located in the center of the cavity 120. Therefore, the control fluxes 176, 186 generated by the coils 172, 182 in the drive 170 and the bearing 180, respectively, are only required to stabilize the axial position of the impeller and overcome disturbance forces.
0115The control flux 186 generated by the bearing 180 may increase the effective attraction between the impeller 130 and the bearing 180 by generating a magnetic field complementary to or vice versa to the magnetic field generated by the permanent magnet 183. , Or can be reduced. This controls the net magnetic field between the bearing and the bearing magnetic material, which in turn allows the position of the impeller 130 to be controlled in either direction within the cavity 120. ..
0116Axial position feedback control of the impeller is achieved in response to impeller displacement, which may be detected by three position sensors 195A, 195B, and 195C, such as eddy current sensors (U5B, Lion Precision, Minnesota, USA). Will be done. These displacement measurements are feedback parameters used by control algorithms that stabilize the system. The control gain is output to the power amplifier, which produces the required current in the corresponding coil to change the magnetic flux density in the magnetic gap, and thus the attractive force, thereby causing impeller flotation. maintain. Similarly, the motor controller generates a three-phase current for the motor coil to achieve synchronous rotation. This rotation may be controlled using a Hall effect sensor or a speed feedback parameter derived from a counter electromotive force (EMF) record.
0117The operating principle of tilt control by magnetic bearings is shown in FIG. 11B. Constant bias magnetic fluxes 177 and 187 are generated by the permanent magnets 174 and 183. When the impeller tilts, the electromagnetic control magnetic fluxes 176 and 186 generated by the axial bearing 180 on the side where the impeller 130 approaches are the reverse magnetic flux 186.<sub>counter</sub>Is reduced by applying, while the magnetic flux on the other side is complementary flux 186<sub>comp</sub>Increased by applying, thereby the recovery tilt force F<sub>Tilt</sub>Occurs.
0118The method by which the impeller position is maintained is described herein with respect to FIGS. 12A-12C. The controller 190 attempts to maintain a stable axial position of the impeller 130 in the center of the pump cavity 120, as shown in FIG. 12A. In this regard, with the impeller centered, the static bias fluxes 177A, 187A generated by these magnets 174, 183 are in equilibrium, so that only the minimum control fluxes 176A, 186A drive. It is required to be produced by 170 and bearings respectively.
0119The introduction of the static disturbance force F, shown in FIG. 12B, attempts to displace the impeller 130 from this central position (in this case towards the cavity 121). A slight deviation in the axial position of the impeller 130 involves a "zero position" controller 190 that increases the magnetic control flux 186B generated by the bearing 180, these forces F are offset and the magnetic bearing The conceptual center "zero" position that results in the increase in magnetic power used by 180 is maintained.
0120However, by implementing the "zero power" controller 190, the conceptual central axial position of the impeller 130 is not fixed at the physical center position within the cavity 130, but in response to external disturbances. It will change.
0121In the previous embodiment, the application of the disturbance force F toward the cavity 121 is offset by an increase in the permanent magnet attraction of the bearing bias flux 187C and a decrease in the subsequent permanent magnet attraction of the motor bias flux 177C. The controller 190 moves the impeller 130 in the opposite direction toward the bearing 180 until. This allows the bearing electromagnetic control flux 186B to be reduced until the motor and control fluxes 176C, 186C are minimized and used again only to stabilize the impeller in the new axial position.
0122As briefly mentioned earlier, controlling a magnetic bearing in zero power mode can be advantageous when operating a cardiac pump in a cardiovascular system that is constantly adapting and changing its physiological parameters. These parameters specifically affect the pressure generation within the cardiac pump, which imposes static disturbance pressure on the impeller in the axial and radial directions.
0123The ability to alter the hydraulic output of VAD or BiVAD is important to enable effective physiological movement within the cardiovascular environment. Changes in perfusion requirements due to changes in the patient's physical condition must be met by a cardiac pump. In addition, the potential for chamber deviation due to overpumping of the cardiac pump must be prevented or corrected immediately after these events occur by reducing the cardiac pump discharge.
0124The most common technique for achieving this hydraulic performance change is to change the impeller speed. However, the axial displacement of the impeller within the pump cavity also changes the hydraulic output of the pump by inducing a leak change in the flow from the high pressure outlet to the low pressure suction port. This technique effectively modifies the energy efficiency of the impeller vane set and is most effective when a semi-open (ie, no upper shroud) impeller is used.
