Patent Application
20130071271
1. Field of the Invention
Embodiments of the present invention relate to a valveless pump design, and more specifically to a pulsatile, non-contact, positive displacement, finger-pump for pumping one or more fluids.
2. Background of the Invention
Pump designs exist for a variety of applications. Regardless of design, however, pumps are fundamentally of two types: direct and indirect (or, “non-contact”). In a direct pump, portions of the pump are in direct contact with the fluid. An automotive water pump, for example, uses a motor or belt-driven impeller to pump cooling fluid throughout the engine. The impeller is connected to the drive (e.g., a pulley or gear) and is separated from the cooling fluid by a seal and a housing. In this configuration, the impeller is in direct contact with the cooling fluid, while the drive pulley or gear is outside the cooling system.
In some applications, however, it is desirable to provide a pump capable of pumping a fluid without contacting the fluid directly. This can be to prevent, for example, contamination or clotting of the fluid (e.g., in medical applications). This may also be desirable to prevent contamination of the pump (e.g., when pumping radioactive cooling fluid in a nuclear reactor, for example). An indirect pump may also be desirable when pumping particularly dirty fluids, such as bilge water, for example, to prevent fouling and/or clogging of the pump.
One application for indirect pumps is in the pumping of fluids for hemodialysis (hereinafter, “dialysis”). End Stage Renal Disease (ESRD) is a disease afflicting hundreds of thousands of patients worldwide. Most patients diagnosed with ESRD undergo dialysis.1 Traditional dialysis generally requires patients to visit a clinic three times a week for three to five hour treatment sessions.2 This results in the patients disrupting their everyday life and restricting activities such as extended travel. In addition, because the treatments are generally administered every other day, high waste levels accumulate, which can make patients experience dizziness and lethargy, amongst other symptoms.3 Limitations in the current technology have spurred the development of portable renal replacement systems. A portable renal system would both free the patients from time required at the clinic but would also allow more frequent treatments, leading to lower average waste levels.4 1A. A. P. B. Kerr, J. L. Flavier, B. Canaud, C. M. Mion, “Comparison of dialysis and hemodiafiltration: A long-term longitudinal study,” Kidney International, vol. 41, pp. 1035-1040, 1992.2J. Daugirdas, “Second generation logarithmic estimates of single-pool variable volume Kt/V: an analysis of error,” Journal of the American Society of Nephrology, vol. 4, pp. 1205-1213, 1993; I. T. Kjellstrand C M, “Daily dialysis: history and revival of a superior dialysis method,” ASAIO Journal, vol. 44, pp. 115-122, 1998 May/June; B. F. Piccoli G. B., Iacuzzo C., Anania P, Iadarola A. M., Mezza E., “Why our patients like daily dialysis,” Dialysis International, vol. 4, pp. 47-50, 2000.3F. Mastrangelo, et al., “Dialysis with increased frequency of sessions (Lecce dialysis),” Nephrol Dial Transplant, vol. 13 Suppl 6, pp. 139-47, 1998.4U. B. Francesco Locatelli, Bernard Canaud, Hans Köhler, Thierry Petitclerc, Pietro Zucchelli, “Dialysis dose and frequency,” Nephrology Dialysis Transplantation, vol. 20, pp. 285-296, 2005; L. A. F. Mastrangelo, M. Napoli, V. DeBlasi, F. Russo and P. Patruno, “Dialysis with increased frequency of sessions (Lecce dialysis),” Nephrol Dial Transplant, vol. 13, pp. 139-147, 1998.
In the case of dialysis system design, maintaining a sterile path for the blood flow is a critical requirement. Any pump used in the system must not introduce harmful agents (e.g., bacteria or viruses) to the blood or the dialysate. Conventional dialysis systems typically use peristaltic roller pumps, a type of rotary, positive displacement pump, because they can maintain a sterile flow path for the blood. These pumps use moving rollers to squeeze a flexible tube forcing the fluid inside to move in one direction. When the roller has passed a section of tubing, the tubing rebounds, or “reinflates,” and enables fluid to flow into the void thus created. Because the flexible tubing is the only component exposed to the fluid, the pump can maintain sterility.
These types of pumps rely on relatively stiff tubing, however, to displace fluid. As a result, these pumps exhibit poor efficiency primarily due to the energy required to compress the (not particularly) flexible tubing. In addition, during dialysis, two pumps are generally used, one for circulating blood and the other for circulating the dialysate. Additional pumps can be used is additional drugs or blood products, for example, are to be administered. Pumps of this type that are capable of producing the necessary flow rates for efficient dialysis tend to be too large and heavy to be used as part of a portable renal replacement system.
