Method for operating a supply device which supplies a liquid to a channel, and supply device, hollow catheter, and catheter pump

Abstract

The present invention relates to a supply device for a channel (8), in particular within a hollow catheter (1), and to a method for operating a supply device of this type that supplies a channel (8) with a liquid and has two pumps (10, 19) arranged at points of the channel distanced from one another, characterised in that the parameter values of at least one operating parameter of both pumps are coordinated with one another in a controlled manner. As a result of the method, interruption-free and precisely controllable operation is to be ensured with simple structural means, in particular in the case of use of wear-free diaphragm pumps.

Claims

1. A hollow catheter comprising: a channel comprising; an inflow channel region and an outflow channel region fluidly connected in series; a supply device for supplying the channel with a liquid, wherein the supply device comprises: a diaphragm pump; and a control device, which controls the diaphragm pump in respect of a generated pressure or a delivery rate to prevent blood from entering the channel; wherein a sensor is assigned to the diaphragm pump, said sensor being a liquid pressure sensor, an electric sensor for detecting a power consumption of the diaphragm pump, or a flow rate sensor.

2. The hollow catheter of claim 1, comprising one or more further diaphragm pumps, wherein: the diaphragm pump and the one or more further diaphragm pumps are arranged at points of the channel distanced from one another; the control device is configured to control each of the diaphragm pump and the one or more further diaphragm pumps individually in respect of the generated pressure and/or the delivery rate; a sensor is assigned to each of the one or more further diaphragm pumps; wherein, each of the sensors is assigned to the one or more further diaphragm pumps is one of a liquid pressure sensor, an electric sensor for detecting a power consumption of the one or more diaphragm pumps, and a flow rate sensor.

3. The hollow catheter of claim 1, wherein a catheter pump comprises the hollow catheter.

4. The hollow catheter of claim 3, wherein the catheter pump is configured for intraventricular operation within a heart.

5. A method for operating a supply device that supplies a channel of a hollow catheter with an inflow channel region and an outflow channel region fluidly connected in series with a liquid, the supply device having at least one diaphragm pump with a sensor assigned thereto, the method comprising: using the sensor, detecting an operating parameter of the at least one diaphragm pump, the operating parameter being one or more of liquid pressure at one or more points in the channel, power consumption of the at least one diaphragm pump, or flow rate; calculating a generated pressure or delivery rate based on the operating parameter detected by the sensor to prevent blood from entering the channel; and controlling the at least one diaphragm pump according to the generated pressure and/or delivery rate.

6. A method for operating a supply device that supplies a channel of a hollow catheter with an inflow channel region and an outflow channel region fluidly connected in series with a liquid and has at least two diaphragm pumps, arranged at points of the channel distanced from one another, the method comprising: detecting a value of an operating parameter for each of the diaphragm pumps, the operating parameter being one or more of liquid pressure at one or more points in the channel, power consumption, or flow rate; and controlling the parameter value of at least one operating parameter for each of the diaphragm pumps in a coordinated manner to prevent blood from entering the channel.

7. The method of claim 6, wherein the controlled parameter values of the diaphragm pumps lie at a ratio to one another that is dependent in a predetermined manner on the detected values of the liquid pressure in the channel.

8. The method of claim 6, wherein the controlled parameter values of the diaphragm pumps are variable over time in accordance with a fixed schema.

9. The method of claim 6, wherein the controlled parameter values of the diaphragm pumps are changed periodically over time following a start-up phase.

10. The method according to claim 6, wherein the controlled operating parameters of the diaphragm pumps are respective values of the liquid pressure generated by the diaphragm pumps, respective delivery capacities of the diaphragm pumps, or respective flow rates of the diaphragm pumps.

11. The method according to claim 6, comprising setting a fixed pressure difference between the diaphragm pumps.

12. The method according to claim 6, comprising setting a fixed flow rate difference between the diaphragm pumps.

13. The method according to claim 12, wherein setting a fixed flow rate difference between the diaphragm pumps comprises setting a fixed flow rate difference of less than 100 millilitres per day, less than 10 millilitres per day, or less than 1 millilitre per day.

14. The method according to claim 6, comprising: controlling the diaphragm pumps to reverse a direction of movement of the liquid.

Description

(1) The invention will be shown in figures of a drawing and described hereinafter on the basis of exemplary embodiments. In the figures

(2) FIG. 1 shows a hollow catheter having a driveable shaft and a supply device according to the invention,

(3) FIG. 2 shows the end of a hollow catheter in a longitudinal section with a distally fastened rotary pump for operation in a blood vessel.

