HK1125322B - Dual-channel pump cartridge and pump for dialysis use - Google Patents

Description

Dual chamber pump cassette, pump and method of use in a wearable continuous renal replacement therapy device

Reference to related applications

This application is a partial continuation of U.S. patent application No.10/940862, filed on day 9/14 2004 of U.S. patent application No.10/940862, which is a partial continuation of U.S. patent application No.10/085349, filed on day 11/16 2001 of U.S. patent application No.10/085349, which are all incorporated herein by reference. This application also claims priority from U.S. provisional patent application No.60/706167, filed on 8/5/2005, 60/706167, which is incorporated herein by reference.

Technical Field

The present invention relates to a pump for a dialysis system, and in particular to a pump and method of using a pump having a fully wearable dialysis or hemodiafiltration system. Furthermore, the present invention relates to a dual channel pulsatile pump cassette, and more particularly to a method and apparatus for valving a dual channel pulsatile pump cassette that provides a reliable open/closed valve and variable semi-circular fluid flow between the channels.

Background

Hemodialysis is a process by which microscopic toxins are removed from the blood using a filtration membrane (e.g., a dialyzer). Typically, hemodialysis is performed intermittently for a period of three to four hours, twice or three times a week. The results of renal replacement therapy in the form of hemodialysis are still not optimistic with respect to the quality of life, morbidity and mortality of these patients. Evolving literature and research suggests that daily dialysis may be beneficial for a variety of biochemical and clinical improvements and quality of life, and may also increase the longevity of patients with end stage renal disease. However, daily dialysis is almost impossible due to manpower and cost limitations. Furthermore, Continuous Renal Replacement Therapy (CRRT) is preferred over intermittent dialysis because more toxins can be removed from the blood using CRRT seven days a week and twenty-four hours a day. Some advantages of CRRT include an expected reduction in morbidity and mortality, a reduction in the amount of medication required, a reduction in fluid intake and dietary restrictions, and a greatly improved quality of life for ESRD patients.

Existing CRRT machines are large, heavy machines that provide dialysis, hemodiafiltration, or a combination of both to each patient on a daily basis. Existing CRRT machines are heavy and must be hung on several feet of tubing at an electrical outlet. In addition, these machines require continuous supply of several gallons of fresh water in order to produce dialysate fluid. Moreover, the patient must remain connected to the existing cumbersome and cumbersome CRRT machine for many hours, limiting his or her ability to perform normal daily activities.

An additional problem with existing dialysis machines is that frequent reconnection to the machine requires access to the blood stream by puncturing the arteriovenous branch. These branches last only for a limited time and are subject to infection, coagulation and other complications, which lead to a large number of hospitalizations and repeated surgical interventions.

On the other hand, implementation of daily dialysis encounters obstacles that make it practically impossible to implement on a large scale. Some of the barriers include the inability or reluctance of most patients to dialyze at home, the lack of nurses and technicians in the dialysis unit to provide more treatment, and the reluctance of government payers to undertake the expense of additional processing steps. Also, its implementation not only takes time, but requires a large capital investment to create additional capacity in the dialysis unit. Although home dialysis may solve this problem, most patients are unable or unwilling to use home dialysis machines. Therefore, there is a need for a solution that can increase dialysis time without incurring additional costs or unnecessary additional manpower.

Continuous Renal Replacement Therapy (CRRT) enables significantly higher doses of dialysis, but is not suitable for treating end-stage renal disease (ESRD) patients because the machines are heavy, are mounted on wall electrical outlets, and require many gallons of water.

Unsuccessful attempts have been made to create either a compliant or commercially available wearable dialysis device or a Wearable Artificial Kidney (WAK), which are capable of providing CRRT. Because of the cumbersome nature of common dialyzers and associated sorbent devices, the concept of a wearable dialysis device or WAK must also become practical for dialysis patients. In view of the above disadvantages, there is a continuing need for a portable, wearable CRRT device that can be used substantially continuously 24 hours a day and 7 days a week. There is also a need for an improved subsystem, such as a pump, sensor, etc., that can be included in the WAK so that the WAK can function properly as a CRRT device.

Disclosure of Invention

One embodiment of the present invention is directed to a dual channel pulsatile pump comprising a dual channel pump cassette having a first channel and a second channel, each channel of said dual channel pump cassette comprising a compressible peristaltic tube having an input end and an output end. An input valve is at an input, the input valve comprising: an input O-ring; an input ball abutting the input O-ring to prevent backflow of fluid through the input valve; and an input spring member that positions the input ball against the input O-ring when peristaltic tubing is compressed, the input spring member allowing the input ball to exit the input O-ring and allow fluid to flow forward through the input valve when the peristaltic tubing is allowed to decompress. There is also an output valve at the output, the output valve comprising: an output O-ring; an output ball abutting the output O-ring to prevent backflow of fluid through the output valve; and an output spring member that positions the output ball against the output O-ring when peristaltic tubing is decompressed, the output spring member allowing the output ball to exit the output O-ring and allow fluid to flow forward through the output valve when the peristaltic tubing is compressed. In an example embodiment, a pump motor part is further included, the pump motor part including: a first oscillatory pushing member for compressing and decompressing the peristaltic tube in the first channel; and a second oscillatory pushing member for compressing and decompressing the peristaltic tube in a second channel, the dual channel pump cassette being removably mounted on the pump motor portion.

Other applications of embodiments of the present invention will become apparent by consideration of the following detailed description and accompanying drawings. It should be understood that the description and illustration of the embodiments of the present invention is not intended to limit the scope of the invention, and that various changes and modifications within the spirit and scope of the invention will occur to those skilled in the art.

Drawings

The present invention will become more fully understood from the detailed description given herein below, taken in conjunction with the accompanying drawings, which are given by way of illustration at least, and are not intended to limit the scope of the present invention. In the drawings:

fig. 1 is a perspective view of a portable CRRT device worn around the waist of a dialysis patient according to the present invention.

FIG. 2 is a front view of the wearable CRRT device of FIG. 1 after detachment from a dialysis patient.

Fig. 3 is a perspective view of the dialyzer section of the wearable CRRT device of the present invention.

FIG. 4 is a perspective view of the additive pump and dialyzer section of the wearable CRRT device of the invention.

Figure 5 is a cross-sectional view of a first embodiment of a dialyzer of the wearable CRRT device of the present invention.

Figure 6 is a cross-sectional view of a second embodiment of a dialyzer of the wearable CRRT device of the present invention.

Figure 7 is a top view of the housing of the dialyzer of the wearable CRRT device of the present invention.

FIG. 8 is a perspective view of a first embodiment of the adsorbent portion of the wearable CRRT device of the invention.

FIG. 9 is a perspective view of a second embodiment of the sorbent portion of the wearable CRRT device of the invention.

FIG. 10 is a perspective view of a variation of the second embodiment of the adsorbent portion of the wearable CRRT device of the invention.

Fig. 11 is a top view of a housing of a sorbent device of the wearable CRRT device of the present invention.

FIG. 12 is a diagram of an example embodiment of a wearable CRRT apparatus.

Fig. 13 is a side view of an example dual channel pulsatile pump.

Fig. 13A and 13B are front and top views of an example pump-motor portion of an example dual chamber pulsatile pump.

Fig. 14 is an oblique view of an example pump-motor portion of an example dual chamber pulsatile pump.

Fig. 15 is a three-view illustration of an example dual channel impulse pump cartridge.

FIG. 16 is a three-view illustration of an example spring used in an example input or output valve.

Fig. 17 is an exploded view of an example dual channel impulse pump cartridge.

FIGS. 18A-18E are graphs showing the results of experiments using examples of wearable CRRT devices.

FIG. 19 is another example embodiment of a wearable artificial kidney.

Figure 20 is a schematic diagram of an example test apparatus for evaluating pump and dialyzer characteristics.

FIG. 21 is a schematic diagram of an example test apparatus for evaluating sorbent characteristics.

FIG. 22 is a graph illustrating the instantaneous flow performance of an example dual channel pump operating at 9 volts DC.

Fig. 23A and 23B provide a comparison of the transient performance of different types of roller pumps.

Fig. 24A and 24B provide a comparison of the instantaneous pressure performance of the example pump v.

Fig. 25A and 25B provide a comparison of the instantaneous pressure performance of an example pump v.

Fig. 26A and 26B provide test results of solute clearance using different types of pumps.

Fig. 27A and 27B provide test results of solute clearance using different types of pumps.

Fig. 28 provides a graph of test results of sorbent capacity of an example wearable artificial kidney.

Fig. 29 provides experimental comparisons of sorbent performance of example wearable artificial kidneys with and without optimization of dialysate pH.

Detailed Description

Referring to fig. 1 and 2, a Continuous Renal Replacement Therapy (CRRT) apparatus 10 is adapted to be worn about a portion of the body of a dialysis patient 15. The CRRT device 10 includes a belt 20, the belt 20 being divided into a plurality of sections including: a dialyzer section 30, the dialyzer section 30 comprising a blood inlet tube 33 leading from the blood vessel and a blood outlet tube leading to the blood vessel; an adsorbent portion 40; an additive pump portion 50; and an electronic control section 60, the electronic control section 60 including a microprocessor and a battery for supplying power to the apparatus 10.

As best shown in fig. 2, the strap 20 includes a pair of end portions 70, 75 that are secured together by a conventional strap fastener 80 (e.g., a buckle, snap, button, or hook and loop fastener). Although the CRRT device 10 is shown in fig. 1 as being worn around the waist of a patient 15, those skilled in the art will appreciate that the device 10 may alternatively be worn around other parts of the patient's body, such as over the patient's shoulders.

Referring to fig. 3, the dialyzer section 30 of the belt 20 includes a plurality of mini-dialyzers 100, 110, 120, 130 that utilize dialysate fluid 140 to remove impurities from the blood 150 of the patient 15. The number of dialyzers 100, 110, 120, 130 in the plurality of dialyzers 100, 110, 120, 130 may be varied in order to reflect different dialysis prescriptions. As best shown in FIG. 3, a plurality of dialyzers 100, 110, 120, 130 are connected in series, so that a common pump forces a patient's blood 150 through blood inlet tube 33, through dialyzers 100, 110, 120, 130, and into blood outlet tube 37. It will be appreciated by those skilled in the art that the dialyzers 100, 110, 120, 130 may also be connected in parallel without departing from the scope of the invention.

During dialysis, the dialysate is pumped in the opposite direction to the blood flow (as indicated by arrows 125, 135, 145) using a conventional pump (not shown). Spent dialysate 140 flows through spent dialysate tube 370 to sorbent section 40. Excess fluid is removed from the spent dialysate 140 by a volumetric device (volumetric)155 and into a waste receptacle 65 which is periodically emptied by the patient through a tap 175. A microprocessor in the electronics section 60 determines the rate and amount of fluid removed by the volumetric pump 155.