0125As shown in FIG. 13B, with reference to a single VAD such as the cardiac pump 100B shown in FIGS. 13A-13C, centered on the impeller 54 within the pump cavity 50 and operating at a set rotational speed around the axis 59. In addition, the desired hemodynamics are generated for the circulatory system in need of support. Shifting the impeller away from the suction port 57 along its axis of rotation 59, as indicated by arrow 51c, without changing the velocity, effectively increases the clearance on the impeller vane. This movement reduces the VAD discharge rate at the discharge port 58, as indicated by arrow 53c in FIG. 13C. The decrease in discharge rate is a direct result of the increase in leakage from the high pressure discharge port 58 to the low pressure suction port 57, as indicated by the arrow 55c. Similarly, shifting the impeller towards the suction port 57, as shown by arrow 51b, has the opposite effect, that is, increasing the VAD discharge rate, as shown by arrow 53b in FIG. 13A.
0126Therefore, the axial movement of the Impeller 130B can be used to deliver a discharge flow from the VAD that is sufficiently variable to meet the cardiovascular physiological requirements of the set rotational speed. This effect is due to the low impeller blade height, the ratio of the impeller blade height to the axial clearance being about 3: 1 (blade height = 1.4 mm, starting axial clearance = 0.5 mm), and sufficient discharge fluctuation. More prominent when achieved by +/- 0.3 mm movement.
0127In certain applications, the ability of the BiVAD system to accurately change the left and right discharges is a relative change in systemic and pulmonary vascular resistance, relative levels of ventricular contractility, reduction of acute lung or systemic congestion. It is important, especially during the postoperative period, to cope with the disease and to prevent suckdown by maintaining proper atrial filling pressure. This requirement is also achieved by axially displacing the impeller 130A within a biluminal cardiac pump, such as the cardiac pump 100A, thereby maintaining a long-term balance of left and right flow rates.
0128With reference to FIGS. 14A-14C, examples of the effect of axial position on relative flow rates for BiVADs such as the cardiac pump 100A are described here. This embodiment is described with respect to the heart pump of FIG. 1A, therefore similar symbols are used to specify similar features.
0129In this embodiment, the impeller 130A is placed in the pump cavity 120 at the axial center of the cavity shown by line 20 and operates around a shaft 19 at a set rotational speed of about 2300 rpm, which is a systemic system. And for the pulmonary system, it results in the production of the desired hemodynamics of 100 mmHg (LVAD) and 20 mmHg (RVAD). Therefore, the discharge amounts through the first and second discharge ports 151 and 152 are in an equilibrium state at about 5 L / min, as shown by arrows 23a and 24a in FIG. 14B. Due to the difference in the natural discharge of the heart caused by the bronchial circulation, the exact discharge from the left cavity is slightly higher than the right cavity.
0130Shifting the 0.3mm impeller along its axis of rotation 19 towards the second cavity 122A, as indicated by arrow 21c, without changing the velocity, causes an axial clearance (0.2mm) on the RVAD vane. Decrease (up to), while at the same time increasing the clearance above the LVAD vane (up to 0.8 mm). This movement improves the RVAD discharge rate through the discharge port 152 as shown by arrow 24c in FIG. 14C, while reducing the LVAD discharge rate via the discharge port 151 as shown by arrow 23c. The decrease in discharge rate is a direct result of the increase in leakage from the high pressure discharge port 151 to the low pressure suction port 141, as indicated by the arrow 25c.
0131Similarly, shifting the impeller from the center position towards the first cavity 121 has the opposite effect, as indicated by arrow 21b. For example, this 0.3 mm movement towards the LVAD cavity 121 while maintaining arterial pressure increases the left discharge rate to 6.4 L / min, as shown by arrow 23b in Figure 14A, while the right discharge rate is. As shown by the arrow 24b, it decreases to 4.6 L / min and shows an instantaneous discharge rate difference of 1.8 L / min (36%). Therefore, this axial movement can be used to address the required variable discharge rates from the left and right hearts at set rotational speeds. This effect is more pronounced when the impeller blade height is low and the ratio of the impeller blade height to the axial clearance is about 3: 1.
0132Graphs showing examples of pump relative performance, in particular pressure and discharge at outlets 151, 152, are shown in FIG. 15 at 1500 and 1510, respectively. An exemplary BiVAD-configured heart pump similar to the heart pump 100A of FIG. 1A is described herein, with additional features shown in Figures 16A-16H. It will also be appreciated that similar features may be implemented in VAD, if appropriate.
0133In this embodiment, the cardiac pump draws suction from the right atrium through the second suction port 142 and from the left ventricle through the first suction port 141 and into the pulmonary artery through the second discharge port 152. It is discharged to the aorta through the first discharge port 141. In this embodiment, the first and second sets of vanes 131, 132 are mounted on a shared rotating hub to form a magnetically and hydrodynamically levitated centrifugal impeller 130A. The vanes 131, 132 of the first and second sets have different outer diameters to generate the pressure required from the systemic and pulmonary systems at a common rotational speed, as described above. The difference in discharge required from the left and right hearts is achieved by changing the axial clearance above each semi-open VAD impeller.