What is needed, therefore, is a small, energy efficient pump capable of pumping one or more fluids. The pump should be an indirect pump such that none of the pump components are in direct contact with the pumped fluid. The pump should be reliable, serviceable, and economical to produce and maintain. It is to such a pump that embodiments of the present invention are primarily directed.
Embodiments of the present invention relate to a valveless pump design, and more specifically to a pulsatile, non-contact, positive displacement, finger-pump for pumping one or more fluids. In its most basic form, the pump can comprise one or more sets of fingers sequentially activated by a drive mechanism. Each finger can have a first, or open, position and a second, or closed position. When a finger is in this first position, the length of tubing proximate that finger can be open and uncompressed. When a finger is in the second position, the length of tubing proximate that finger can be partially or completely compressed.
The fingers can be activated, or moved, by a drive mechanism in a sequential manner. The fingers can act sequentially along the length of tubing to create positive displacement of the fluid inside the tubing. As the fingers are sequentially activated, they alternately compress and release the tubing to create positive displacement of the fluid in one direction. In some embodiments, in the closed position, the finger can fully compress the tube such that no fluid will flow (forward or backward) through that particular portion of tubing. In a preferred embodiment, at least one finger in the set of fingers is in the closed position at all times to substantially prevent backflow in the tubing.
Embodiments of the present invention can comprise a pump system for pumping one or more fluids comprising a housing, which can include a first tubing holder for holding a length of flexible tubing and a first axle disposed in the housing. The pump system can further comprise a plurality of fingers, each finger comprising an axle hole located on a distal end of each finger, wherein each finger is pivotally coupled to the first axle. The pump system can further comprise a drive system configured to move each of the plurality of fingers sequentially from a first position to a second position. In some embodiments, each of the fingers can at least partially occlude a portion of the flexible tubing in the second position. In a preferred embodiment, each of the fingers substantially completely occludes the portion of the flexible tubing in the second position.
In some embodiments, the tubing holder can comprise a channel and a plane surface and each of the plurality of fingers can compress the flexible tubing between the finger and the plane surface when the finger is in the second position. In some embodiments, the plane surface can further comprise a resilient cover to prevent over compression of the flexible tubing in the second position.
In some embodiments, the drive system can comprise a camshaft including a central shaft and a plurality of cam lobes, and a motor for driving the camshaft in a rotary motion. The plurality of cam lobes can be a modified eccentric design. In some embodiments, the motor can drive the camshaft directly. In other embodiments, the motor can drive the camshaft with a belt. In still other embodiments, the camshaft can further comprise a driven gear mounted to the central shaft and the motor can comprise a drive gear mounted to the motor and in communication with the driven gear to drive the camshaft in a rotary motion. The driven gear can also comprise a drive gear system comprising two or more gears in communication with each other and with the driven gear to drive the camshaft in a rotary motion.
Embodiments of the present invention can also comprise a pump system wherein the drive system comprises a plurality of shape memory alloy (SMA) wires. In this configuration, each of the plurality of fingers can comprise a first SMA wire coupled to the finger above the axle hole and a second SMA wire coupled to the finger below the axle hole. In this configuration, the first and second SMA wires can be individually activated to move the finger back and forth between the first position and the second position. First instance, in some embodiments, the first SMA wire can move the finger from the first position to the second position and the second SMA wire can move the finger from the second position to the first position. In a preferred embodiment, at least one of the plurality of fingers is in the second position at any given time to prevent backflow in the flexible tubing.
Embodiments of the present invention can further comprise a system for pumping two or more fluids comprising a housing, with a first side, a middle, and a second side. The housing can further comprise a first tubing holder for holding a first flexible tube, a second tubing holder for holding a second flexible tube, a first axle disposed on a first side of the housing, and a second axle disposed on the second side of the housing. The system can include a first set of two or more fingers, while each finger can comprise an axle hole, and can be pivotally coupled to the first axle and a second set of two or more fingers, again where each finger can comprise an axle hole, and can be pivotally coupled to the second axle.
The system can include a camshaft, which can comprise a central shaft and a plurality of cam lobes, and can be disposed in the center of the housing. In a preferred embodiment, the camshaft sequentially moves the first set of fingers and the second set of fingers from a first position to a second position and the fingers at least partially occludes a portion of the first flexible tubing in the second position. The sequential occlusion of the first and second tubes by the fingers can pump fluid through the tubes. In some embodiments, each cam lobe can have a duration angle of between approximately 30 degrees and 180 degrees of rotation.
The system can also include a motor for driving the camshaft in a rotary motion. In a preferred embodiment, the motor is an electric motor.