Referring also to fig. 3, the blood inlet tube 33 includes a side aperture 180 through which anticoagulant is pumped into the blood by an anticoagulant pump 190. A common anticoagulant is infused into the blood 150, including but not limited to: heparin, prostacyclin, low molecular weight heparin, hirudin and sodium citrate. As best shown in fig. 4, the blood outlet tube 37 includes a side hole 200 for injecting an additive that is forced into the blood 150 by a plurality of additive pumps 270, 280, 290, 300. Piston, suction, piezoelectric, micro or very small roller pumps may be used for this purpose. Such pumps may all be classified as micro pumps. Each additive pump 270, 280, 290, 300 forces a controlled amount of each additive into the blood 150, wherein the injection rate of each additive is electronically controlled by a microprocessor in the electronic control portion 60. In a known manner, the doctor can use the electronic control portion 60 to set the injection rate of each additive so as to correspond to a predetermined dose of each additive. Because the additives cannot be mixed together prior to infusion into the blood 150, they have a separate circuit 305. Typical additives include, but are not limited to: sodium citrate, calcium, potassium and sodium bicarbonate.

Referring to fig. 5, in a first dialyzer embodiment, each dialyzer 100, 110, 120, 130 is a common dialyzer comprising a plurality of cylindrical hollow fibers 310 through which blood 150 is circulated. As indicated by arrows 320, 330, the dialysate fluid 140 circulates around the outer walls 350 of the hollow fibers 310 in a direction across the blood flow inside the hollow fibers 310, as indicated by arrows 325, 335. The outer walls 350 of the hollow fibers 310 are semi-porous so that impurities can move from the blood 150 into the dialysate 140. Fresh dialysate 140 flows from the sorbent section 40 through a dialysate inlet tube 360 and into a series of dialyzers 100, 110, 120, 130. The used dialysate 140 then flows out of the series of dialyzers 100, 110, 120, 130 through the used dialysate outlet tube 370 and into the sorbent section 40. The dialysate inlet tube 360 includes a side aperture 380 (shown in fig. 3) for injecting additives that can be forced into the blood 150 by the aforementioned additive pumps 270, 280, 290, 300, and thus the rate of injection is electronically controlled by the microprocessor in the electronic control portion 60. Referring to fig. 6, in a second dialyzer embodiment, each dialyzer 100, 110, 120, 130 comprises a plurality of parallel plates 390 of semi-porous material, wherein the dialysate fluid 140 circulates on one side of the parallel plates 390 and the blood 150 circulates in a direction on the other side of the parallel plates 390.

Referring to fig. 7, each dialyzer 100, 110, 120, 130 is a mini-dialyzer having a flexible housing 400 for conforming to the contours of the patient's body. In addition, the body-side wall 410 of each housing 400 is concave to further correspond to the body curve of the user. The housing 400 may be made of any suitable material having suitable flexibility to conform to the body part in which it is used. Suitable materials include, but are not limited to: polyurethane and polyvinyl chloride.

Referring to fig. 8-10, in sorbent portion 40, spent dialysate 140 flows from dialyzer portion 300 through spent dialysate tube 370 and into a plurality of sorbent devices 420, 430, 440, 450, 460, as indicated by arrows 415. The regenerated dialysate 140 then flows through the tube 360 and back into the dialyzer section 30, as indicated by arrow 465. Preferably, the sorbent device 420, 430, 440, 450, 460 comprises a series of sorbent cartridges 420, 430, 440, 450, 460 for regeneration of the used dialysate 140. By utilizing the sorbent cartridges 420, 430, 440, 450, 460 to regenerate dialysate, the example CRRT device 10 requires only a small amount of dialysate in a single pass through the hemodialysis device. Importantly, each sorbent cartridge 420, 430, 440, 450, 460 is a small sorbent cartridge 420, 430, 440, 450, 460 containing a different sorbent.

Referring to fig. 8, in the first embodiment of the sorbent section 40, there are 5 sorbent cartridges 420, 430, 440, 450, 460, including an activated carbon cartridge 420, a urease cartridge 430, a zirconium phosphate cartridge 440, a hydrous zirconium oxide cartridge 450, and an activated carbon cartridge 460. Those of ordinary skill in the art will appreciate that these sorbents are similar to those used in commercially available recirculating dialysis (REDY) systems. However, in REDY systems, the sorbent is a single cartridge of multiple layers. Instead, the sorbent of the present invention is part of different sorbent cartridges 420, 430, 440, 450, 460, such that each cartridge 420, 430, 440, 450, 460 can be conveniently replaced and the other cartridges 420, 430, 440, 450, 460 independently disposed when needed. As known to those of ordinary skill in the art, activated carbon, urease, zirconium phosphate, hydrous zirconium oxide, and activated carbon are not the only chemicals that may be used in the CRRT device 10 of the present invention. Indeed, any number of additional or alternative adsorbents may be used without departing from the scope of the present invention.

Referring to fig. 9 and 10, in a second embodiment of the sorbent section 40, there are a plurality of sorbent cartridges 500, 510, 520, 530, wherein each cartridge 500, 510, 520, 530 comprises a plurality of sorbent layers 540, 550, 560, 570, 580: an activated carbon layer 540, a urease layer 550, a zirconium phosphate layer 560, a hydrous zirconium oxide layer 570, and an activated carbon layer 580. The cassettes 500, 510, 520, 530 may be in series, as shown in fig. 9, or may be in parallel, as shown in fig. 10. In this embodiment, the number of sorbent devices may be varied to correspond to different dialysis protocols.

Referring to fig. 11, each cassette 500, 510, 520, 530 is a small cassette having a flexible or curved housing 600, the housing 600 being adapted to conform to the contours of a patient. In addition, the body-side wall 610 of each casing 600 is concave so as to further correspond to the body curve. The housing 600 may be made of any suitable material having suitable flexibility to conform to the body part in which it is used. Suitable materials include, but are not limited to, polyurethane and polyvinyl chloride.

Referring to FIG. 12, another exemplary embodiment of a wearable CRRT apparatus is shown. The wearable CRRT device 700 is built into or is part of a patient wearable strap, belt or other wearable device 702. The strap 702 may include a pair of end portions 704, 708, the pair of end portions 704, 708 being adapted to be secured together by a fastening device (not expressly shown). The end portions/fastening means 704, 708 may be any number of fastening means suitable for securing the ends of a belt or strap together, and are not limited to: snaps, buttons, buckles, clips, braids, hooks and loops, zippers, buckles, and the like. One embodiment of the CRRT device may be envisioned in the shape of a ammunition or military-type supply belt, which may also be in the shape of a belt pack. The example wearable CRRT device 700 is worn by the patient on or under other clothing.

Microcontroller 714 is used to control and monitor various aspects of wearable CRRT device 700. Microcontroller 714 is preferably a low or very low power microprocessor, but essentially any microcontroller can be used to operate in the example wearable CRRT device 700. One of the many functions that a microcontroller has is to monitor the battery 716. The example CRRT apparatus 700 will operate continuously for at least 5 to 10 hours using less than 10 continuous watts of power. Preferably less than 3 continuous watts. The weight of embodiments of the present invention is less than 10lbs, preferably less than 5lbs, during operation.

The battery 716 is removably mounted in the wearable CRRT device 700. The battery 716 may be rechargeable and may be charged by a charging device (not shown) while remaining in the wearable CRRT device 700 or when disconnected from the wearable CRRT device 700. Preferably, the battery 716 can store sufficient energy to power the wearable CRRT device 700 for at least five (5) or more hours of continuous, non-intermittent device operation. The microcontroller itself or through additional circuitry monitors the state of charge of the battery 716. When the microcontroller 714 determines that the battery 716 is low on charge or less than a predetermined estimated amount of remaining operating time (e.g., 1 hour remaining), the microcontroller 714 may initiate an alarm condition via the alarm circuit 718. The alert circuit 718 may provide any combination of audio, video, or physical alerts. The physical alarm signal may comprise a vibration or a less stinging type of vibration to the patient. The alarm state or warning may be displayed on the display 720 using liquid crystal, light emitting diodes, or other low power display technology. The alert state may also turn off all or a predetermined portion of the example wearable CRRT device 700.

Humidity sensor 722 is also in electrical communication with microcontroller 714. The humidity sensor 722 is used to detect high humidity, condensation, or liquids present inside a package or cover (not expressly shown) on the wearable CRRT device 700. The packaging or cover over the example CRRT device 700 may be plastic, cloth, rubber, a polymer product, or other suitable material. The cover may cover a portion of the wearable CRRT device 700 and allow access to various portions of the device, such as the display 720 and the user/physician controls 723.

High humidity, coagulation, or the presence of liquid inside the wearable CRRT device 700 may indicate a patient blood leak, dialysate leak, or other fluid leak. By detecting humidity, the humidity sensor 722 provides a signal to the microcontroller 714 and activates an alarm via the alarm circuit 718. Also, the pump 724 may be shut off by the microcontroller 714 to help reduce the loss of blood, dialysate, or other fluid. The microcontroller may also turn off the micro-pump (described below).

The microcontroller 714 may also cause the communication device 725 on the device to contact medical assistance or other medical assistance organization. The communication device may include a wireless paging cordless telephone or other mobile communication circuitry. The communication device 725 may also provide the geographic location of the example wearable CRRT device 700.

The pump 724 is an electric pump. The pump 724 may be two pumps 724a and 724 b. The two pumps 724a and 724b may each be powered by the same or separate electric motors. The pumps 724a and b are powered by the rechargeable battery 716. Also, microcontroller 714 can be used to adjust various pumping variables. Possible adjustable pumping variables include, but are not limited to, adjusting pump stroke, volume per stroke, speed, torque, pumping rate (i.e., number of pump cycles per minute), pump pressure differential between the input and output of the pump, and pump stop and cycle time.

The example wearable CRRT device 700 has two fluid circuits: a blood circuit 727 and a dialysate circuit 729. A dual channel pulsatile pump 724 can be used in the example embodiments. Typically, the pulsating pump has a cartridge for the rubberizing of each channel. The cartridge has an input valve at the input side of the cartridge and an output valve at the output side of the cartridge. Fig. 12 shows a single-direction, double-pulsation pump 724. A bi-directional, double pulsation pump may also be used. A bi-directional channel pump is preferred in order to reduce the bending of the tubing used in the fluid circuit.