0134The levitation system incorporates a fluid journal bearing 115, while swirl chambers 111, 112 are provided as part of the housing 110, thereby assisting the transfer of fluid to outlets 151, 152. The spiral chamber may be any combination of spiral / single, split / double, or circular / concentric types, with the latter circular spiral chamber type being preferred. The reason is that this configuration produces a stabilizing radial hydraulic pressure for optimum journal bearing function.
0135As outlined above, axial hydraulic pressure is imposed on the impeller 130A by the accumulation of pressure in the left and right cavities. The difference in these pressures acting on the top and bottom of the impeller will generate the resulting force in the positive or negative axial direction.
0136For dual cardiac support conditions, the most important parameter to maintain is the balance of left and right discharges from each support pump. Discharge volume inconsistencies can lead to a possibly catastrophic situation of left or right chamber deviation. The application of the zero power controller and its ability to adapt in favor of device performance are described below for some predictable physiological conditions encountered by BiVAD.
0137Figures 17A-17D show the resulting axial force generation on the double-sided impeller 130A under various general conditions. During the expansion phase, the force acting on the surface of the semi-open LVAD impeller decreases from the outer diameter to the inner diameter in response to pressure generation along this path. This is balanced by the force acting under the impeller 130A, which consists of a section exposed to high LVAD outlet pressure (decreased less than the degree of diameter reduction) and an RVAD impeller section exposed to the low pressure generated within the RVAD cavity. To do.
0138This equilibrium is disrupted during contraction, as shown in Figure 17B. In this example, left ventricular pressure acts over the entire LVAD impeller surface and cannot be matched by lower RVAD pressure. Due to this disturbance of blood and the impulse nature of damping, the actual forces encountered by magnetic bearings are lower than static differences suggest.
0139The effect of increased systemic vascular resistance (SVR) on pressure generation and thus force generation is described in Figures 17A, 17B, and 17C. Increasing the SVR acts to increase the pressure in the first cavity 121 of the LVAD, which in turn creates a force towards the second cavity 122A of the RVAD. The same force obtained is generated when pulmonary vascular resistance (PVR) is reduced. This logic is used to generate the opposite force towards the LVAD cavity as the SVR decreases and / or the PVR increases.
0140Finally, the force generated on the impeller at the moment of departure from the left chamber is described in Figure 17D. In this example, the pressure generated in the inlet cannula drops well below 0 mmHg and the outlet drops to a correspondingly lower pressure value. This combination contributes to a large axial force towards the LVAD cavity.
0141Figures 18A-18C describe the adaptation of the cardiac pump 100A to changes in relative vascular resistance. The examples presented are for a relative increase in SVR compared to PVR, but can describe a relative decrease in PVR compared to SVR. The opposite will happen with a relative decrease in SVR or increase in PVR.
0142During normal operation as shown in FIG. 18A, the impeller 130A is centered within the cavity 120 and the left and right discharges from the first and second suction ports 151, 152 are balanced as indicated by arrows 36a, 35a. In state. Since the impeller 130A is located in the center of the pump cavity 120, the bearing and motor permanent magnet (PM) bias forces indicated by arrows 34a, 31a are also in equilibrium. The electromagnetically controlled flux forces from the drive 170 and bearing 180 are minimal, as indicated by arrows 32a, 33a, and are required only to stabilize and compensate for dynamic disturbances.
0143However, the increase in SVR shown in FIG. 18B caused a decrease in LVAD discharge from the discharge port 151 and an increase in LVAD cavity pressure indicated by arrow 36b, which, as previously mentioned in FIG. 17C, was the RVAD cavity 122A. Provides a static axial hydraulic pressure vector towards. To keep the impeller 130A in the center position, the magnetic bearing flux must increase in magnitude (and therefore power) to generate an opposite magnetic field opposite the magnetic field of the permanent magnet 183, as indicated by arrow 33b. It does not, thereby providing resilience.
0144However, this increased SVR must be overcome to restore LVAD discharge. Implementing the "zero power" controller described above allows the impeller, LVAD cavity 121, until the disturbance forces indicated by arrow 30 are balanced by the increase in permanent magnet bias flux from the drive 170, as indicated by arrow 31c. Will be moved towards. At this equitable position, the electromagnetic control flux is once again minimized and is only required in the case of dynamic disturbances, thus reducing power consumption. However, most importantly, the LVAD discharge rate from the discharge port 151 increases as indicated by arrow 36c, while the RVAD discharge rate from the discharge port 152 indicated by arrow 35c decreases slightly and is balanced. Be maintained. Should it be necessary to maintain an absolute discharge level, the rotational speed may also be increased to increase both LVAD and RVAD discharges simultaneously.