In some embodiments, the flexible tube can comprise inlet connectors and outlet connectors for connecting the tubes to external tubing. The inlet connector can comprise an inlet port that can be disposed proximate a top portion of the inlet connector and the outlet connector can comprise an outlet port disposed proximate a bottom portion of the outlet connector. In some embodiments, the inlet connector can also comprise a steeper change in cross-sectional area from a first side to a second side than the outlet connector.
In some embodiments, each of the fingers can comprise a first shoulder on a first side of the axle hole and a second shoulder on the second side of the axle hole to reduce friction. the fingers can also be separated by one or more washers disposed between each of the fingers to reduce friction.
Embodiments of the present invention can also include a method of manufacturing a pump. The method can comprise providing a housing with a first end, a second end, a first side, a middle, a second side, and one or more tubing holders; mounting a first set of one or more fingers on a first axle; mounting the first axle proximate the first side of the housing; mounting a camshaft, comprising one or more cam lobes, in the middle of the housing proximate the first set of fingers; and mounting a motor proximate the camshaft to rotate the camshaft. When mounted each of the one or more cam lobes can be disposed proximate one of the first set of one or more fingers such that, as the cam rotates, each cam lobe can activate a respective finger. Activating can comprise, for example, moving the respective finger from a first, open position to a second closed position.
In some embodiments, the method can further comprise mounting a second set of one or more fingers on a second axle and mounting the second axle proximate the second side of the housing. Again, each of the one or more cam lobes can be disposed proximate one of the first set of one or more fingers and one of the second set of one or more fingers such that, as the cam rotates, each cam lobes can activate a respective finger from the first set and a respective finger from the second set. In a preferred embodiment, the cam lobes are disposed on the camshaft such that the cam lobes activate the first set, the second set, or both sets of one or more fingers sequentially.
In some embodiments, the method can further comprise affixing a drive gear to the motor and affixing a driven gear to the camshaft. In this configuration, the drive gear rotates the driven gear. in other embodiments, the motor can be directly attached to the camshaft. In still other embodiments, the motor and the driveshaft can be coupled using a drive belt.
In some embodiments, the housing can further comprise a first adjustment means to enable the distance between the tubing holder and the camshaft to be adjusted. In other embodiments, the housing can further comprise a second adjustment means to enable the height of the camshaft to be adjusted. In still other embodiments, the first and second sides of the housing can be coupled to the housing with hinges that can enable insertion of tubing in the tubing holder. The hinges can be independently adjustable to enable different sizes of tubing to be inserted in the tubing holders.
These and other objects, features and advantages of the present invention will become more apparent upon reading the following specification in conjunction with the accompanying drawing figures.
a depicts a side, perspective view of a finger, in accordance with some embodiments of the present invention.
b depicts an end view of a finger, in accordance with some embodiments of the present invention.
Embodiments of the present invention relate to a valveless pump design, and more specifically to a pulsatile, non-contact, positive displacement, finger-pump for pumping one or more fluids. In its most basic form, the pump can comprise one or more sets of fingers sequentially activated by a drive mechanism. Each finger can have a first, or open, position and a second, or closed position. The fingers can be activated in a sequential manner to create a pumping action through flexible tubing.
To simplify and clarify explanation, the system is described below as a pump system for use with a portable renal replacement system. One skilled in the art will recognize, however, that the invention is not so limited. The system can be used to in a variety of fluid pumping applications, especially those applications in which the fluid is preferably pumped without direct contact with the fluid. In other applications, the pump could be used for pumping non-fluids, such as for example, slush, sewage, or in food processing applications. These applications can include, for example and not limitation, medical, nuclear, and sewage applications. In addition, the system is described below for use in pumping two fluids simultaneously. Obviously, the system can be adapted to pump one fluid or more than two fluids without departing from the spirit of the invention.
The materials described hereinafter as making up the various elements of the present invention are intended to be illustrative and not restrictive. Many suitable materials that would perform the same or a similar function as the materials described herein are intended to be embraced within the scope of the invention. Such other materials not described herein can include, but are not limited to, materials that are developed after the time of the development of the invention.
As described above, a problem with conventional fluid pumps is that pumps with higher efficiency tend to be direct fluid pumps, i.e., the impeller, or other pump components, are in direct contact with the fluid being pumped (or, “pumped fluid”). Indirect pumps, like the roller pumps generally used for dialysis, on the other hand, tend to be heavy and inefficient. These pumps tend to rely on stiff-walled tubing, which requires larger, more powerful motors and increases frictional losses. What is needed is an energy efficient, compact, indirect pump capable of pumping one or more fluids that is compatible for use with both flexible and stiff-walled tubing. It is to such a pump system that embodiments of the present invention are primarily directed.