The motor and transmission in the pulsating pump press the rubberized tubular portion of the cartridge. The pressing of the cartridge causes the contents of the cartridge to be squeezed and expelled outwardly from the outlet valve. When the pump motor rotates and causes the mechanism of the pump to release pressure from the rubber-treated portion of the cassette, the output valve closes and the input valve opens to allow fluid (blood or dialysate) to enter the cassette so that fluid can be expressed from the output valve in the next pump cycle. The input and output valves are one-way valves, thereby allowing one-way fluid flow through the cartridge. Other configurations of the pulsating pump are possible. The example pumps 724a, 724b provide blood flow rates (pulsatility) at about 15 to 100 ml/min. The approximate size of the exemplary dual pulsatile pump 724 is 9.1 × 7.1 × 4.6cm and weighs less than 400 grams. The approximate dimensions of the dual channel pump cassette were 0.72 inches high, 3.82 inches long, and 1.7 inches wide. Example embodiments may be any one or more of these dimensional measurements plus or minus 50%.

The example pulsatile pump uses less than 10 watts of energy and can provide lower battery power and a pump occlusion warning signal to microcontroller 714. Low power pulsating pumps of 5 watts or less may also be used.

The pulsatile pump can be adjusted so that the phases of the pulses or cycles of the two pulse chambers will be 180 ° out of phase or any predetermined number of degrees out of phase in order to utilize the pulses of the pump to help maximize the dialysis treatment performed in dialyzer 730. The relative directional flow of blood and dialysate through the dialyzer may become more efficient with different phase settings of the pumps 724a and b.

Referring to fig. 13, 14 and 15, wherein fig. 13 illustrates a side view of an exemplary dual chamber or dual channel pulsatile pump 800. The dual channel pulsatile pump can include at least two different components: a pump-motor portion 802 and a dual channel pump cassette 804. Fig. 13A and 13B provide only exemplary front and top views of the pump-motor portion 802. Fig. 14 provides an isometric view of an example pump-motor portion 802. Fig. 15 provides three views (side view a, top view B, and end view C) of an example dual channel pump cassette 804. The example dual channel pump cassette 804 has two components that simultaneously provide synchronized pulsatile motion of one or two different liquids. This detachable passive dual channel pump cassette 804 is configured to fit into the structure of a dual channel pulsatile pump 804 such that the mechanical pump element 806 of the pump-motor portion 802 can cause the chamber pumping elements 907 to expand and contract in a rhythmic fashion, drawing fluid into the flexible chamber 908 and then pushing it out the other end of the chamber. The example dual channel pump cassette 804 is somewhat isolated from the pump-motor portion 802. In a two-chamber or dual channel embodiment, the dual channel pump cartridge 804 produces a first flow of liquid in a first half cycle and a second flow of liquid in a second half cycle. For example, blood flow may be pumped from the first channel 902 in a first half-cycle and dialysate may be pumped from the second channel 904 in a second half-cycle.

The driving of the two chambers of the removable cartridge 804 is performed by an oscillation 905 (see arrow in fig. 13A) of a pushing mechanism 906 of the mechanical pump part, the pushing mechanism 906 being connected to an output shaft 910 of a fixed or variable speed motor 912 by a drive linkage 914, the drive linkage 914 being capable of transforming the rotational movement of the output shaft 910 into an oscillating movement 905 of the pushing mechanism 906. The rotational motion of the pump output shaft 910 is caused by a gear structure that produces both circular and non-circular motion. This combination of the motor-driven pushing mechanism 906 and the dual channel pump cassette 804 can be powered by a variety of power sources, including batteries, DC power, or AC power.

The dual channels 902, 904 of the removable cartridge 804 may be configured to provide the same or opposite flow directions of the two liquids. Figure 5 shows the flow of two channels 902, 904 in the same direction. One channel may be made in an inverted or swiveling manner such that cassette 804 causes fluid to be pumped through channels 902 and 904 in opposite directions. Each channel 902 and 904 of the removable cartridge 804 has a release mechanism 920 at the channel inlet 916 and outlet 918, the release mechanism 920 adjusting the flow direction according to the configuration. The release mechanism 920 may be of the one-way valve type configured to allow fluid flow in one direction when a positive pressure is applied, or in another configuration, only when a negative pressure is applied. An example release mechanism 920 is shown that may include an occluder-retainer combination (described in more detail below) configured to allow flow in one direction when a positive pressure is applied. Example embodiments of the removable cartridge may include a biocompatible material. Fluid flow regulation is performed substantially independently of the density, viscosity and physical properties of the flowing liquid. When used to pump blood, the removable cartridge has a low hemolysis characteristic.

The example fluid pump-motor portion 802 has multiple or single pumping members 906 that simultaneously provide synchronized pulsatile motion of one or more different separate fluids in each cartridge channel 906. With the pumping chamber 908, a degree of isolation from the mechanical pump element 806 is obtained in each channel of the dual channel pump cassette 804. The detachable portion 804 of the dual channel pulsatile pump 800 fits into the structure of the pump-motor portion 802 such that the mechanical push mechanism 906 can expand and contract the chamber pumping elements 907 in a rhythmic fashion, drawing liquid into the input side 916 of the flexible chamber 908 and then pushing it out of the output end 918 of the chamber. In the two-chamber embodiment shown, the mechanism produces a first flow of liquid in the first half of the cycle and a second flow of liquid in the second half of the cycle. For example, blood flow in the first half cycle and dialysate in the second half cycle.

In another embodiment of the pump-motor portion 802, the mechanical pump component is connected to the output shaft of a fixed or variable speed, reversible or irreversible motor through a drive linkage, which is a derivative of a crankshaft (not specifically shown), capable of converting the rotational motion of the output shaft 910 into an oscillating motion of a pushing mechanism. The example structure includes circular and non-circular motion.

In another embodiment of the pump-motor section, the oscillating action of the pushing mechanism can be adjusted so that pushing does not occur in alternate half cycles only, or 180 degrees out of phase. In contrast, dual pushing mechanism 906 may be adjusted or configured to push or compress the flexible chambers of the first and second channels from in-phase (i.e., simultaneously) to 180 degrees out of phase (i.e., alternating half cycles).

In experiments using embodiments of the present invention, an example dual channel pulsatile pump 800 was used. The custom, battery-powered, or AC-to-DC powered pump may be the mechanical motor of the example wearable renal replacement device of various embodiments of the present invention and is used to simultaneously provide pulsatile fluid flow of blood and dialysate through the example renal replacement therapy device. The dual channel pulsatile pump 800 includes two components: a micro-motor portion 802 and a flow cartridge 804, as described below:

3 watt DC micromotor: the specification of the motor is as follows:

	manufacturer of	FAULHABER, Germany (available from MicroMo, Inc., Florida, USA)
				Part number	1331012S
	Voltage	Rated at 12 volts
				Terminal resistor	13.3Ω
	Output power	2.62 watts
				Efficiency of	77％
	Rotate at a high speed to	12000RPM
				Torque to	0.354oz-in
	Current flows to	0.300 ampere
				Gear head	15/5
	Estimating lifetime	300 hours of continuous operationAccording to the manufacturer)

Dual channel flow box 804: the gear head of the micropump was modified to accommodate an oscillating mechanism that was connected to a custom made dual channel flow cassette, enabling simultaneous pulsatile flow of blood and dialysate at a controlled flow rate of 40-100cc/min per channel. There is no connection between the two channels and the design of the dual channel pump cartridge 804 enables the two fluids to flow in the same direction or in opposite directions depending on the needs and configuration/location of other system components. The dual channel flow cassette 804 includes standard FDA approved PVC tubing, standard FDA approved Delrin stopper, standard FDA approved Latex (Latex) retainer, all encased in a Lucite resin (Lucite) housing that fits over the micromotor housing, as shown in fig. 13.

The valves at each end of each channel 902, 904 of the flow cartridge 804 are one-way valves that open when pressure in the positive fluid flow direction is applied to them. Suction valve 922 includes an inlet 924, an O-ring 926, a manifold head inlet 928, a blocking ball 929, a merzides spring 930, and a peristaltic tubing fitting 932. When peristaltic tubing 934 mounted on peristaltic tubing fitting 932 is expanded or depressurized after being pressed by pushing mechanism 906, suction is created on suction valve assembly 922 such that ball 929 presses against merzides spring 930 and flexes spring 930, thereby allowing blocking ball 929 to exit O-ring 926 and fluid to flow through suction valve assembly 922 into peristaltic tubing flexible chamber 908.

Referring to fig. 16, a schematic front and side view of a meisszid spring and various exemplary dimensions are shown. The mercedes spring 930 is made of a polymer, much like the O-ring 926, but may also be made of metal, plastic or a composite material. The mercedes spring 930 comprises two concentric rings: an inner ring 1000 and an outer ring 1002. Three spokes 1004, each extending radially from the inner ring 1000 to the outer ring 1002, connect and stabilize the positioning of the inner and outer rings. The inner ring 1000 acts as a stabilizer and positions the blocking balls 929. It will also be appreciated that the inner ring 1000 helps reduce the rapid or oscillatory movement of the occluding ball 929 during fluid flow, thereby providing a higher fluid flow rate than other example occluding valves. The inner ring 1000 of the merzides spring also centers the ball so that the valve can quickly close when the fluid flow decreases. Internal ring 1000 also serves to center ball 929 on O-ring 926 when suction valve 922 is in the closed position, and to significantly limit the amount of fluid backflow through valve 922 or prevent all possible backflow when the state of peristaltic tube 934 changes from a decompressed state to a compressed state and when pushing mechanism 906 pushes against peristaltic tube 934. In another embodiment, the meissies valve 930 may be replaced by a valve having an inner ring 1000, an outer ring 1002, and (only) two radial spokes 1004 (the radial spokes 1004 extend from the outer ring 1002 to the inner ring 1000). When two spokes 1004 are used, the spokes may be positioned substantially opposite or 180 degrees from each other. Inner ring 1000 and outer ring 1002 are shown as concentric rings, but alternative embodiments of the present invention may have inner and outer rings that are not concentric.

Output valve 936 is connected to peristaltic tubing 934 by another peristaltic tubing coupling 932 and O-ring 926 acts as a sealing seat for ball 929 inside outlet manifold 938. Second mercedes spring 930 flexes when pushed by ball 929 to allow fluid to flow through output valve 936 when peristaltic tube 934 is compressed by pushing mechanism 906. Furthermore, the dual concentric rings 1000, 1002 of the merzides spring 930 resist flutter of the ball 929 during fluid flow, thereby maximizing fluid flow and maintaining smooth fluid flow through the output valve 936. The tremor of the ball 929 is also undesirable because it can damage blood cells as blood flows through the example valve of the pulsatile pump. Also, the merseiders valve 930 closes when the compression of the peristaltic tubing stops and allows the ball 929 to smoothly seat against the O-ring 926, thereby limiting or eliminating backflow through the valve and thus limiting damage to blood cells passing through the blood channel of the dual channel pulsatile pump cassette 800.