0145Figures 19A-19C describe the adaptation of the cardiac pump 100A to chamber deviation. The examples presented describe the consequences of left chamber deviation. However, the opposite property occurs in the case of right chamber deviation.
0146In the embodiment of FIG. 19A, the left heart chamber deviation impairs the flow of blood entering the suction port 141 on the left side of the cardiac pump as shown by 47, and the LVAD discharge rate from the discharge port 151 as shown by arrow 46a. Brings a serious reduction. As shown in Figure 17D, an axial force vector towards the LVAD cavity occurs subsequently, which force vector must be offset by an increase in the bearing magnetic control flux as shown by arrow 43a, thereby centering. The impeller position is maintained.
0147However, the "zero power" control method also automatically positions the impeller axial position towards the RVAD cavity 122A, as shown in FIG. 19B, until the bearing permanent magnet bias force indicated by arrow 44b balances the disturbance pressure 40. Adjust. This action returns the electromagnetic control flux generated by the drive 170 and bearing 180 to the minimum level, as indicated by arrows 43b, 42b, thus reducing power consumption. Further, the RVAD discharge amount from the discharge port 152 increases, and the pressure difference from the LVAD suction port 141 to the discharge port 151 decreases. This continues to shift blood into the lung circuit and, as a result, into the left heart chamber, thus reducing the deviation of the left heart.
0148The LVAD and RVAD cavity pressures then return to normal, eliminate disturbances and allow the impeller 130A to automatically translate back to the center of the cavity 120, which in turn is arrow 46c, As shown by 45c, the balance of LVAD and RVAD discharge amounts from the discharge ports 151 and 152 is restored.
0149In another embodiment, a single VAD cardiac pump, such as the cardiac pump 100B, is designed to provide ventricular assist to one side of a failed heart. In this embodiment, axial hydraulic pressure is imposed on the impeller 130B by the accumulation of pressure in the first and second cavities 121, 122B. The difference in these pressures acting on the top and bottom impeller surfaces will generate the resulting force in the positive or negative axial direction. The pressure difference is caused by the flow between the suction port 141 and the discharge port 151, resulting in a pressure gradient in front of and behind the surface of the impeller 130B including the vanes, and the pressure gradients on the other surfaces are substantially constant.
0150Figures 20A-20D show the resulting axial force generation on the impeller under various general conditions. During the expansion phase, as shown in FIG. 20A, the force below the impeller 130B is only partially balanced by the force acting on the top of the impeller, resulting in a net force towards the suction port 141 as indicated by arrow 2000. This force decreases from the outer diameter to the inner diameter as pressure is generated along this path. This imbalance is reduced during contraction, as shown in FIG. 20B, when left ventricular pressure acts across the VAD impeller surface.
0151The effect of increased systemic vascular resistance (SVR) on pressure generation and thus force generation is described in Figure 20C. Increasing the SVR acts to increase the overall pressure in the pump cavity, which in turn increases the inconsistency of forces acting below and above the impeller 130B, towards the suction port 141. The net power indicated by the arrow 2010 is brought about. Finally, the force generated on the impeller during the moment of chamber deviation is described in Figure 20D. In this example, the pressure generated in the inlet cannula drops well below 0 mmHg and the outlet drops to a correspondingly lower pressure value. This combination contributes to the increase in axial force towards the suction port 141, indicated by arrow 2020.
0152For single heart support applications, the two parameters that require the most control are the discharge pressure (and therefore the discharge rate) and the correction of cardiac deviation in response to changes in vascular resistance. The ability of a zero power controller to adapt favorably to cardiac pump performance is described below for some predictable physiological conditions encountered by a single VAD.
0153Figures 21A-21C describe the adaptation of cardiac pump 100B to increased vascular resistance. In FIG. 21A, during normal operation, the cardiac pump circulates the discharge pressure and discharge provided by the cardiac pump outlet 151, with the impeller positioned approximately centered within the cavity 120, as indicated by arrow 63a. It is configured to meet the physiological requirements of the system. In this configuration, with the impeller 130B located in the center of the pump cavity 120, the bearing 180 and drive 170PM bias forces indicated by arrows 67a and 64a are in equilibrium. The electromagnetically controlled flux forces from the drive 170 and bearing 180 are minimal, as indicated by arrows 65a, 66a, and are required only to stabilize and compensate for dynamic disturbances. ..