Embodiments of the present invention can comprise a non-contact fluid pump design that can minimize total pump size while maximizing pump efficiency. As shown in
In some embodiments, the plurality of fingers 130 can be positioned between the cam lobes 115b and the tubes 150. The fingers 130 can be mounted on shafts 145 located on the distal end of fingers 130 enabling each finger to pivot between a first, open position and a second, closed position. As each finger 130 is actuated by its respective cam lobe 115b, the finger 130 pivots between the first position, in which the tube 150 is open and flowing and the second position, in which the tube 150 is partially or completely compressed. In some embodiments, there can be a gap between each pair of fingers 130 to maximize pumping efficiency along the tube 150 and reduce friction by preventing direct contact between the fingers 130 themselves. In a preferred embodiment, a low friction washer is positioned between each finger 130 to further minimize friction.
In some embodiments, a motor 155 can be used to drive the camshaft 115. The motor 155 can be, for example and not limitation, an electric, pneumatic, or hydraulic motor. In some embodiments, the motor 155 can be positioned under the camshaft 115, though other configurations are contemplated. In some embodiments, a drive gear on the motor shaft can be used to drive a driven gear on the camshaft 115 (gears not shown for clarity). In other embodiments, the motor 155 can be configured to drive the camshaft 115, for example and not limitation, directly or with a chain or belt.
As shown in
As one would expect, each cam lobe 115b has a different orientation on the shaft 115a. In a preferred embodiment, the cam lobes 115b are positioned such that they sequentially compress the tubing 150 to create a pumping action.
The eccentric cam design has the advantage of enabling control of tube 150 occlusion time when compared to, for example, a conventional eccentric cam design that provides only one instance of occlusion (i.e., a the top of the lobe). In an exemplary embodiment, the pump can utilize seven cam lobes 115b that each actuate a finger 130 on each side of the housing 120 (i.e., each cam lobe 115b actuates two fingers 130). In this embodiment, the five middle cam lobes 115b can have 60 degrees of duration, the first cam lobe 115b can have 90 degrees of duration, and the last cam lobe 115b can have 180 degrees of duration. This configuration provides extra time for the tube 150 to refill after compression and minimizes backflow. Of course, other camshaft 115 configurations can be used for different pumping requirements (e.g., more even cam lobe 115b duration could provide a substantially constant flow rate). In some embodiments, for example, the cam lobe 115b configuration can be changed to provide a pulsatile flow similar to the natural flow of blood in the body caused by the heart.
In some embodiments, as shown in
In this configuration, a first end of both wires 315 can be anchored to an anchor point on the housing 305. The second end of a first wire 340a can then be affixed to a finger 330 below the pivot shaft 335, while the second end of a second wire 340b can be affixed to the finger above the pivot shaft 335. The wires 340a, 340b can then be activated using an electrical current or heat, as applicable, which causes the wires 340a, 340b to retract.
In use, one of the wires 340a, 340b can be activated to pull a finger 330 into the first position, while the other wire 340a, 340b can be activated to pull the finger 330 into the second position. The mounting direction, i.e., which side of the housing 120 anchors the wires 340a, 340b, determines which wire 340a, 340b pulls the finger 330 closed or open. In a preferred embodiment, the wires 340a, 340b are anchored to the far side of the housing to maximize their length, thus maximizing their retraction.
As shown in more detail in
In some embodiments, each finger 130 can further comprise a rounded back profile 415. The rounded profile 415 can reduce the contact area between the finger 130 and the cam lobes 115b reducing friction. The fingers 130 can further comprise a longitudinal ridge 420. The ridge 420 can reduces the contact area between the finger 130 and tube 150, which can also reduce friction. Reduced contact area between the finger 130 and the tube 150 can reduce the area of the tube 150 that is completely occluded when the finger 130 is in the second, or closed, position. The use of a small occlusion area does not affect the flow rate provided the tube 150 is completely occluded by the contact area 420 and relatively little fluid remains between the contact areas 420 of each finger 130 during use. In addition, when used for pumping blood, for example, the small occlusion area tends to reduce blood cell damage.
As shown in
In some embodiments, therefore, an adjustment means 510 can be provided. The adjustment means can comprise an adjustable mounting means 510 for the tube channel 120a/plane surface 120b (collectively, tube holder 505). In other words, the tube holder 505 can be mounted to the pump housing 120 in such a way that its distance from the camshaft 115 can be adjusted (e.g., using slotted mounting holes 510). In other embodiments, the pump system 100 can comprise a resilient cover 515 (e.g., a rubber plate) disposed over the plane surface 120b. The resilient cover 515 can enable the tube 150 to be completely occluded, while absorbing the over compression and reducing energy consumption. The adjustment means 510 can enable individual adjustment to facilitate the use of different sizes of tubing 150 at the same time (e.g., large tubing on one side and small tubing on the other).