An example dual channel pulsatile pump 800 including a pump-motor portion 802 and a dual channel pump cassette 804 was tested for in vitro reliability for one week on applicants' equipment at the Cedars-Sinai medical center in los angeles, california. Deionized water having the same density and viscosity as the dialysate is used in the dialysate channels. A mixture of approximately 40% glycerol (viscosity 3.5cp) in water at room temperature was used as the blood replacement fluid in the blood channel. Flow rate data were obtained and recorded by tapping a supersonic flow probe (type 2 XL) connected to a T110 supersonic flow meter (transonic system, ixa, new york) to produce a ± 1 volt output signal that was input to a LabView-based data acquisition system (integrated in a separate workstation).

The hemolysis characteristics of the model dual chamber pump cassette were estimated according to the protocol based on ASTM standard F-1841-97. The maximum working blood flow rate was 78ml/min and the pressure differential was 100mmHg, with free hemoglobin plasma adjusted for relatively static control averaging about 5 mg/dL/hr. These results show that hemolysis of the exemplary dual channel pulsatile pump is within acceptable boundaries when manufactured into its final modified form (giving a wider range of published data). To check this point, the same hemolysis protocol was also used for three commercially available dialysis roller pumps. The corresponding free hemoglobin plasma was approximately 3mg/dl/hr and 450ml/min flow. The hemoglobin destruction rates corresponding to the exemplary dual channel pulsatile and roller pumps were 25 and 15mg/hr, respectively, when a total of 5dL (500ml) of fluid was available for each pump for the six hour test. Since a particular pump running at a particular flow rate will logically damage hemoglobin at the same mg/hr rate, regardless of how much blood is obtained, this estimate may be well used for patients with 5 liters of blood. When an average of 25mg/hr is practically acceptable (blood flow rate of about 78ml/min), the patient will lose about 0.6 gm/day of hemoglobin. It will be appreciated that the average red blood cell weight is about 0.055 nanograms, and that healthy people will produce about 2 billion per day, and therefore 11gm per day. However, this would not be the case in ESRD patients who are not supplemented with, for example, Epogen et al. It is expected that a dual channel pump user will lose approximately 1.7 gm/day of red blood cells based on a 0.6 gm/day loss of hemoglobin and a hemoglobin/red blood cell ratio of 0.35. The estimated hemolysis will not cause any problems when ESRD patients maintain at least 130 gm/liter of red blood cells (considering only 0.25% of 650gm red blood cells in total).

The results from the hemolysis test described above also show a standardized index of hemolysis (NIH) of approximately 0.4 mg/dL. This NIH, also known in the literature as the Haemolytic Index (HI), is a measure of the added grams of free hemoglobin plasma per 100L of blood pumped, corrected for plasma volume using a hematocrit and standard flow rates and circulation time. The commonly reported HI value is between 0.1 and 0.2mg/dL because standard dialysis roller pumps operate multiple times at higher flow rates than the exemplary dual channel pump.

However, it should be appreciated that other types of pumps 724 (FIG. 12) than pulsatile pumps can be successfully used or incorporated into the wearable ultrafiltration device embodiments, provided that the power consumption is low enough for battery operation. The power consumption of the pump will be less than about 5 watts, and it is even better between 1 and 3 watts. Two separate single channel pumps may be used. Other types of pumps include, but are not limited to: reciprocating pumps, piston pumps, roller pumps, centrifugal pumps, piezoelectric pumps or other conventional pumps. Regardless of the pump used, the pump 724 will have a manually or electrically adjustable flow rate in the range between 20ml/min and 120ml/min, and the power required is between 1 and 5 watts.

Fig. 17 provides an exploded view of an example dual channel pulsatile pump cassette 804 according to an embodiment of the present invention. The center panel 1100 serves as a floor for the dual fluid channel. Each of the fluidic channels in this example embodiment includes an inlet manifold 1102 and an outlet manifold 1104. Peristaltic tubing 1105 provides a squeezable or collapsible flexible chamber that is elastically deformed by the urging mechanism of the mechanical pump element. Tangential ball seat 1106 is used to place ball 1108 in a valve-closed position. An O-ring may be disposed with or in place of the tangential ball seat 1106. The blocking ball 1108 is used to block or allow flow through the valves. The pellets may be made of plastic, polymer or metal. At the inlet end of the barbed fitting 1110, an adhesive barb 1112 may be used to help connect the fluid conduit to the inlet side of the fluid channel. Similarly, on the outlet side of outlet barb fitting 1114, bondable barb 1112 may be used to connect the fluid pipe to the outlet side of the channel. Peristaltic tubing connections 1116, 1118 are provided for the inlet side 1116 and the outlet side 1118 of the peristaltic tubing 1105. Also, the mercedes spring 1120 is positioned such that: the concentric inner ring is caused to move perpendicular to the outer ring in the direction of fluid flow when a positive fluid pressure in the direction of fluid flow is placed on the ball 1108. In the absence of positive pressure, the Messels spring 1120 presses the ball 1108 against the tangential ball seat, thereby sealing the passageway and preventing fluid flow in the opposite direction.

Referring to fig. 12, microcontroller 714 may display pump status or other pump related information on display 720. User controls 723 (as buttons, switches, slide controls, knobs, connectors, or infrared receivers, not specifically shown) may be used to enable a patient, doctor, nurse, technician, or computer-based device to adjust various settings and controls of the example ultrafiltration device 700. Further, the communication means 725 may be used for receiving control settings and transmitting information over a wireless paging or other telecommunication communication channel. For example, adjustments in the pumping rate, torque, valve opening size, output pressure, flow rate, rpm, and on/off of the pump 724 may be monitored or controlled via the user interface 723 or the communication device 725.

Referring first to the exemplary blood circuit 727, blood from a patient enters the blood circuit 727 through the blood inlet tube 726. An input blood pressure sensor 728 measures the input blood pressure and provides an input blood pressure signal to microcontroller 714 (connections to the microcontroller not specifically shown). The input blood pressure may be the mean pressure of the blood prior to entering the pump 724 a. The blood is then pumped by pump 724 a.

After the blood passes through the main pump 724a, it continues in the blood circuit 727 through the blood inlet tube 726. An input blood pressure sensor 728 measures the input blood pressure and provides an input blood pressure signal to microcontroller 714 (connections to the microcontroller not specifically shown). The input blood pressure may be the mean pressure of the blood prior to entering the pump 724 a. The blood is then pumped by pump 724 a.

After the blood passes through the main pump 724a, it continues in the blood circuit 727. A reservoir 734 containing a blood diluent or anticoagulant (e.g., heparin or other acceptable anticoagulant additive) is connected to the blood circuit by a micro-pump 736. Prior to the dialyzer 730, a micro-pump 736 supplies the fluid content of the reservoir 734 to the blood circuit 727 in a continuous or discontinuous occurrence of measurement. (the reservoir 734/pump 736 combination may be connected to the blood circuit before the pump 724 a.) the micro-pump 736 is of a type that is capable of pumping microscopic or minute amounts of fluid per minute. Micropumps can typically pump fluids at rates ranging from 0.1 to 400ml/hr (milliliters per hour). The micro-pump requires about 1 to 500 milliwatts to operate. Various pumps are currently contemplated as micropumps, including but not limited to: piezoelectric pumps, solenoid pumps, micro-piston pumps, peristaltic pumps, nanotechnology related pumps, microtechnology/micromachine pumps, syringe-type pumps, roller pumps, centrifugal-type pumps or diaphragm-type pumps.

The blood diluent and/or anticoagulant can mix or combine with the blood in the blood circuit at any point between the inlet of the blood inlet tube 726 and the blood input side of the dialyzer 730.

The reservoir 734 may have a fluid level sensor 735 or other type of sensor to detect the amount of fluid available in the reservoir 734. Sensor 735 provides a signal to microcontroller 714 indicating the amount of fluid in reservoir 734. When the fluid level or amount of fluid in the reservoir 734 is below a first predetermined amount or volume, the microcontroller 714 sends a warning signal to the warning circuit 718. The microcontroller 714 may also turn off the ultrafiltration device 700 when the fluid level in the reservoir 734 is at or below the first predetermined level and at a second predetermined level.

The combination of the reservoir 734 and the micro-pump 736 loads the blood diluent or anticoagulant into the blood flowing in the blood circuit 727. Also, a diluent or anticoagulant is injected into the blood before dialyzer (or hemofilter) 730 (and in some embodiments before blood pump 724a) to help reduce blood clots that may occur in hemofilter 730 (and possibly blood pump 724 a).

A second pressure sensor 733 detects the pressure in the blood circuit after the blood pump 724a (but before the dialyzer 730). The pressure readings are fed to a Microcontroller (MC)714, which monitors the readings.

The dialyzer 730, represented as a single dialyzer, may be a single or multiple dialyzers as previously described. The dialyzer may be in the form of a cassette that can be "snapped in" or inserted into and removed from the blood/dialysate circuit by a doctor, nurse, or technician. The dialyzer may comprise a dialysis surface area of 0.2 to 1 square meter. During dialysis, the flow in the blood circuit 727 is in the opposite direction as in the dialysate circuit 729 to aid in maximizing dialysis treatment. Furthermore, the pulses of pumps 724a and b may be in phase or out of phase, also helping to produce maximum dialysis treatment.

After dialysis in dialyzer 730, the blood exits dialyzer 730 and flows through third pressure sensor 737. The third pressure sensor 737 provides a pressure signal to the microcontroller. The combination of the first, second and third sensors provides a differential pressure measurement that can be analyzed by the microcontroller 714. For example, when the pressure differential across dialyzer 730 is too high, it may mean that there is a clot in dialyzer 730 first, or that its working blood flow is too high. Thus, a warning condition may be initiated, or the pumping rate or torque of the blood pump 724a may be adjusted by microprocessor control. When the pressure at the sensor drops below a predetermined pressure, it may indicate a fluid leak or air in the blood circuit 727. Microcontroller 714 can shut down all or predetermined portions of wearable CRRT device 700 in response to measured pressure being below a predetermined pressure level.

The blood is returned to the patient through blood outlet tube 740. As shown in fig. 4, a side port 200 may be included so that additional electrolytes, drugs, blood additives, vitamins, or other fluids may be added to the blood in the blood circuit 727 through the reservoir/micro-pump combination before the blood is returned to the patient through the blood outlet tube 740.

Referring also to fig. 12, an example dialysate circuit will be described. The fourth pressure sensor 750 measures dialysate pressure on the input side of the dialysate pump 724b and provides a pressure reading to the microcontroller 714. The dialysate pump 724b, like the blood pump 724a, is preferably part of the dual pump arrangement 724 described above, but may also be a separate pump arrangement.

Clean, fresh dialysate from the sorbent filter 769 flows in the dialysate circuit 729 through the dialysate pump 724 b. The dialysate pump 724b can pump dialysate at a flow rate ranging from near zero to 150 ml/min. An exemplary normal operating flow rate for the dialysate pump is between 40 and 100 ml/min.