(4) FIG. 3 shows a cross section through a hollow catheter,

(5) FIG. 4 shows a cross section through a further hollow catheter,

(6) FIG. 5 shows a method sequence for a method for operating a supply device,

(7) FIG. 6 shows a graph reproducing the course over time of flow rates in three different variants, and,

(8) FIG. 7 shows a graph reproducing the course over time of the liquid pressure.

(9) FIG. 1 shows a hollow catheter 1 in an illustration interrupted in the longitudinal direction, wherein an end 1a which is proximal in medical use is illustrated in the lower region and a distal end 1b is illustrated in the upper region. By way of example, an implantable blood pump may be provided at the distal end of the hollow catheter 1, especially for operation in a blood vessel and/or a heart chamber.

(10) A rotatably driveable shaft 2 extends within the hollow catheter 1. This serves for example to drive a blood pump and is connected at its proximal end 2a to a drive motor 3. The shaft 2 can be introduced in the region of a feedthrough 4 into a coupling housing 5, wherein the feedthrough 4 is configured in such a way that a medium is prevented from passing along the shaft into the coupling housing 5 or cut from the coupling housing 5 by a seal.

(11) However, the solution is also conceivable that the rotating drive movement is transmitted by means of a magnetic coupling through a closed wall of the coupling housing 5 in that a first magnet element 6 is magnetically coupled within the coupling housing to a second magnet element 7 fastened outside the coupling housing 5 on a shaft end connected to the motor 3. The shaft 2 then has an interruption between the motor 3 and the further extension thereof in the coupling housing 5, and the corresponding wall of the coupling housing 5 is formed continuously and without an opening. The magnet elements 6, 7 are illustrated in a dashed manner in FIG. 1 as an alternative.

(12) The driveshaft 2 is produced for example from litz wires, in particular in twisted or stranded form, or is formed as a helical spring or is formed in a combination of both variants by a core surrounded by a helical spring on the one hand so as to be able to transfer high rotational speeds in the region of a few thousand revolutions per minute and on the other hand so as to be flexible during this process.

(13) In order to cool such a shaft during operation on the one hand and on the other hand so as to reduce the friction by lubrication, a cooling and lubricating liquid is usually provided within the channel 8 formed in the hollow catheter 1 and is advantageously biocompatible. The liquid is fed to the coupling housing 5 via an inflow channel 9 and is transported along the channel 8. For this purpose, the inflow channel 9 is connected to a first pump, which is formed in the exemplary embodiment as a diaphragm pump 10. Diaphragm pumps in this context have the property of being controllable in a very reliable and reproducible manner in order to be able to control generated pressures and flow rates in an exact manner. The use of magnetically actuated diaphragm pumps has proven to be particularly advantageous in this context. A magnet device 10a is therefore illustrated in FIG. 1, which serves as a drive for the diaphragm pump 10, wherein the magnet device 10a is actuated by an electric control device 11.

(14) The diaphragm pump 10 sucks liquid from an inflow reservoir 12, as is illustrated by the arrow 13, and transports this at an adjustable flow rate and an adjustable pressure into the coupling housing 5 via the inflow channel 9. The liquid spreads in the coupling housing 5 and in particular moves in the direction of the arrow 14 along the channel 8 in the direction of the distal end 1b of the hollow catheter. The movement along the channel 8 can be assisted for example by the rotation of the shaft 2, when this has an at least partially helical outer contour and rotates in a suitable direction of rotation.

(15) Although the rotation of the shaft 2 may assist the movement of the liquid along the channel 8, it is possible in some exemplary embodiments to determine the contribution of the rotation of the shaft to the delivery capacity so as to thus adapt the delivery capacity of the pump(s). In other words, the delivery capacity provided on account of the rotation of the shaft is compensated for by an adaptation of the delivery capacity of the pumps. The determination of the delivery capacity on the basis of the rotation of the shaft may then also be interpreted as a disturbance variable, which is compensated for by the adaptation of the delivery capacity of the pumps in order to ensure a predetermined delivery capacity through the channel. The delivery capacity of the shaft 8 may be dependent inter alia on the rotational speed of the shaft, possibly wear on the shaft, the deflection of the catheter, or the like. Although these variables can be determined, a compensation of the resultant delivery capacity of the shaft by the pump is often easier.

(16) Flow rates in the range of microliters or millilitres per hour can usually be set by means of the actuation of the diaphragm pump 10.

(17) In order to be able to suitably control or regulate corresponding flow rates and/or pressures, at least one suitable sensor 15 is provided in the channel 8 and is connected by means of a communication line 16 to the control device 11. The sensor 15 may be formed for example as a pressure sensor, as a flow rate sensor, or as a combined sensor for detecting the pressure and the flow rate.