In some embodiments, the position of the camshaft 115 can be adjusted to provide complete occlusion. In other words, the plane surfaces 120b can be angled such that changing the height of the camshaft 115 changes the distance between the camshaft 115 and the plane surfaces 120b. In some embodiments, therefore, the camshaft 115 can be provided with an adjustment slot 520 to enable the height of the camshaft to be adjusted. In this manner, provided the tubes 150 are relatively the same size, larger tubes 150 can be accommodated by raising the camshaft 115 and smaller tubes 150 can be accommodated by lowering the camshaft 115. This assumes, of course, that the plane surfaces are angled outward from bottom to top, but the reverse is also contemplated.
In still other embodiments, the plane surfaces 120b can be hinged where they attach to the housing 120. In this configuration, the plane surfaces 120b can be pivoted away from the camshaft 115 to facilitate insertion of the tubing 150. In some embodiments, the hinges can have multiple positions to enable the distance between the plane surfaces 120b and the camshaft 115 to be adjusted. Again, this provides the ability to accommodate different sizes of tubing. The hinges can be independently adjustable to facilitate the use of different sizes of tubing 150 at the same time (e.g., large tubing on one side and small tubing on the other).
The pump system 100 is compatible with a variety of tubing types. In a preferred embodiment, thin-walled, flexible tubing (e.g., Penrose drain tubing) is used to reduce, among other things, the compression forces, pump size, and energy consumption. This flexibility is generally only required, however, for the portion of the tube 150 that is inserted into the pump 100. For the other portions of the fluid circuit, using more durable tubing, such as the tubing commonly used for IVs, may be a better choice. The use of highly flexible tubing throughout the fluid circuit can cause, for example, excessive flow rate variation. To enable modularity and flexibility, tubing connectors can be used to connect the flexible tubing used at the pump to, for example, standard IV tubing.
As shown in
In some embodiments, the present invention can comprise a method for efficiently determining pump parameters given a set of performance requirements (e.g., fluid volume per unit of time). The performance of a finger pump can be controlled using several design parameters, such as, for example and not limitation, the characteristics of the tube and the actuator, which can also affect the size of the pump. And, while some of these design parameters have simple relationships when taken separately, many parameters are inter-related, making it difficult to predict how the pump will perform for a given configuration and set of actuator characteristics.
An understanding of the relationship of these parameters for the creation of an analytical pump model can nonetheless enable the systematic pump minimization while providing the desired performance. Required parameters can represent, for example, experimental settings, tube and actuator characteristics, design requirements, power consumption, and pump size information. Several of those relationships can be embedded in an analytical pump model and are discussed below.
The flow rate can be calculated by multiplying the effective cross sectional area of the tube with finger width and the number of strokes. When the tube is inserted into the pump, its cross section tends to become a long oval that fits a right trapezoid space between the fingers and the housing. The oval area can be calculated by multiplying the shape factor n/4 with tube width and squeeze distance as the oval area that fits in a rectangle. This is not, however, the only factor that determines the effective cross sectional area as peristaltic pumps are subject to head change, back flow loss, and friction loss, among other things. To consider these factors, a lumped constant called the oval constant, can be used to account for pump head changes, back flow loss, friction loss, and the shape factor shown below in Eq. 1.
effective tube cross section(cm2)=oval constant×tube width(cm)×squeeze distance(cm) (1)
The oval constant can be found by matching the model prediction with the experimental data (e.g., rpm vs. flow rate). If the pump is operating with high outlet pressure, for example, the flow rate will decrease and the oval constant will be smaller than the default value.
The efficiency of the pump can be calculated by dividing the fluid power by the brake power to pump the fluid, as shown in Eq. 2.
Fluid power refers to the theoretically calculated power required to transport the fluid at a given flow rate with a given flow pressure, as shown in Eq. 3.
fluid power(W)=flow rate×flow pressure (3)
For use in a dialysis application, the flow pressure can be assumed to be approximately equal to “normal” blood pressure (e.g., approximately 100 mmHg). This enables Eq. 3 to be simplified as shown in Eq. 4.
Finally, brake power refers to the power required to operate the pump. Brake power can be calculated, for example, by multiplying the required battery voltage and current.