0154An increase in vascular resistance (VR) shown in FIG. 21B causes an increase in VAD outlet pressure, and thus cavity pressure, indicated by arrow 63b, as indicated by arrow 68, as described in FIG. 20C. , Brings an increase in static axial hydraulic pressure towards the first cavity 121. In order to keep the impeller 130B in the center position, the magnetic bearing flux must increase in magnitude (and therefore power), as indicated by arrow 66b, which provides resilience.
0155However, in a single VAD application, this increase in outlet pressure due to increased VR should be reduced by the innate baroreceptor reflex. Discharge rate changes can be achieved by velocity changes, while the discharge rate balancing problem is counteracted by the remaining functioning ventricular ability to balance the discharge rate. Implementing a "zero power" controller as described above causes the impeller to move away from the first cavity 121 until the disturbance force 68 is balanced by the increase in PM bias force from the bearing 180, as indicated by arrow 67c. Will be moved like this. At this equitable position, the electromagnetic control flux is once again minimized and is only required in the case of dynamic disturbances, thus reducing power consumption. However, most importantly, the VAD discharge pressure indicated by arrow 36c returns to a low value, while the discharge volume decreases (due to increased vascular resistance). The opposite characteristic is observed for the decrease in vascular resistance (impeller movement towards the suction port) encountered during exercise. Therefore, automatic impeller movement maintains vascular pressure and therefore increases the discharge rate to suit.
0156However, this increase in discharge can result in a situation of excessive pumping and thus chamber deviation. Figures 22A-2C describe the adaptation of the cardiac pump 100B to this event.
0157When a heart chamber deviation occurs as shown in FIG. 22A, this impairs blood flow into the suction port 141, as shown by arrow 79, and the VAD discharge rate at the discharge port 151, as shown by arrow 73a. Brings a serious reduction in. As described in FIG. 20D, an axial force vector towards the suction port 141 (which must be offset by an increase in the bearing magnetic control flux indicated by arrow 76a) is generated to maintain the centered impeller position.
0158However, the "zero power" control method automatically positions the impeller axial position away from the suction cavity 141 until the bearing PM bias force indicated by arrow 77b in FIG. 22B balances the disturbance pressure indicated by arrow 78. Adjust to. This action returns the electromagnetically controlled magnetic flux indicated by arrows 75b, 76b to a minimum level and thus reduces power consumption. Therefore, the pressure difference from the VAD inlet to the outlet is reduced, thus avoiding the deviation of the left heart. The VAD cavity pressure then returns to normal, eliminates disturbances, allows the impeller to automatically translate back to the center of the cavity, and normal discharge as shown by arrow 79c in Figure 22C. Return to.
0159Experiments were performed using a heart pump similar to the heart pump shown in Figure 6A, which was configured to operate according to the control process of Figure 9. Cardiac pumps simulate various hemodynamic conditions such as pulmonary hypertension, LVAD inflow obstruction / left ventricular suckdown, and systemic hypertension, thereby assessing pump responsiveness. Coupled to a fluid circulation loop designed to enable.
0160All parameters and hemodynamics obtained are described in Table 2 and shown in Figures 23A-23E, a detailed description will continue to be provided. The simulated state is summarized as follows.
01611. Normal state (Fig. 12A) 2. Pulmonary hypertension (Fig. 12B) 3. Right impeller movement to compensate for PHT (12C) 4. Normal condition 5. Suction event (Fig. 19A) 6. Right impeller movement to compensate for suction (19B) 7. Normal state (Fig. 18A) 8. Generalized hypertension (Fig. 18B) 9. Left impeller movement to compensate for SHT (18C)
0162<tables num="2"><img id="000005" he="140" wi="159" file="JP5719829B2_D0001.tif" img-format="tif" img-content="drawing" /></tables> In the first example, pulmonary hypertension was simulated. In fact, these events that occur within the human body have the potential for reduced flow through the pulmonary system (and thus reduced venous return to the left heart) and suction of the left heart chamber. It will bring about an inevitable increase. Examples of cardiac pump responses in these situations are as described above with respect to FIGS. 12A-12C.
0163With reference to FIG. 12A, in the initial "normal" state (state 1), the impeller 130A is centered by normal pressure and a flow rate is generated. It allows a diameter of 50 mm within the 1st LVAD cavity 121 and a diameter of approximately 25 mm within the 2nd RVAD cavity 122A (vanes have a height of 1.4 mm), as well as +/- 0.5 mm axial impeller movement. Achieved using an impeller 130A with a starting axial clearance of 0.5 mm.
0164In this case, the liquid pressure caused by the differential pressures from the RVAD and LVAD cavities 122 and 121, respectively, is equal to + 4.2N towards the LVAD cavities 121. This is instantly balanced by the magnetic bias force 187A caused by the permanent magnets in the magnetic bearing 180. The bearing current is only required for the disturbance force and therefore the bearing electromagnetic force 186A, and therefore the power usage, is equal to zero as shown in Table 2.