(18) In the shown exemplary embodiment the sensor 15 is assigned to the first diaphragm pump 10 and detects the pressure generated by this first pump and/or the corresponding flow rate.

(19) In accordance with the exemplary embodiment of FIG. 1 the channel 8 is divided in the longitudinal direction into a first channel region 8a, through which liquid flows in the direction from the coupling housing 5 to the distal end 1b of the hollow catheter 1 in the direction of the arrow 14, and a second channel region 8b, which is formed as a return channel. The two channel regions 8a, 8b are thus connected in series and together form the channel 8.

(20) The return channel 8b may be separated from the first channel region 8a for example by a partition wall 17, which is illustrated in FIG. 3, or the second channel region/return channel 8b may be formed by a cannula 18, which extends within the hollow catheter 1. This variant is illustrated in FIG. 4 in cross section.

(21) The return channel 8b is formed in accordance with FIG. 1 in such a way that it causes a return flow of the liquid into the coupling housing 5 and from there into a second diaphragm pump 19. The second diaphragm pump 19 may advantageously be formed as a magnetic diaphragm pump having a magnet device 19a which is actuated by the control device 11 and forms the drive of the diaphragm pump 19. The diaphragm pump 19 suctions the liquid from the return channel 8b and guides this via a discharge channel 20 into a discharge reservoir 21.

(22) The control device 11 is additionally connected to a second sensor 22, which, similarly to the first sensor 15, can be formed as a flow path sensor and/or as a pressure sensor and is assigned to the return channel 8b and therefore to the second diaphragm pump 19. By way of example, the flow rate of the return channel 8b or the suction pressure of the second diaphragm pump 19 can be detected by the second sensor 22. The parameters detected by the second sensor 22 are fed via a second communication line 23 to the control device 11.

(23) The control device 11 is in turn connected to an electric supply connection 11a, which supplies the control device with a low DC voltage (low voltage). The control device 11 generates pulses, which are fed to the magnet devices 10a, 19a in order to drive the first and second diaphragm pump 10, 19. The flow rates and/or pressures generated by the first and second diaphragm pump 10, 19 can be controlled by means of the frequency and the stroke of the pulses generated by the control device 11.

(24) FIG. 2, as an example for a use of a hollow catheter having a driveable shaft, shows an implantable blood pump 24, which is formed as a rotary pump having a rotor 25 with delivery elements. The rotor 25 is directly connected to the shaft 2, which is mounted in the housing 27 of the blood pump at the distal end of the rotor 25 in a rotary bearing 26. The blood pump 24 suctions blood via suction openings 28 at its distal end in the direction of the arrows 29, 30 and transports this externally past the hollow catheter 1 via an annular channel 32 formed by an outflow tube 31 into a blood vessel (not illustrated).

(25) The shaft 2 is mounted at the end of the hollow catheter 1 in a bushing bearing 33, which on the one hand is to allow high rotational speeds, and on the other hand is to be as tight as possible in order to prevent or to limit a liquid exchange along the shaft 2. In particular, blood is to be prevented from passing from the interior of the housing 27 of the blood pump 24 into the hollow catheter 1, i.e. into the channel 8.

(26) In FIG. 2 a partition wall 17 is illustrated in a dashed manner in order to indicate the separation between the first channel region 8a of the channel 8 and the second channel region/return channel 8b. The inflow of the liquid through the first channel region 8a in the direction of the arrow 34 to the distal end of the hollow catheter 1 and the return flow in the direction of the arrow 35 through the second region 8b of the channel 8 is thus made possible. The rotary shaft 2 can thus be supplied with the liquid along its entire length.

(27) In order to prevent the inflow of blood into the channel 8, an overpressure of the liquid in the interior of the hollow catheter 1, i.e. in the channel 8, can be set, which causes liquid to flow at a very slow flow rate from the channel 8 into the housing of the blood pump 24, as indicated by the arrows 36, 37. By way of example, an outflow rate of a few microliters or millilitres per day can be set here, which represents a difference between the feed rate in the first channel region 8a and the return rate in the return channel 8b. This difference can be set and measured as the difference of the delivery rates between the first pump 10 and the second pump 19.

(28) A flow diagram for a method for operating the shown supply device is illustrated in FIG. 5. In a first step 38 a venting of the channel 8 inclusive of the coupling housing 5 is performed in that liquid is fed by means of the first pump 10. Following the venting of the channel 8 and of the pumps, of which the speed is adjustable, the direction of movement (forwards/backwards) in which the liquid is to be moved through the channel 8 is determined in a second step 39. The diaphragm pumps 10, 19 and the reservoirs 12, 21 can allow both directions of movement of the liquid. The pressures generated by the diaphragm pumps 10, 19 are set depending on the direction of movement of the liquid.