The size of the pump, especially one for portable use can be an important calculation for the evaluation of a design. The overall volume of the pump, for example, can be calculated by multiplying the pump depth, width, and height. Pump depth, for example, can be defined as the length of the pump in the flow direction, and is related to the width and number of fingers that force the flow in the flow direction, as shown in Eq. 5.
pump depth(cm)=no. of fingers×finger width(cm)+α (5)
Where α can account for the additional length due to the housing of the pump. This information can provide useful values, for example, after designing an actual pump with a given setting.
The model can include one or more of the parameters provided below in Table 1. Each parameter can be designated as either an input, intermediate calculation, or output. Input parameters include, among other things, the type of tube, requirements, and actuator characteristics (shown in normal fonts). Output parameters, on the other hand, can provide information related to the pump design specifications (shown in bold). Finally, intermediate parameters can be used to calculate the output values from input values (shown in italics).
Effective tube cross section
fluid power (W)
Required flow velocity (cm/s)
finger width (cm)
Flow amount of a finger stroke (ml)
No. of cycles per minute (/min)
No. of finger strokes (/min)
signal delay (s)
Required force generation per
finger (N)
Total required force generation
(N)
Current (A)
Ampere hour (Ah)
Watt hour (Wh)
Battery size constant (cm
3/Wh)
brake power (W)
Efficiency
pump width (cm)
pump height (cm)
pump depth (cm)
pump body volume (cm
3)
Battery volume (cm
3)
Total volume (cm
3)
Based on the parameters' relationships, a finger pump model can be created in a simulation model such as, for example and not limitation, Simulink®. The pump model can include quantifiable design parameters that affect the design of the pump. The model can, in turn, provide quantifiable design output values including, but not limited to, the expected power consumption and the overall size of the pump. In addition, the model can be adapted for different flow circuits and different design requirements such as, for example and not limitation, target flow rates.
For efficiency, the pump model can utilize one or more assumptions. For example, the compression sequence of the fingers can be changed by varying the cam design for each finger. For simplicity, however, the model assumes that no such difference exists. In use, as long as a suitable average value is used, the results of the model are sufficiently accurate.
The model also assumes that the motor fits in the housing in the space under the camshaft. In other words, if the depth of the pump is, in reality, shorter than the motor, a portion of the motor will extend out of the housing. As a result, the size change according to the change in the force requirement for an actuator is not embedded in the model, though this could be added by, for example, comparing the trend of the motor size change due to the change in the force requirement.
The relationship between the force generation and power consumption is not embedded in the model because this is typically motor dependent. This could nonetheless be added for a known motor or for a series of motors. Without this information, suitably accurate information can nonetheless be provided by the model using average resistance per power output to predict power consumption. In addition, the force required to squeeze a tube does not increase linearly when the compression area increases, most notably when the compression area gets smaller. The model nonetheless assumes a linear force-compression relationship for simplicity. Thus, a finger width of greater than 0.4 cm, for example, is preferable.
The creation of analytical pump models with limited number of design parameters with oval constant allowed an exploration of all the different parameters associated with pump design. Modeling suggests that increasing squeeze distance, number of fingers, and applied voltage while using a flexible tube reduces the size of the pump. The model is validated experimentally in section 5, below.
To verify the accuracy of the pump model discussed above, the model was tested while varying pump model parameters. The pump model can be validated using the pump system 100, as described, with water experiments. Blood flow results can prove the validity of the pump system 100 for use in, for example and not limitation, a portable renal replacement system.
To test the effect of the cross-sectional area, one Penrose drain tube was replaced by two small tubes disposed on the same side of the pump (i.e., at different heights in the same tube holder 505). Water was used as the pumped fluid.
Although the two tubes were the same size, due to their location in the tubing channel 150 (i.e., one mounted higher than the other) and the geometry of the pump, the top tube cross-sectional area was approximately 1.586 times bigger than the bottom tube. As shown in Table 2, below, when 10.8V was applied to run the pump at approximately 38 rpm, the flow rate difference closely matched the area ratio indicating that the cross-sectional area increases the flow rate linearly.
Because flexible tubes are used, a supply side head is preferably provided to fill the tube (i.e., rather than having the fingers open the tube). In addition, it is preferable for the pump to be able to generate stable flow rates regardless of the level of fluid in the supply source (e.g., a dialysate bag). Experiments were conducted with water with a dialyzer in the circuit to provide the same setting as a portable renal replacement system.
Four different water levels were tested by starting from a full reservoir (about 550 ml) and running the pump for 1 minute each time. Resulting flow rates were 132, 127, 125, and 125 ml/min, respectively. The result validated that in the given system configuration, the head difference in the supply source had a negligible effect on flow rate. In addition, the same flow rate is expected for both single pass mode and recirculation mode.