0165The occurrence of pulmonary hypertension is subsequently simulated (state 2), the occurrence of pulmonary hypertension is described in Figures 12B, and the results are shown in Figures 23A-23E. Hemodynamic results are an immediate decrease in flow through the pulmonary system and an immediate decrease in left atrial pressure from 9 to 7 mmHg. In this case, the fluid pressure F towards the LVAD / motor is the increased pulmonary vascular resistance 0.00168 Nscm<sup>5</sup>(168dynes.s.cm<sup>5</sup>) Increases by 0.8N to + 5.0N. To keep the impeller 130A in the center position, the controller 190 inputs current to the magnetic bearing 180 to generate an additional 0.8N electromagnetic force 186B and therefore all external at the expense of an additional 0.672W bearing power. Balance the forces. If this condition remains, the reduced flow through the pulmonary system can further reduce left atrial pressure and induce aspiration in the left heart.
0166To avoid this situation, controller 190 then "automatically responds (automatically). Generates a "response)" state (state 3) and Figure 12C, which causes the impeller 130A to automatically move -0.2 mm towards the RVAD cavity. RVAD then improves its hydraulic efficiency, increasing pulmonary arterial pressure, pushing more flow into the left heart and stopping the decrease in left atrial pressure. This reduces the hydraulic pressure on the impeller 130A by 0.8N and returns it to + 4.2N. The movement of the impeller 130A towards the magnetic bearing 180 causes the magnetic bias force 187C generated by the permanent magnets in the magnetic bearing 180 to increase in magnitude to 5.3N. Therefore, a small canceling current of -0.25 amperes is required to flow through the bearing coil in order to reduce the bearing force to + 4.2N and restore the force balance. This results in a bearing power of 0.26W, which is less than the bearing power before moving (0.672W), thus demonstrating the effect of the controller 190 on minimizing the bearing power when operating in the "zero power" mode described above. These results show that the pump uses minimal power when the pump's hemodynamic response works as expected and is in the autoresponder state (state 3) required to offset the pulmonary hypertension state. Emphasize what to do.
0167In the second example, LVAD inflow failure was simulated. In practice, these events that occur within the human body will result in a decrease in flow through the pulmonary system (due to increased PVR) and an inevitable increase in the likelihood of aspiration of the left heart chamber. Examples of cardiac pump responses in these situations are as described above with respect to FIGS. 19A-19C.
0168The cardiac pump 100A and the circulatory system are in an initial "normal" state (state 4), which is substantially the same as in the normal state (state 1) described above, and are therefore not described in more detail.
0169The occurrence of left ventricular aspiration / disorder is then simulated as described in Figure 19A (state 5). The result of hemodynamics is an immediate decrease in left atrial pressure. In this case, the hydraulic pressure 40 towards the LVAD / motor increases. To keep the impeller in the center position, the controller inputs an electric current to the magnetic bearing to generate an additional electromagnetic force (43a).
0170The controller 190 then seeks the minimum power position, which moves the impeller 130A in the direction of the RVAD cavity 122 by -0.1 mm. At this position, 12N of electromagnetic force 43a is still required to balance all external forces at the expense of 20W bearing power. As a result, LAP decreases from 9mmHg to 3mmHg. If this condition persists, reduced pulmonary flow may further reduce left atrial pressure, inducing complete aspiration in the left heart.
0171The controller 190 then performs the automated response state (state 6) shown in FIG. 19B to move it to a position of -0.5 mm towards the RVAD cavity 122, thereby minimizing the power used by the magnetic bearings. To do. Then, LVAD reduces its hydraulic efficiency, which reduces both aortic pressure and flow, while RVAD improves its hydraulic efficiency, which increases both pulmonary arterial pressure and flow, with the left heart. Extrude a large flow rate to stop the decrease in left atrial pressure and raise it back to 9 mmHg, thus correcting the aspiration event.
0172This movement reduces the force on the rotor by 1.9N (40) to + 10.1N. The impeller movement towards the magnetic bearing causes the magnetic bias force 44b generated by the permanent magnets in the magnetic bearing to increase in magnitude. A small canceling current of only -0.75 amperes is required here to flow through the bearing coil in order to reduce the bearing force to + 10.1N and restore the force balance. This results in a bearing power of 2.36W, which is less than the bearing power before moving (20W), thus demonstrating the effectiveness of a zero power controller that minimizes bearing power.
0173Further slight movements back towards the LVAD cavity 121 would then fine-tune the impeller operating position by increasing the permanent magnet bias force, returning a true zero power reading.