Embodiments of the wearable CRRT device 700 are designed to work with less than 1 liter of dialysate. Preferably, embodiments only require 300ml to 400ml in the closed dialysate fluid circuit 729 to work. Embodiments designed for young adults or children may work with about 100 to about 300ml of dialysate. The combination of dialysate and filter 769 enables embodiments to circulate dialysate for at least 24 hours before the filter needs to be replaced. Moreover, because fewer liters of total dialysate are needed in the closed dialysate circuit 729, sterile or ultra-pure dialysate can be economically used in WAK example embodiments or wearable CRRT devices 700.

In normal or larger dialysis machines, about 90 liters of dialysate is typically used per patient run. Typically, filtered water (rather than ultrapure water) is used because of the amount of water required to produce the dialysate. Filtered water is much cheaper than ultrapure or sterile water. The filtered water used in dialysis machines allows some bacteria to be in it. These bacteria are larger than the size of the pores of the membrane used in dialyzer 730. Because the bacteria are larger than the size of the pores, the bacteria cannot cross the membrane into the blood.

In contrast, medical research has provided some results where the use of non-sterile dialysate (dialysate containing filtered water, bacteria, toxins or microorganisms) is inappropriate. Microorganisms and bacteria produce waste, toxins or toxic substances in the dialysate. Waste products from bacteria may pass through the dialyzer aperture into the patient's blood, although the bacteria are not actually able. These toxins are in some cases referred to as endotoxins. Endotoxin passing from the dialysate to the blood has an adverse effect on the health of the patient. Endotoxin can cause nausea in patients.

Because the wearable CRRT device 700 of the example embodiment requires less than 1 liter of dialysate, ultrapure or sterile water can be economically used in the manufacture of dialysate.

The dialysate leaves the dialysate pump 724b and passes through another pressure sensor 752, which pressure sensor 752 measures the dialysate pressure at the input side of the dialyzer 730. The dialysate circuit 729 passes dialysate into the dialyzer 730 so that the dialysate preferably moves in a direction opposite to the direction of blood flow through the dialyzer. When the dialysate is in the dialyzer 730, waste and toxins in the blood pass through the membrane of the dialyzer, thereby cleaning the patient's blood.

The dialysate exits the dialyzer 730 and flows through another pressure sensor 754. The pressure sensor 754 on the output side of the dialyzer 730 sends a signal to the microcontroller 714 indicating the pressure of the dialysate. This pressure helps indicate clogging, leakage or other emergency of the dialyzer.

The dialysate circuit 729 sends spent dialysate containing toxins or contaminants to a first series of dialysate filters 769. The filter may filter or react with predetermined substances in the dialysate to recirculate the dialysate in the dialysate circuit for continued use.

In the exemplary embodiment, first filter 760 includes urease. Urease filters the spent dialysate and a further function is to break down the urea removed from the blood in dialyzer 730. When urease breaks down urea, at least two undesirable by-products are produced. Typically, these two by-products are ammonium (ammonia) and carbon dioxide.

The dialysate with ammonia and carbon dioxide exits the first filter 760. Urea is substantially removed from the dialysate, but ammonia and carbon dioxide also need to be removed from the dialysate. The dialysate, ammonia, and carbon dioxide enter a second filter 762. The second filter 762 contains a compound comprising zirconium or zirconium phosphate (i.e., ZrPx). The zirconium in the second filter 762 captures the ammonia. One of ordinary skill in the art of dialysate chemistry will recognize that a variety of chemicals and their derivatives can be used to achieve the same or similar results.

The zirconium filter (second filter 762) will eventually become saturated with ammonia. When becoming saturated with ammonia, the zirconium filter will become less efficient at removing ammonia from the dialysate. It is not advantageous to enable ammonia or ammonium to be circulated through the dialysate circuit 729. Thus, in the example wearable CRRT 700, a sensor 764 is disposed in the dialysate circuit 729 to detect the presence of ammonia in the dialysate. The sensor 764 may be a ph sensor, an ammonia specific sensor, or a conductivity sensor. When an ammonia sensor is used, it will detect whether a predetermined amount of ammonia is present in the dialysate. When a ph sensor is used, it will detect whether the ph of the dialysate is biased more than normal toward a basic predetermined amount. When ammonia is present, the dialysate will become more alkaline. It will be appreciated that depending on the actual chemicals and adsorbents used in the filter, the dialysate may become more acidic and therefore the sensor will be used to detect this acidity. When a conductivity sensor is used, it will detect a change in conductivity of the dialysate.

The sensor 764 is in electrical communication with the microcontroller 714. An alarm state is activated by the microcontroller 714 when a signal read by the microcontroller 714 or provided to the microcontroller 714 from the sensor 764 indicates that the second filter 762 (zirconium filter) is not adsorbing most or a predetermined amount of ammonia in the dialysate. The warning state will indicate to the user that one or more filters (cartridges) need to be replaced. The alert state may also shut down predetermined functions of the wearable CRRT device 700. For example, one or more of the pumps 724 may be turned off, or the pumping rate of one or more of the pumps and the micropump may be reduced. Decreasing the pumping rate may increase the amount of ammonia adsorbed by the zirconium-based filter in the adsorbent filter portion 769.

A sensor 764 for detecting the presence of ammonia in the dialysate is arranged behind the second filter 762 containing zirconium phosphate. The sensor 764 may be disposed after the third filter 766 containing hydrous zirconia or the fourth filter 768 being a carbon filter. One or more sensors in the dialysate circuit will detect pressure, pH, ammonia, flow rate, temperature, or other physical properties. The sensor will provide a signal to the microcontroller indicating that the dialysate circuit needs maintenance.

A third example filter 766 is a hydrous zirconia (ZrOx) filter that can further remove contaminants and ammonia from the dialysate. The bubbler degasser or valve assembly 770 may be part of the filter (i.e., 762, 766, or 768), or a separate element (as shown) that removes air, carbon dioxide, and other bubbles from the dialysate. Importantly, a limited amount of air bubbles pass through dialyzer 730. Thus, the bubbler(s) 770 should be located before the pump 724b, but after the filter that can cause bubbles to form in the dialysate.

The fourth example filter 768 contains carbon and is used to further remove impurities from the dialysate by adsorption. As mentioned above, the filter is preferably designed as a filter cartridge. The cassettes may be inserted into and removed from the wearable CRRT device 700 by a patient, doctor, technician, or nurse. Each filter cartridge 760, 762, 766, 768 may contain a layer or combination of chemicals or adsorbents. In fact, an exemplary embodiment may have a single filter cartridge that contains multiple layers of the desired substances to clean and regenerate the dialysate after passing through the dialyzer 730. The filter cartridges may each include a bubbler device, or the bubblers 770 may be separate elements in the dialysate circuit 729.

Example wearable CRRTIn the device 700, the cartridges may be replaced by the patient daily or every other day. The dry weight of each cartridge will be less than half a pound. The combined total dry weight of all cartridges will be less than two pounds. Each filter cartridge may have internal dimensions of about 4cm by 10cm, or about 400cm for each adsorbent material³±100cm³The volume of (a). Regardless of the quantity, combination, used, the total volume of all adsorbent materials may be about 400cm³±2000cm³In the meantime, the method is described. In an example embodiment, the filter cartridges may be changed one at a time in a day or longer.

An additive reservoir 772 and a micro-pump 774 may be connected with the dialysate circuit 729 after the filter cartridge 769 but before the pump 724 b. Although not specifically shown in fig. 12, a plurality of reservoirs 772 and micropumps 774 can be connected with the dialysate circuit 728. Micropump 774 may be any micropump described above with respect to micropump 736. Here, the micro-pump 774 and reservoir 772 may add chemicals and additives to regenerate the dialysate and prolong its ability to function as dialysate. The example wearable CRRT device 700 may have only 300ml to about 1 liter of dialysate in the dialysate circuit 729. Importantly, the sorbent portion 769 is capable of cleaning and regenerating the dialysate as it circulates around the dialysate circuit 729.

The example wearable CRRT device 700 may also remove ultrafiltrate or fluid from the patient's blood. The patient's kidney does not function properly. After the dialysate exits the dialyzer 730, preferably before the dialysate enters the filter cartridge 769, the ultrafiltrate/dialysate and other contaminants and fluids obtained through the dialyzer 730 can be removed from the dialysate circuit 729 through a valve 776 and deposited in a fluid bladder 778. The fluid bladder 778 may hang below the wearable CRRT device 700 (not specifically shown) and may be capable of storing from about 0.1 to 2 liters of fluid. A fullness sensor associated with the fluid bladder 778 is in electrical communication with the microcontroller 714 to enable generation of a warning state when the fluid bladder 778 is at a predetermined fullness.

The fluid bladder 778 may also be contained in the wearable CRRT device 700 as an empty cartridge that is filled by a combination of micro-pumps and valves 776. The fullness sensor 780 can assist the microcontroller in determining the fullness of the cassette bladder 776, which will turn off the ultrafiltrate supply micro-pump 776 and provide a signal to the user indicating that the cassette needs to be emptied. The fluid bladder or cartridge 778 may contain an adsorbent material (not specifically shown) for adsorbing the fluid present to the bladder 778. The adsorbent material may be cotton, a polymer, a sponge, a compressed material, a powder, a gel, or other material that adsorbs fluid and/or limits sloshing in a soft capsule or cartridge. The bladder may be designed to expand when it is full. The bladder may press against a microswitch 780 (not expressly shown) when it is full, providing a signal to the microprocessor 714.

The fluid bladder or cartridge 778 may have means 782 thereon for evacuating the fluid bladder which may be a cap, plug, valve, removable internal bladder or other form.

Referring again to the blood circuit in fig. 12, a reservoir/micro-pump combination 784 (piezo pump, solenoid pump, syringe pump, etc.) may be connected to the output side of the blood circuit dialyzer 730, 727. One or more micropumps may be coupled to the fluid reservoir 784. Added heparin, electrolytes, blood additives, drugs, vitamins or hormones may be added to the dialyzed blood returning to the patient's body. The reservoir/micropump combination may be monitored and controlled by a microcontroller and may be regulated by a user controller 723 or by instructions received by the communication device 725.

Exemplary embodiments of the wearable CRRT device may provide a variety of treatments to a patient from basic dialysis functions to more complex medical dialysis, ultrafiltration, and drug therapy.

As previously mentioned, the evolving literature suggests that increasing dialysis time (with longer and more frequent dialysis treatments) can improve treatment outcomes in End Stage Renal Disease (ESRD) patients with respect to quality of life and expected morbidity and mortality.