The pump 100 can pump two different fluids at the same time. For a portable renal replacement system, for example, the two fluids can be blood and dialysate. These fluids may have two different viscosities, however; thus, the effect of viscosity on flow rate was also tested. To provide a variety of fluid viscosities, glycerol mixtures were prepared with different water-glycerin ratios. Table 3 shows the flow rates for four separate experiments for each viscosity.
As shown in the table, the flow rate was virtually constant for different viscosities. Based on the experimental data, a difference in viscosity of up to 5.5 mPa has little or no effect on the flow rate. Based on this data, it follows that the flow rate will be the same for both the blood (1.84 mPa) and dialysate, which is mostly water (1.005 mPa). It also shows that the oval constant can be kept the same as long as the fluid viscosity does not vary too widely (i.e., at least within a 5.5. mPa range).
The effect of the outlet, or disposal side, head change on the flow rate was tested using in vitro blood experiments. The blood flow rate was tested while both dialysate and blood were running through a dialyzer. Four different disposal side heads (24 cm, 33 cm, 42 cm, and 50 cm) were tested with different motor speeds (27, 34, and 42 RPM) for each configuration. The test results for the 50 cm head at 27 rpm are not provided because the fluid was not able to climb up the height with the slow flow rate. The results are plotted in
As shown, a lower disposal side head yielded higher flow rates than a higher head. In addition, flow rates increased with increased rpm. This tends to indicate that the disposal side head should be considered when designing a pump for use with, for example and not limitation, a portable renal replacement system configuration.
In addition, these results indicate that the head effect has a considerable effect on the oval constant. Since the oval constant can be determined according to the flow rate change when all the other design variables are kept the same, it shows that the oval constant for 24 cm height, for example, will be about 4 times larger than that for 50 cm (e.g., 90 ml/min vs. 22 ml/min). The flow results also can be sensitive to whether or not the tubes are completely occluded during operation. This indicates that manufacturing tolerances should be controlled, or, as mentioned above, means should be provided for adjustment of, or flexibility in, the tube holders 505 to ensure complete tube occlusion. It should be noted, however that although the flow rate is affected by complete occlusion, the system nonetheless yields very repetitive flow rates at a given setting. In other words, the flow rate is repeatable whether the tube is completely occluded or not.
Pump 1 was used for in vitro blood experiments to prove that it can be used for dialysis treatments. The pump was used to dialyze blood and the results were compared to known dialysis models that have been proven to match well with both in vitro experimental data and published patient data.6 6J. C. Olson, “Design and modeling of a portable hemodialysis system,” Master of Science in Mechanical Engineering, Mechanical Engineering, Georgia Institute of Technology, Atlanta, 2009.
Porcine blood was obtained during desanguination at a local abattoir, and Citrate and Heparin were added to the blood immediately after collection at a ratio of 100:10:890 (35% Citrate solution: Heparin: blood). The porcine blood was then transported to the laboratory in a thermally insulated container. Dialysate was prepared by mixing concentrates with deionized water at a ratio of 22:38:940 (acid concentrate: bicarbonate concentrate: deionized water).
The individual solute clearance, KD, is the typical measure of clearance of an individual solute. Manufacturers provide a value of KDA, the clearance multiplied by the surface area, in the specifications for their dialyzers. A dialyzer with a small KoA removes waste more slowly than the one with a larger KoA. Thus, KoA is an important dialyzer characteristic for the prediction of the waste level over time. KoA is different, however, for each solute for different flow rates.
Unfortunately, a typical manufacturer reports only KoA for the flow rate around 500 ml/min. This flow rate is consistent with conventional non-portable dialysis machines designed to treat dialysis patients quickly at an outpatient dialysis clinic. The use of a constant, portable dialysis system, on the other hand requires much lower flow rates. As a result, a single pass mode experiment was conducted first to find the KoA value. 500 ml of blood was used with the blood flow rate of 105 ml/min and dialysate flow rate of 100 ml/min. Creatinine levels were measured over time with an iSTAT analyzer (Abbott). The data points for both single pass mode and recirculation mode are plotted in
A recirculation mode experiment was performed with 450 ml of blood, with a flow rate of 74 ml/min, and 550 ml of dialysate, with a flow rate of 75 ml/min. Again, the data points matched the prediction values closely with the same KoA value found in the single pass mode, demonstrating that the pump works also well for the recirculation mode.