0174In the third example, systemic hypertension was simulated. In fact, these events that occur within the human body reduce the flow through the systemic system (and thus the venous ring flow to the right heart) and inevitably increase the likelihood of aspiration of the right heart chamber. Will bring. Examples of this are shown in Figures 18A-18C.
0175In FIG. 18A, the cardiac pump 100A and the circulatory system are in an initial "normal" state (state 7), which is substantially the same as in the normal state (state 1) described above, and thus in more detail. I can't say it.
0176The development of generalized hypertension is then simulated as shown in Figure 18B (state 8). The result of hemodynamics is an immediate decrease in right atrial pressure. In this case, the hydraulic pressure towards the LVAD cavity 122 decreases. Therefore, a larger electromagnetic power is required by the magnetic bearing 180 in order to increase the magnetic bearing force 33b so as to maintain the set central impeller position.
0177In this way, the impeller 130A may be allowed to move -0.1 mm towards the RVAD cavity 122. In this case, the fluid pressure towards LVAD is an excessively increased systemic vascular resistance 0.0295 Nscm.<sup>5</sup>(2950dynes.s.cm<sup>5</sup>), It is further reduced by 1.5N to + 2.7N. To keep the impeller 130A in this position, the controller inputs a canceling current into the magnetic bearing to reduce the bias electromagnetic force 34b by the required 1.5N, thus adding an additional 2.36W bearing power. Balance all external forces at the expense. If this condition persists, the reduced LVAD discharge may further reduce the right atrial pressure and induce aspiration in the right heart.
0178To avoid this situation, the controller 190 then performs an automated response state (state 9) as shown in Figure 18C, which causes the impeller 130A to automatically only +0.5 mm towards the LVAD cavity 121. Moving. The LVAD then improves its hydraulic efficiency, increasing both aortic pressure and flow, pushing more flow into the right heart and stopping the decrease in right atrial pressure. This further reduces the hydraulic pressure on the rotor by 0.2N (30) to + 2.5N. The movement of the impeller towards the drive 170 causes the magnetic bias force 34c generated by the permanent magnets in the magnetic bearing to decrease in magnitude. In this case, the total liquid pressure is almost balanced by the permanent magnet bias force, and only 0.26 W magnetic bearing power is required to balance the force. Therefore, the magnetic bearing current 33c essentially returns to the minimum, and thus the minimum power state is observed.
0179The results described above emphasize that the control process can therefore automatically adjust the cardiac pump discharge rate in response to the state of the system in which the cardiac pump operates. In particular, the system uses axial magnetic bearings and drives to levitate and rotate the impeller of a centrifugal blood pump operating under zero power control. This control process acts to automatically adjust the axial position of the impeller in response to changes in the axial hydraulic pressure applied to the impeller. This technique allows the cardiac pump to simulate the flow balancing characteristics of Frank-Starling's cardiac law and therefore automatically adapt to changes in atrial pressure (preload) and vascular resistance (afterload). To do. This is particularly advantageous for preventing perhaps catastrophic deviations of the left or right heart chamber and for overcoming changes in vascular resistance. One exemplary application of the controller relates to its use in a biventricular auxiliary heart pump. This cardiac pump includes an impeller with left and right vanes positioned on a shared rotating hub that has completely surfaced in the blood. This levitation system incorporates an electromagnetic motor and axial magnetic bearing system for axial levitation and drive, while radial bearings are achieved using fluid journal bearings. This journal bearing is completely cleaned by a unique shunt flow from left to right cavity. The left and right vanes have different outer diameters to generate the pressure required from the systemic and pulmonary systems at a common rotational speed.
0180The instantaneous discharge difference required to balance the discharge from the left and right hearts is achieved by changing the axial clearance above these semi-open vanes. Therefore, the axial movement towards the left cavity will reduce the axial clearance above the left vane set and thus increase the left heart discharge. At the same time, the gap above the right vane set will increase, thus reducing the right heart discharge. Similarly, moving to the right cavity will provoke the opposite effect.
0181This allows the cardiac pump to automatically adjust the discharge volume of the left and right cavities to minimize the possibility of deviation of the left and right heart chambers and to compensate for relative changes in vascular resistance. To do. For example, in the case of a heart chamber deviation, the hub will translate away from the deviation side, thus reducing suction at the suction port while increasing discharge from the opposing pump as well. Upon encountering an increase in relative vascular resistance, the impeller will translate towards that side with increased afterload, thus ensuring that the affected side overcomes the resistance and maintains discharge balance. to enable.
0182This enables Frank-Starling-like control of this discharge balance, which is automatically achieved by incorporating a zero-power magnetic control algorithm. In another exemplary application, controller 190 is used for a single ventricle assisted heart pump. This cardiac pump includes an impeller with a single set of vanes positioned on a rotating hub that floats completely in the blood. This levitation system also incorporates an electromagnetic motor and axial magnetic bearing system for axial levitation and drive, while radial bearings are achieved using fluid journal bearings. This journal bearing is completely cleaned by the impeller through holes that allow blood flow along the underside of the impeller.