However, performing treatment in this manner would be complicated by the lack of readily available economic sources to pay for increased time or more frequent dialysis treatments. Moreover, even though there may be money for paying more dialysis time or treatment, there is currently limited additional care or technical manpower to provide more additional care. Furthermore, the structure of the additional device will need to accommodate all these additional requirements. Due to budget restrictions in healthcare budgets in most countries, the variation in any or all of these things is small. Moreover, very few dialysis patients are eligible for home self-treatment by means of non-wearable dialysis devices.

Embodiments of wearable CRRT devices are typically worn by a patient on a belt or strap and may be used for continuous renal replacement therapy 24 hours a day, 7 days a week. This embodiment can provide a significantly higher dose of dialysis than is currently typically performed by dialysis equipment in intermittent doses, while significantly reducing costs associated with human use and other medical treatments.

Recently, the examples of the present invention were tested to evaluate the efficiency and viability of the present invention in uremic pigs. The efficiency of the example wearable CRRT device was estimated by obtaining the amount of urea, creatinine, potassium, phosphorus and ultrafiltrate removed, which would standardize the volume status and the above chemicals in uremic persons (when the device would be worn continuously). Moreover, the efficiency of the device is tested by the dialysis dose obtained, which will be equal to or higher than the dose obtained by intermittent daily dialysis, as measured by creatinine clearance, urea clearance and urea Kt/V per week.

An example embodiment of a wearable CRRT device for use in testing includes a blood circuit and a dialysate circuit. The blood circuit and dialysate circuit flow through a small dialyzer that utilizes polysulfone hollow fibers. The dialysis surface area of the dialyzer is about 0.2 meters. The blood circuit has an orifice for continuously adding heparin into the circuit before the dialyzer. The blood and dialysate will be pushed through their required circuits by a dual channel pulsatile pump powered by a replaceable battery. The dirty or used dialysate leaving the dialyzer will be circulated through a series of filter cartridges containing urease and sorbent similar to that described by Marantz and coworkers and widely used in the known REDY system. Ultrafiltrate is removed from the dialysate circuit by a valve arrangement. The removed ultrafiltrate is directed to and stored in a plastic bag which will be periodically emptied after the volume measurement. Sensors connected to the microprocessor monitor various aspects of the example apparatus.

Six farm-grown pigs (each weighing approximately 150lbs) were anesthetized and uremia was induced by surgical ligation of both ureters. After 24 to 48 hours, the animals were reananced, a double lumen Mahurkar catheter was inserted into the jugular vein, the catheter was connected to the exemplary CRRT device, and each animal was dialyzed for 8 hours. At the end of 8 hours animals were subjected to a gentle treatment (euhemized).

Blood samples were drawn from the arterial line inserted into the carotid artery and measured for CBC, urea, creatinine, sodium, potassium, chloride, CO₂Phosphorus, calcium and magnesium. The same chemicals are measured in the dialysate circuit on the input side of the dialyzer and on the output side of the respective filter cartridge.

The results of the test experiments are as follows. No adverse events were observed in the animals in the test trials. The average blood flow rate in the blood circuit was 44ml/min and the average dialysate flow rate was 73 ml/min. The results of the test trials are summarized in tables I and II.

Table I: the amount of fluid removed from each animal over 8 hours (in ml)

	Pig C (g)	Pig D (g)	Pig E (g)	Pig F (g)	Pig G (g)	Pig H (g)
							1 hour	400	100	100	100	150	180
2 hours	700	200	200	200	220	200
							3 hours		300	200	300	380	350
4 hours	800	400	250	400	500	700
							5 hours		500	300	500	600	710
6 hours		500	500	800	690	1410

7 hours		620	600	1000	700	1400
							8 hours		800	1000	1150	800	1400
Average	100	100	124	144	100	175

Table II: experimental data obtained from six pigs using an exemplary CRRT apparatus

	Creatinine clearance (ml/min)	Total creatinine (g) removed (8 hours)	Urea clearance (ml/min)	Total urea (g) removed (8 hours)	Std (Kt/V) Urea weekly	Phosphorus (g) (24 hours)	Potassium (m mol) (24 hours)
								Pig C	20.10	0.91	29.40	7.61	6.50	2.30	266.11
Pig D	21.10	0.76	26.80	5.75	6.20	2.60	259.91
								Pig E	23.50	1.14	27.30	5.37	6.10	2.67	303.54
Pig F	23.50	1.14	27.30	5.37	6.00	2.44	270.50
								Pig G	22.30	0.95	25.70	6.46	5.20	2.41	236.97
Pig H	22.30	1.02	26.30	6.24	5.80	2.42	277.01
Average	22.13± 1.34	0.99± 0.15	27.13± 1.27	6.13± 0.85	5.97± 0.44	2.47± 0.14	260.67± 27.05

The volume of fluid removed varied anywhere from 0 to about 700ml/hr over the course of the test. The limiting factor in removing larger amounts of fluid per hour is that the blood flow in the dialyzer gradually decreases as the fluid removal rate increases. The blood flow rate, directly normalized to the ultrafiltration (fluid removal) rate, will decrease. However, it is not difficult to keep the fluid removal at 100 ml/hr. The amounts of urea, creatinine, and phosphorus are further represented in fig. 18A through 18E. The amount of potassium and phosphorus removed represents 24 hours per treatment. The daily removal of potassium was 260.67. + -. 27.05 mmol/24 h. The daily removal of phosphorus was 2.47. + -. 0.14 g/24 h. The mean creatinine clearance obtained by this example is 22.13 ± 1.34 ml/min. The mean urea clearance was 27.13. + -. 1.27ml/min and the Kt/V of urea per week was 5.97. + -. 0.44.

The absence of complications in the test trial indicates that the example wearable CRRT device may operate without possible complications. The wearable CRRT devices of the example embodiments do not exhibit any complications that are different from those associated with existing large-scale dialysis machines used in the industry. The relatively low flow rates of the blood circuit and the dialysate circuit mitigate various complications that arise in some dialysis systems. The CRRT device of the example embodiments may be varied to increase blood flow rates to a range from about 50 to 120 ml/min. The variation includes at least one of: increasing the size of the dialyzer, increasing the flow of the double pump, and adjusting the pump gearing, delivery and valves.

The ability of the exemplary wearable CRRT device to stably remove fluid from the vascular space (in an amount similar to the volume of fluid physiologically removed by normal kidneys) enables the treating physician to maintain good patient (euvolomic) regardless of the amount of fluid ingested by the patient. Moreover, eliminating excess fluid may also result in better control of the patient's hypertension. The sodium concentration in the ultrafiltrate withdrawn is approximately equal to the sodium concentration in the plasma of the patient. Thus, removal of about 0.5 to 3 liters of ultrafiltrate a day by the exemplary CRRT device will result in removal of about 10 to 20 grams of salt per day. Removal of sodium from the patient's body by embodiments of the present invention may help to better control the patient's hypertension and also allow for a wider range of salt intake for ESRD patients. Therefore, the quality of life of the patient can be improved by increasing the types of foods that the patient can eat. Furthermore, eating various foods may result in improved nutrition for the patient.

Also, the removal of potassium and phosphorus from the patient's blood by the exemplary wearable CRRT device will further help eliminate oral intake restrictions on both elements, and no oral phosphate cement is required.

The test results show that the removal of creatinine and urea and the higher dialysis dose (expressed as two clearance rates and weekly urea Kt/V) make it possible to obtain all the advantages of the intermittent daily dialysis doses currently provided. At the same time, tests have demonstrated that the medical manpower used and other costs associated with long-term dialysis can be reduced.

Referring now to fig. 19, another example WAK1200 is shown incorporated into a fully wearable belt 1202. WAK1200 is a lightweight (between 0.5 and 5 pounds), belt-type, battery-operated 1204 WAK device 1200 that includes the following major components:

A. a dual channel pulsatile pump 1206 that pushes blood 1208 and dialysate 1210 through WAK 1200;

B. a high flow AN-69 dialyzer 1212 with a 0.6 square meter membrane surface (Hospal Industrie, france);

C. a dialysate regeneration system 1214 including 1 to 5 specially designed cartridges and/or filters containing adsorbents and a reservoir 1216 of electrolyte additives and heparin 1220 and a pH control circuit 1218;

E. auxiliary micro-pumps 1222, 1224 for feeding heparin, Mg, Ca, K and sodium bicarbonate and removing excess ultrafiltrate, all at a pre-specified flow rate; and

F. microprocessor-based control and monitoring devices (not expressly shown).

The dual channel pulsating pump 1206 uses a 3 watt DC micro motor (faulhahber, germany). The dual channel pulsatile pump 1206 has an oscillating mechanism and a dual channel flow cassette that enables simultaneous pulsatile flow of blood and dialysate at a controllable rate of 40-100 ml/min. When one channel pushes fluid out of its compression chamber (as the "systolic phase"), the other chamber fills its compressible chamber (as the "diastolic phase"), thereby creating a peak pressure in one channel and a pressure in the other channel at its valley or lower.

The example WAK of fig. 19 uses a compact larger diaphragm surface (0.60 square meters), high flow AN69Multiflow^TM60 dialyzers, instead of Hemophan for the test described above(0.22 square meter dialyzer). The pH meter electrodes 1220 are disposed in the dialysate circuit and are connected to a pH control system 1218. The pH control system 1218 is configured to initiate sodium bicarbonate infusion of the dialysate 1210 prior to the dialysate entering the dialyzer in order to maintain the pH of the dialysate at about 7.4. Urea and creatinine for this exampleClearance and standard weekly urea Kt/V were calculated as follows:

clearance ═ blood flow × [ Δ solute ]/[ solute ingress ] (1);

standard weekly Kt/V-effective clearance x time/total body fluid (2);

where [ Δ solute ] is the difference between the concentrations in and out of the dialyzer. The "time" in the actual test was 480 minutes and was extrapolated 10080 minutes per week. The "total body fluid" was estimated to be 60% of body weight. It should be noted that this calculation uses the average effective clearance averaged over the entire time period of the week.

Pump/dialyzer features

To test the example WAK1200 using different pumping devices, the following pumps were used: a dual channel pump (10 × 7 × 5cm, 380 g) 1206 of example WAK; minimump^TM(30 × 15 × 15cm, 4650 grams, minntech, MN), similar to the roller pump used in HD; MasterFlex(30X 18X 12cm, 2950 g, Cole-Parmer, IL), similar to the CRRT blood roller pump; and Profile^TM(8X 7cm, 65 g, Meikoset, Corp., Japan).