Since the concept of using a finger pump for a portable renal replacement system is demonstrated, the effect of rpm change on the flow rate was tested to find the oval constant that represents the experimental setting. Pump 2 (a second prototype pump) was used, which employed a stronger, faster motor. The experiment results are plotted in
The experimental data was then used to calibrate the model for the predefined experimental settings. As explained above, the oval constant is used as a lumped constant to account for head changes and/or back flow loss. For a given pump design, however, the oval constant changes substantially linearly with the flow rate change. Thus, the oval constant can be determined by finding the value at which the model prediction closely matches experiment data. Since the main required flow rate for this pump is about 100 ml/min, the oval constant that closely matches that region is found.
As with Pump 1, the effect of outlet, or disposal side, head change on the flow rate was tested with pump 2, and the results are plotted in
With the oval constant found using pump 2, the pump model was used to explore the design space and find an optimized pump design for a portable renal replacement system. The problem is formulated as below.
According to the problem formulated above, the design space was explored in the region where the input variables ranged as shown in Table 4.
The maximum squeeze distance was chosen as 1.5 cm because squeeze distances larger than this tend to cause the flexible tube to shift out of position. The minimum voltage value was chosen to ensure that the motor reaches desired minimum operating speed, and the maximum value is equivalent to the motors nominal voltage. The increment of the voltage was chosen based on the per cell voltage (1.5V) of commercially available batteries.
Other design variables were set to default values. For example, tube width was set to 2.5 cm since the pump design that has this motor under the cams yields 2.5 cm space for the tube. The target flow rate was set to 130 ml/min. The Actual flow rate required was 100 ml/min, but a 130 ml/min target flow rate was used to provide a 30% safety margin. As a result of the design space exploration, many options were found, and five candidate pump designs are shown in Table 5.
Comparing efficiency and total pump volume, we find that the smaller pump volume tends to yield lower efficiency. As shown, one can choose any of these designs to achieve a flow rate of 130 ml/min. Thus, the ultimate decision can be based on the predominant design factor such as, for example and not limitation, overall pump size, efficiency requirements, or pumping volume. In this case, the second design yields a relatively small pump volume while retaining good efficiency.
Pump 3 (a third pump prototype) was built according to the Candidate 2 design from Table 5 above and tested to determine if its performance matches model predictions. Since an oval constant value of 0.36 was used for the optimization in the previous section, the same value was used in the model for preliminary validation. A 1.6×0.8 cm Penrose drain tube was used for the experiments and the results are shown in
The size reduction of the pump yielded an unexpected result when compared to pump 2. As shown, the flow rate did not increase linearly but exponentially as rpm increased. Comparison of the experimental results with the model results [OC=0.3 (1.6×0.8 cm) and OC=0.36 (1.6×0.8 cm) curves] shows that pump 3 provides flow rates similar to the model with OC=0.3 at low speeds, similar to the model with OC=0.36 at medium speeds, and greater than the model with OC=0.36 at higher speeds. As shown, the pump can achieve a 100 ml/min flow rate at 7.9V (66 rpm, 3.2% efficiency) and 130 ml/min flow rate at 9V (76 rpm, 3.2% efficiency). The electrical current varied between 0.05˜0.12 A.
The expected performance of the pump with 2.5×0.8 cm tube is also plotted in
While several possible embodiments are disclosed above, embodiments of the present invention are not so limited. For instance, while several possible configurations have been disclosed (e.g., a camshaft driven design and a SMA driven design), other suitable materials and configurations could be selected without departing from the spirit of the invention. In addition, the location and configuration used for various features of embodiments of the present invention can be varied according to, for example, the intended use of the pump, or a particular pumping volume requirement, viscosity, or chemical resistance requirement. Such changes are intended to be embraced within the scope of the invention.
The specific configurations, choice of materials, and the size and shape of various elements can be varied according to particular design specifications or constraints requiring a device, system, or method constructed according to the principles of the invention. Such changes are intended to be embraced within the scope of the invention. The presently disclosed embodiments, therefore, are considered in all respects to be illustrative and not restrictive. The scope of the invention is indicated by the appended claims, rather than the foregoing description, and all changes that come within the meaning and range of equivalents thereof are intended to be embraced therein.
The present application is a national phase entry and claims the benefit of PCT Patent Application Serial Number PCT/US2011/028885, and entitled, “Valveless Pump,” filed 17 Mar. 2011, which claims the benefit of U.S. Provisional Patent Application Ser. No. 61/314,631 filed 17 Mar. 2010, and entitled “Pump Design for a Portable Renal Replacement System,” which are hereby incorporated by reference in their entirety as if fully set forth below.
| Filing Document | Filing Date | Country | Kind | 371c Date |
|---|---|---|---|---|
| PCT/US11/28885 | 3/17/2011 | WO | 00 | 9/17/2012 |
| Number | Date | Country | |
|---|---|---|---|
| 61314631 | Mar 2010 | US |