0183Pressure control is achieved by an axial clearance above these semi-open vanes. Therefore, the axial movement towards the suction port will reduce the axial clearance above the vane and thus increase the pressure generated in the pump.
0184This allows the cardiac pump to automatically adjust the pressure generated by the pump to minimize pressure changes in the circulatory system while minimizing the possibility of chamber deviation. For example, in the case of a heart chamber deviation, the hub will translate away from the suction port, thus reducing suction at the suction port. Upon encountering an increase in vascular pressure, the impeller will translate away from the suction port, thus providing pressure relief.
0185Therefore, the system described above may provide a VAD capable of automatically controlling the discharge rate in response to the needs of the circulatory system. It uses an axial magnetic bearing system that runs a zero power controller that adjusts the axial position of the impeller in response to hydraulic pressure. In one embodiment, the motor and bearing arrangement configurations each provide a net suction force to the impeller 130, allowing the impeller to be brought to an equilibrium point where its position depends on the relative pressure in the pump cavity. With proper placement configuration, this can be used to provide relative discharge control for BiVAD applications and relative pressure control for VAD applications.
0186In one embodiment, the combination of axial movement (+/- 0.3 mm) and small impeller blade height (1-2 mm) produces a sufficient change in discharge hydraulics at a constant rotational speed. This may provide sensitivity to atrial (preload) and arterial (afterload) pressures, similar to Frank-Starling's law, as well as allow correction of chamber deviations. This is because the cardiac pump maintains a suitable left / right discharge balance, as in the baroreceptor reflex, depending on the atrial pressure in the BiVAD embodiment and the arterial pressure set in the left or right VAD embodiment. To enable.
0187Therefore, this avoids the need to change the centrifugal pump rotation speed to generate the pump performance changes used in conventional cardiac pumps. This performance change receives feedback from hardware sensors and also uses software algorithms to avoid complex and active physiological control algorithms that control pump speed, while circulatory physiology. It is advantageous because it meets the requirements. These hardware sensors pose additional reliability issues that limit the long-term durability of the cardiac pump, while software estimation introduces complexity.
0188The control process also addresses the issue of discharge balance from the left and right pumps in both chamber auxiliary systems, especially from the single rotation impeller system. Therefore, the control process can automatically and passively adjust the hydraulic output of its rotational impeller without changing the rotational speed, and it does not rely on feedback from hemodynamic sensors or software estimates. Provides a controller for. Instead, the impeller modifies its output in response to changes in preload and afterload, similar to Frank-Starling's heart law.
0189In one embodiment, the control process is accomplished by incorporating an axial magnetic motor and bearings that implement a zero power controller that automatically adjusts the axial position of the centrifugal impeller in the pump cavity. The zero power controller responds to changes in axial hydraulic pressure encountered as the pump's preload and afterload change. Performance changes due to axial movement are most effectively observed when the impeller incorporates a set of semi-open (no shroud) blades.
0190The control process also provides the ability of the two-chamber auxiliary system to automatically adjust its discharge rate to maintain a suitable left / right discharge rate balance. This is achieved more efficiently with a single rotation type centrifugal heart pump that implements a "zero power" controller and is also designed to provide biventricular assistance.
0191Those skilled in the art will appreciate that numerous modifications and changes will be revealed. All such modifications and alterations manifested to those skilled in the art should be considered to fall within the spirit and scope of the invention, which appears widely before being stated.
0192For example, the functionality provided by a separate motor and bearing configuration is a combinatorial configuration where one end of the housing contains a set of passive attraction magnets while the other end of the housing contains a combinatorial motor and bearing windings. Can be achieved using the configuration.
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| EP2419160A1 | European Patent Office (EPO) | A1 | |
| US2012095280A1 | United States of America | A1 | |
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Numbers
- Publication
- 5719829
- Publication, DOCDB
- 5719829
- Publication, EPODOC
- JP5719829B
- Application
- 2012504998
- Application, DOCDB
- 2012504998
- Application, EPODOC
- JP20120504998
Titles2
- Japanese
- 心臓ポンプコントローラおよび心臓ポンプを制御する方法
- English
- Heart pump controller and how to control the heart pump
Classification
- CPC, 13
- F04D13/0633
- F04D29/048
- F04D29/042
- A61M2205/3334
- A61M60/422
- A61M60/546
- A61M60/232
- A61M60/237
- A61M60/178
- A61M60/531
- A61M60/183
- A61M60/822
- A61M60/148
- IPC, 1
- A61M1 10