First, heparinized pig blood 1250 is dialyzed against isotonic saline as dialysate 1252 (fig. 20) using the aforementioned pump and the following dialyzers: hemophan(Model 100HG, Gambro, Germany), AN69(Multiflow 60, Hospital France) and F3(Fresenius, germany). Urea, creatinine and phosphorus were added to the blood to raise BUN to 60mg/dL, creatinine to 10mg/dL and K to 6m mol/L. These additions were based on calculating total blood and individual chemistryWeight of drug, and use of i-STATA portable clinical analyzer (i-STAT Corporation, NJ.) to record the actual components. Both fluids flowed in an open-ended loop at 37 ℃. Various pumps operate at a variety of speeds. After the flow at each speed was stabilized, the oscillation frequency, speed, flow rate and pressure provided by the pump in the circuit were recorded. Utilization of i-STAT in blood and dialysate samplesPortable clinical analyzers (i-STATcorporation, NJ) and 990HitachiThe analyzer was automated to analyze in duplicate. Gravity is used for continuous, pulse-free flow of blood and dialysate. All experiments were repeated 5 times.

Characteristics of the adsorbent

The features of the dialysate regeneration system 1214 of the example WAK were examined below, as shown in fig. 21. The adsorbents used include immobilized urease 1260, zirconium phosphate 1262, hydrous zirconium oxide 1264, and activated carbon 1266.

The spent dialysate 1266 was passed through the system without change and addition of urea, creatinine, potassium and calcium (to achieve 50mg/dL BUN, 4mg/dL creatinine, 4mEq/L K and 9mg/dL Ca). As the dialysate is circulated through the system, approximately 100ml/hr of the fluid is ultrafiltered. The dialysate was monitored for free ammonia. The solute clearance VS pump speed and flow rate of the dialysate are compared. All experiments were repeated 5 times.

Then, 3gm/L of urea/water solution was circulated through the combination of immobilized urease and zirconium phosphate using a continuous roller pump and a pulsating flow (WAK pump). The effluent was monitored for pH and ammonia. All experiments were repeated 5 times.

In vitro study

Additional studies and tests in vitro were performed using the example WAK 1200. All Animal studies were approved by Institutional Animal Care and Use Committee of Cedars SinaiHospital and according to the NIH guidelines for Animal testing. The ureters were surgically ligated into 5 anesthetized pigs. The next day, they were dialyzed through the example WAK1200 using a double lumen catheter for 8 hours.

Results

Flow/pressure waves generated by a pump

The effect of the pulsating nature of the different pumps described above is shown in figures 22 to 24. The dual channel pulsatile pump 1206 of the example WAK provides pulsatile flow of both blood and dialysate (fig. 22). The flows on the two separate channels are half a cycle out of phase.

The pump 1206 of the example WAK provides a higher amplitude and a much higher pulse frequency than commonly used roller pumps. FIG. 22 shows an exemplary pump 1200 with a stroke capacity of approximately 0.8ml, a frequency of 110 pulses per minute (bpm), a peak instantaneous flow of 350ml/min running blood, and an average of 95 ml/min. The corresponding data for roller pumps for the same blood flow are 8ml, 13bmp and 120ml/min, respectively (see FIG. 23A). Similarly, the corresponding data for the CRRT roller pump are 0.7ml, 125bpm and 140ml/min, respectively (see FIG. 23B).

The pressure waves generated by the example WAK pump 1206 can have a higher across-the-membrane pressure gradient (TMP) than the TMP generated by the roller pump (fig. 24A and 24B). The pressure waves in the combination of CRRT-roller pump moving blood and centrifugal pump moving dialysate are at the same average flow rate. Analysis of these figures showed that TMP varies from-10 to +150mmHg at the blood inlet port (fig. 24A) and from-50 to +50mmHg at the blood outlet port of the dialyzer (top of fig. 25A) by WAK pump. The corresponding values for the roller pump/centrifugal pump combinations were 0 to +60 and-10 to +20mmHg, respectively (fig. 24B and 25B).

Solute clearance through dialyzer

Another result of the testing and testing shows that the pulsating flow of the example dual channel pulsating pump 1206 produces a higher clearance rate than a continuous, steady, pulse-free flow. The data also shows that the handheld battery operated pump of the example WAK is as effective as the more heavy duty CRRT roller pump in producing solute clearance through commercially available dialyzers, even though commercially available dialyzers are designed for much higher blood and dialysate flow rates, typically in the 300-.

Adsorption efficiency by dialysate regeneration system

The adsorption efficiency of the stabilized, pulse-free flow using pulse flow VS generated by WAK and a conventional roller pump is shown in fig. 28. The amount of urea removed per gram of adsorbent is significantly higher in the case of pulsatile flow than in the case of non-pulsatile flow. This is an unexpected result. The effect of optimizing the dialysate pH to 7.4 is shown in figure 29. When the pH was optimized to 7.4, the amount of urea removed per gram of zirconium increased significantly. This is also an unexpected result.

Animal research

The results of two animal studies performed with the first example WAK (using a dialyzer with a smaller surface area, which is used for animal studies) are summarized in table III and compared with the results obtained with the other example WAK 1200.

Wearable artificial kidney

Dialysis in anesthetized uremic pigs for 8 hours

Results		WAK1200	Unit of
				Effective urea clearance	24.1±2.4	39.8±2.7	mL/min
Effective creatinine clearance	25.1±2.3	40.9±2.3	mL/min
				Total urea removed	12.4±2.8	15.3±4.4	g
Total creatinine removed	0.9±0.2	1.7±0.2	g
				Total phosphorus removal	0.8±0.2	N/A	g
Total potassium removed	80.5±19.5	150.5±16.7	m mol
				Extrapolation standard Kt/V	6.9±1.9	7.7±0.5

Table iii results of animal studies with two different exemplified WAKs

Conclusion

Flow/pressure performance

The blood and dialysate flowing in the example WAK or wearable CRRT device each oscillate at a frequency of 1.2-2.4 Hz. This half sine wave motion provides an alternating stroke volume of about 0.8ml in each channel. Therefore, the temperature of the molten metal is controlled,

V_m0.8 or V,/pi F_m＝2.5F (3)

Wherein, V_mIs the maximum instantaneous volume in the F Hz cycle. For a blood flow of 95ml/min and a period of 0.55 seconds, F is 1/0.55 seconds is 1.82Hz, and V is_m＝(2.5)(1.82) ＝4.55cm³。

HemophanThe dialyzer was used to estimate 1830 hollow fibers (100 micron lumen radius, 6.5 micron wall thickness, 18.5cm length). The cross-sectional area and the lumen volume of the blood flow were estimated to be 0.575cm each²And 10.6cm³. Therefore, the blood lumen volume is estimated to increase intermittently by (4.55) (100)/10.6-42.9%. For AN69Dialyzer, which will be 4390 hollow fibers (120 micron lumen radius, 50 micron wall thickness, 15cm length), thus blood flow cross-sectional area and lumen volume are 1.986cm each²And 29.8cm³. Thus, the blood lumen volume will increase intermittently by 15.3% to (4.55) (100)/29.8. These estimated volumetric oscillations are due to the fact that the capillaries are not elastic and the liquid is essentially incompressibleWill be almost completely converted into pressure waves. Using blood movement viscosity v is 0.03cm²Second for Hemophan(r ═ 0.010cm) and AN69(r-0.012 cm) dialyser, corresponding Womesley number r (2 π F/v)^1/20.195 and 0.235, respectively. The Poiseuille theorem cannot be used for the example WAK because the instantaneous lumen volume changes are large and the Womeresley number is not negligible. This is in contrast to the large dialysis machines currently in use, where the flow in the dialyzer is parabolic and therefore in a steady state. Thus, Poiseuille's theorem is considered applicable to ordinary machines, but not to the example WAK.

Example WAK simultaneously pumps a pulsatile flow of blood and dialysate (fig. 22). The two fluid streams are half-cycle out of phase, allowing the push-pull mechanism to traverse the dialyzer membrane and act on the sorbent particle surfaces. The blood flow wave generated by the example dual pulsatile pump has a higher amplitude and frequency than the blood flow wave generated by the roller pump. This advantageously allows a large difference in the amount and direction of TMP along the dialyzer fibers and radically changes the clearance and the mass transfer between the blood and the dialysate. It is clear that the push-pull mechanism created across the membrane by the example dual channel pulsatile pump can significantly improve the performance of the dialyzer.

The pressure waves generated by the example dual channel pulsatile pump represent a different and unique pattern, which is quite different from the pressure and TMP patterns seen in conventional dialysis machines. Through our experiments, a pulsatile profile of TMP (pressure across the membrane) was found between the blood and dialysate and time traces. In contrast, there is no such pulsatile change performance in current dialysis machines, where TMP is time independent. Also, we note that the pivot point (TMP ═ 0 when the blood and dialysate pressure lines cross) is greatly shifted to the right (i.e., convection from blood to dialysate primarily) by the exemplary pulsatile flow.

Solute clearance through dialyzer

Fig. 26A and 26B show that pulsating flow produced by WAK or a conventional roller pump results in a higher clearance than steady flow. Fig. 27 and 28 demonstrate that the clearance values for the example WAK pump are comparable to the clearance values produced by the roller pump at the same average flow rate. The former shows the results produced by AN69 dialyzer and the latter is F3A dialyzer.

Also, even though the example dual channel pulsatile pump provides the same effective solute clearance as current roller pumps, it operates efficiently with a smaller 9 volt battery, whereas a conventional roller pump requires connection to an electrical outlet that provides 110 or 220 volts. Also, the weight of an example dual channel pulsating pump may be between 300 and 450 grams, while a conventional roller pump would weigh 8 to 12 times.

Mass transfer performance

Any mass transfer in the system can be considered to be non-steady state due to the inherent fluctuations of the pulsating flow. This means that instead of using the usual Fick first law,

one will use Fick's second theorem, which in a simplified form is

For cylindrical geometries such as hollow fibers, and

for spherical geometries, for example in adsorbent particles.

In these equations, d andsimple and local difference symbols, respectively, N is the weight flux [ mass/area/time ]]And D is the diffusion coefficient [ length squared/time ]]C is the concentration [ mass/volume ]]And r is the (radial) diffusion distance [ length ]]. Fick's second theorem involves transient-convective-concentration changes. D is independent of the ambient or transient pressure in the liquid system, but it depends on the pore structure of the powder/particles in the adsorbent bed. The first part of equation (5) is the natural starting point for theoretical analysis of hollow fiber dialyzers, and the second part is the sorbent powder and particles used in the dialysate regeneration system. However, sorbent beds cause more complex mass transport due to ion exchange, permeation and filtration; they are designed empirically.

Dialyzer performance

Solute clearance rate: it is clear that the substance transfer in the dialyzer is increased by the pulsating flow. The reasons for increased mass transfer are believed to be greater fluid energy (i.e., greater average pressure), improved convective and diffusive mass transfer, and avoidance of molecular channeling and membrane delamination. It is also believed that a negative pressure gradient between dialysate pores enhances ultrafiltration.

And (3) ultrafiltration: the graphs produced jointly by the manufacturer of the dialysers represent equations (1) and (2), but do not include the effect of convection due to ultrafiltration. The effect was estimated to be at least 3% by continuous flow convective mass transfer as shown below.