(29) In a third step 40 is it decided whether the capacities (performances) of the pumps are to be set manually. If the pumps are to be manually set, the rest of the process proceeds via the path 40a, and in a step 46 the pressures and/or flow rates of the two pumps are set. This variant is usually selected when the flushing rate, i.e. the flow rate through the channel 8, is to be small and constant.

(30) If a manual actuation is to be selected, the continued path proceeds via the arrow 40b, and the automatic actuation of the pumps is started in a fourth step 41. For this purpose the pressure is firstly detected at the two pressure sensors 15, 22 in the step 41, a pressure difference is calculated from this, and from this the actuation of the pumps 10, 19 is calculated in a fifth step 42 by corresponding pulses of the control device 11. Here, the sought pressure difference may also be variable over time, for example varying periodically.

(31) In a sixth step 43 the generated pressure difference is compared with the target pressure difference. If the actual pressure difference corresponds to the target pressure difference, the pressure difference or a flushing rate calculated herefrom for example is thus indicated in a seventh step 44 and the method is ended in an eighth step 45. The ending of the method means that the supply device is in a stable operating state and the pumps 10, 19 are actuated and operate accordingly. If it is determined in the sixth step 43 that the actual pressure difference does not correspond to the target pressure difference, the method jumps back via the path 43a to the fourth step 41, in which the pressure difference is measured, and from this the new actuation of the pumps is determined in a regulation step.

(32) Instead of the pressure measurements and corresponding pressure regulation of the differential pressure, the flow rate may also be measured and a corresponding flow rate difference can be set as control variable.

(33) A typical course over time of flow rates in three exemplary variants is shown in FIG. 6. The flow rate is indicated in volume per time on the y-axis of the graph, whereas the time is plotted on the x-axis. A first curve 48 by way of example shows the flow rate, measured by the sensor 15 or the sensor 22, wherein the flow rate is constant over a large portion of the time, but the flow rate is changed from time to time, for example every twenty seconds or in each case after a few minutes, by a temporary increase 49, 50 of the flow rate. There is thus no stationary flow formed in the channel 8, which might leave certain regions of the channel untouched as what are known as deadwater regions, such that liquid located there does not move on further. A change to the flow rate generates turbulence and non-stationary flow conditions, which then also determine the deadwater areas and exchange the liquid there.

(34) A further object of a corresponding control of the flow rate is to prevent particles located in the liquid and which for example are created by abrasion of the rotating shaft 2 from being moved on further where possible, such that these do not exit through the bearing 33, illustrated in FIG. 2, in the region of the blood pump and cannot enter into the body of a patient.

(35) If the flow rates detected by the two sensors 15, 22 are plotted in the same graph, an increased flow rate with a particularly noticeable difference of the flow rates can be set for example particularly in the regions 49, 50, which indicates that some of the liquid in these regions 49, 50 exits intermittently in very small quantities from the channel 8 and passes into the interior of the pump housing of the blood pump and thus flushes away any quantities of the blood deposited there from the bearing 33.

(36) In a second variant 51 of the course of the flow rate, this is varied periodically around a constant course 52, for example in the form of a sine curve. A constantly changing flow with flow conditions that likewise change constantly and guarantee a liquid exchange in all regions of the channel 8 is thus provided.

(37) In the third variant, which is illustrated in the curve 53, apart from temporary periodic increases 54 of the flow rate, the flow direction is also reversed, shown on the basis of the example of the reduction 55 of the flow rate. The reversal of the flow causes a change to the flow direction of the liquid in the channel 8 and therefore likewise the exchange of liquid in deadwater areas. Such a reversal of the flow direction can occur for example at intervals from five to ten minutes.

(38) In FIG. 7 measured pressure values are plotted on the y-axis over time t, wherein a first curve 56 indicates the pressure in the region of the sensor 15 and a second curve 57 indicates the pressure in the region of the sensor 22. It can be seen that in two regions 58, 59 the pressure is temporarily increased by the first diaphragm pump 10, whereas the pressure in the region of the return line, detected by the sensor 22, remains constant. This causes liquid to flow off into the pump housing through the bearing 33 in the regions of the increased pressure 58, 59, thus relieving the pressure in the channel 8.

(39) As a result of the above-described invention, a supply device in the form of a flushing device for a hollow catheter for a blood pump is provided, in which few wear parts are used and therefore stable operation with low liquid losses thus can be ensured over a long period of time.