Clearance-clearance at zero UF +0.46 × UF (6)

In our previous animal study with the first test model of example WAK, mean clearance and ultrafiltration were 24.6 and 1.67ml/min (100ml/hr), respectively. Thus, a clearance of 1.67ml/min for UF and 0.46 × UF of 0.768ml/min, i.e. 0.768/24.6 of 0.031 or 3.1%.

This would be the case when there is only continuous flow convection. However, due to the pulsating nature of the flow, a much higher convective mass transfer percentage can be expected. Intermittent changes in the direction of the dialyzer TMP gradient indicate that the example CRRT device can actually perform hemodiafiltration, as these changes will clearly result in bi-directional flow transport of water and solutes across the membrane.

Adsorption efficiency in adsorbents

An exemplary sorbent system using a zirconium phosphate cation exchange structure comprises zirconium, phosphorus, oxygen, nitrogen, and hydrogen, and immobilized urease/alumina powder for decomposing urea. The amount of adsorbent used until saturation was reached was measured using different amounts of immobilized urease (typically 35-80 microns, 1.22gm/ml loss) and zirconium phosphate (typically 25-45 microns, 1.22gm/ml loss). Using a non-pulsating or pulsating (WAK pump) flow, the average effective velocity (flow rate divided by cross-sectional area) is the optimum parameter related to the amount of urea pumped per fixed amount of adsorbent. Figure 29 shows that pulsating flow increases the amount of urea extracted per gram of zirconium phosphate.

We have found that when the pH of the dialysate is optimised to 7.4, the result is an increase in adsorbent capacity. This can be explained by the massive adsorption of sodium and by Na⁺To H⁺Early replacement/release. This would explain lower adsorption early at lower pH, since acidity can significantly reduce ammonia adsorption.

Animal research

The results of the initial animal study by the first example of WAK are summarized in table 2 and compared with the results obtained for the second example of WAK. Creatinine clearance increased from 25.1 + -2.3 to 40.9 + -2.3 ml/min, and urea clearance increased from 24.1 + -2.4 to 39.8 + -2.7 ml/min. The extrapolated weekly Kt/V increases from 6.9 + -1.9 through the first example WAK to 7.7 + -0.5 through the second example WAK.

This clearance is about 2 higher than the usual values for ordinary dialysis machines, even for daily dialysis, the reported value is close to 6 higher. Therefore, the second example WAK is defined as being on the straight line std (Kt/V) of Gotch of 8 × sp (Kt/V).

From the above experiments and tests, it has been found that, unlike the hypothesis, pulsatile flow outperforms stable or continuous flow in solute-producing mass transfer and clearance. The example WAK dual channel pulsatile pump produces higher peak blood flow rates and clearance rates similar to current pumps that are too heavy, too energy inefficient, and too bulky to be part of a wearable device. In contrast, the example WAK pump is lighter and its smaller size and geometry (see example dimensions of fig. 14 and 15) can be worn. The change in TMP and higher amplitude of pulsations during the pulsation of the example dual channel pulsating pump will produce very high mass transfer rates, primarily by convection. Thus, the example WAK provides hemodiafiltration, rather than pure hemodialysis. The half-cycle difference between the blood and dialysate pulsations, the amplitude, and the varying frequency of the TMP (produced by the example dual channel pulsatile pump) produce a "push-pull" flow through the dialyzer membrane. The pulsating flow of the example dual channel pump increases the superficial velocity of the dialysate in contact with the sorbent, thereby further increasing the amount of solute adsorbed by the sorbent. Maintaining the dialysate pH at 7.4 improves the adsorption efficiency compared to not maintaining the dialysate pH. The high flux membrane and larger dialyzer surface improve the performance of the example WAK without adding bulk or weight. The results further demonstrate that embodiments of the present invention may be selected to optimize ESRD patient care to provide 168 hours of weekly high clearance dialysis. Current logistical requirements for daily intermittent dialysis can also be eliminated and logistical problems simplified, with hopes for lower cost, improved quality of life and reduced mortality for ESRD patients.

Many variations and embodiments of the above described invention are possible. While particular embodiments of the present invention have been illustrated in the accompanying drawings and described in the foregoing detailed description, it will be understood that the invention is not limited to the embodiments described, but is capable of additional modifications, changes and substitutions without departing from the scope of the invention as set forth and defined by the following claims. It is therefore to be understood that the scope of the present invention encompasses all such structures and is limited only by the following claims.

Claims (14)

1. A dual channel pulsatile pump comprising:

a dual channel pump cassette having a first channel and a second channel, each channel of said dual channel pump cassette comprising:

a compressible peristaltic tube having an input end and an output end;

an input valve at an input, the input valve comprising: an input O-ring; an input ball abutting the input O-ring to prevent backflow of fluid through the input valve; and an input spring member that positions the input ball against the input O-ring when peristaltic tubing is compressed, the input spring member allowing the input ball to exit the input O-ring and allow fluid to flow forward through the input valve when the peristaltic tubing is allowed to decompress, the input spring member comprising:

an outer ring having a first diameter;

an inner ring having a second diameter less than the first diameter; and

at least two but not more than three spokes extending radially from the inner ring to the outer ring; the inner ring being laterally movable relative to a plane extending perpendicular to the axis of the outer ring; and

an output valve at the output, the output valve comprising: an output O-ring; an output ball abutting the output O-ring to prevent backflow of fluid through the output valve; and an output spring member that positions said output ball against said output O-ring when peristaltic tubing is decompressed, said output spring member allowing said output ball to exit said output O-ring and allow fluid to flow forward through said output valve when said peristaltic tubing is compressed; and

a pump motor part including: a first oscillatory pushing member for compressing and decompressing the peristaltic tube in the first channel; and a second oscillatory pushing member for compressing and decompressing the peristaltic tube in a second channel, the dual channel pump cassette being removably mounted on the pump motor portion.

2. The dual channel pulsatile pump of claim 1 wherein: the dual channel pump cassette dimensions are less than 0.72 inch high, 3.82 inches long, and 1.7 inches wide, with each measurement plus or minus 50 percent.

3. The dual channel pulsatile pump of claim 1 wherein: the double-channel pulsating pump has the size of 9.7 multiplied by 7.1 multiplied by 4.6cm and the weight of less than 400 grams plus or minus 50 percent.

4. The dual channel pulsatile pump of claim 1 wherein: the inner ring is a stabilizer for the ball when fluid flows through the input valve.

5. A fully Wearable Artificial Kidney (WAK) device, comprising:

a blood circuit;

a dialysate circuit;

a dual channel pulsatile pump comprising:

a first channel for pumping blood through the blood circuit and a second channel for pumping dialysate through the dialysate circuit, each of the first and second channels comprising:

a compressible peristaltic tube having an input end and an output end;

an input valve at an input, the input valve comprising: an input O-ring; an input ball abutting the input O-ring to prevent backflow of fluid through the input valve; and an input spring member that positions the input ball against the input O-ring when peristaltic tubing is compressed, the input spring member allowing the input ball to exit the input O-ring and allow fluid to flow forward through the input valve when the peristaltic tubing is allowed to decompress, the input spring member comprising:

an outer ring having a first diameter;

an inner ring having a second diameter less than the first diameter; and

at least two but not more than three spokes extending radially from the inner ring to the outer ring; the inner ring being laterally movable relative to a plane extending perpendicular to the axis of the outer ring; and

an output valve at the output, the output valve comprising: an output O-ring; an output ball abutting the output O-ring to prevent backflow of fluid through the output valve; and an output spring member that positions said output ball against said output O-ring when peristaltic tubing is decompressed, said output spring member allowing said output ball to exit said output O-ring and allow fluid to flow forward through said output valve when said peristaltic tubing is compressed; and

a single electric motor and transmission that alternately compresses and decompresses peristaltic tubing in said first channel and peristaltic tubing in said second channel; the dialysate circuit, the blood circuit, and the dual channel pulsatile pump are all worn entirely by the user.

6. A fully Wearable Artificial Kidney (WAK) device as claimed in claim 5, wherein: the dual channel pulsatile pump provides a blood flow rate through the blood circuit of between 15 and 100ml/min pulsatility.

7. A fully Wearable Artificial Kidney (WAK) device as claimed in claim 5, wherein: the dual channel pulsating pump provides alternating pulses in the first and second channels that are 180 degrees out of phase.

8. A fully Wearable Artificial Kidney (WAK) device as claimed in claim 5, wherein: the dual channel pulsatile pump is capable of pumping from 0 to 350ml/min of blood in the blood circuit.

9. A fully Wearable Artificial Kidney (WAK) device as claimed in claim 5, wherein: the dual channel pulsatile pump is capable of pumping from 0 to 350ml/min of dialysate in the dialysate circuit.

10. A fully Wearable Artificial Kidney (WAK) device. The method comprises the following steps:

a blood circuit;

a dialysate circuit;

a dual channel pulsatile pump comprising:

a first channel for pumping blood through the blood circuit and a second channel for pumping dialysate through the dialysate circuit, the first and second channels providing out-of-phase pump circulation; the first channel includes:

a compressible peristaltic tube having an input end and an output end;

an input valve at an input, the input valve comprising: inputting a ball seat; an input ball abutting the input ball seat to prevent backflow of fluid through the input valve; and an input spring member that positions the input ball against the input ball seat when the peristaltic tubing is compressed, the input spring member comprising:

an outer ring having a first diameter;

an inner ring having a second diameter less than the first diameter; and

at least two but not more than three spokes extending radially from the inner ring to the outer ring; the inner ring being laterally movable relative to a plane extending perpendicular to the axis of the outer ring; when the peristaltic tubing is capable of being decompressed, the input spring member allows the input ball to exit the input ball seat and fluid to flow forward through the input valve; and

an output valve at the output, the output valve comprising: an output O-ring; an output ball abutting the output O-ring to prevent backflow of fluid through the output valve; and an output spring member that positions the output ball against the output O-ring when peristaltic tubing is decompressed, the output spring member allowing the output ball to exit the output O-ring and allow fluid to flow forward through the output valve when the peristaltic tubing is compressed.

11. A fully Wearable Artificial Kidney (WAK) device as in claim 10, wherein: the first and second pump cycles are 180 degrees out of phase.

12. A fully Wearable Artificial Kidney (WAK) device as in claim 10, further comprising: a dialyzer in the blood circuit and in the dialysate circuit such that blood and dialysate pass through the dialyzer in opposite directions.

13. A fully Wearable Artificial Kidney (WAK) device as in claim 10, further comprising: a pH control system that maintains the pH of the dialysate at a predetermined pH.

14. A fully Wearable Artificial Kidney (WAK) device as in claim 10, wherein: the fully wearable artificial kidney device provides hemodiafiltration, while not purely providing hemodialysis.