Enhanced clearance in an artificial kidney incorporating a pulsatile pump

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(12) Patent Specification do SE 534 780 C2

Sweden (21) Patent application number: 0950468-9 (45) Patent granted: 2011-12-20 (41) Application generally available: 2009-08-11 (22) Patent application received: 2007-11-19 (24) Maturity date: 2007- 11-19

Completed international patent application with number: PCT / US2007 / 085131 (86) International filing date: 2007-11-19 (83) Deposit of microorganism: - (30) Priority information: 2006-11-17 US 60/866357 2007-11-19 US 11/942626 (51) International class:

A61M1 / 16 (2006.01)

B01D 61/32 (2006.01)

Waltham MA 02451 (73) Patent Holders:

Fresenius Medical Care Holdings, Inc., 920 Winter Street, US

<td>(72) Inventor:</td><td>Edmond RAMBOD, Los Angeles US Victor GURA, Beverly Hills US</td>

<td>(74) Agents:</td><td>Albihns.Zacco AB, Box 5581, 114 85 Stockholm SE</td>

<td>(54) Name:</td><td>Purification in an artificial kidney containing a pulsatory pump</td>

<td>(56) Publications cited:</td><td> ___</td>

(47) Summary:

A continuous renal replacement therapy (CRRT) device is provided which weighs between 2 and 10 pounds. The CRRT device may be portable, mobile or fully supported on the patient. Blood and dialysate are each pumped in a pulsating or pulsatory manner through a dialyzer such that a substantial portion of the peak pulse of the blood flow coincides with a substantial portion of the low pulse portion of the dialysate flow. A pressure difference between a dialysate inlet of the dialyzer and the blood inlet of the dialyzer changes periodically from a high pressure difference of between 70 and 120 mmHg for a first period of time and a low pressure difference of -10 to 10 mmHg for a second period of time. The frequency of the high and low pressure differential cycles is between about 0.5 and 4 Hz.

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* -102

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SUMMARY

A continuous renal replacement therapy (CRRT) device is provided which weighs between 2 and 10 pounds. The CRRT device may be portable, mobile or fully supported on the patient. Blood and dialysate are each pumped in a pulsating or pulsatory manner through a dialyzer such that a substantial portion of the peak pulse of the blood flow coincides with a substantial portion of the low pulse portion of the dialysate flow.

A pressure difference between a dialysate inlet of the dialyzer and the blood inlet of the dialyzer changes periodically from a high pressure difference of between 70 and 120 mmHg for a first period of time and a low pressure difference of -10 to 10 mmHg for a second period of time. The frequency of the high and low pressure differential cycles is between about 0.5 and 4 Hz.

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CLEANING IN AN ARTIFICIAL NJURE CONTENTS

PULSATORY PUMP

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a PCT application from patent application US 11/942,626, filed November 19, 2007, titled IMPROVED CLEANING IN AN ARTIFICIAL NJURE CONTAINING A PULSATORY PUMP (Atty. Dkt. No. NQCR-28505), which is a continuation in -part of ongoing patent application US 10/940,862, filed September 14, 2004, entitled SUSTAINABLE DEVICE FOR CONTINUOUS NURSERY REPLACEMENT THERAPY (Atty. Dkt. No. NQCR-27527), which is a continuation in-part of U.S. Patent Application No. 10/085,349, filed November 16, 2001, entitled PORTABLE NURSING REPLACEMENT THERAPY (Atty. No. NQCR-27513), all of which hereby inserted as a reference. This application further requests priority from the ongoing provisional application US 60/866,357, filed November 17, 2006, entitled IMPROVED CLEANING IN AN ARTIFICIAL NJURE CONTAINING A PULSATORY PUMP (Atty. Dkt. No. NQCR-28 079), which is also incorporated herein by reference.

FIELD

The present invention is directed to dialysis systems, and more particularly to dialysis systems which may be fully and continuously supported by a patient or dialysis system which is portable or mobile.

BACKGROUND

Hemodialysis is a method by which microscopic toxins are removed from the blood using a filter membrane such as a dialyzer. Typically, hemodialysis is administered to a patient in periodic three- to four-hour sessions. Each session takes place two or three times a week. There is a growing amount of research that prefers daily dialysis as increased dialysis time improves results, both in terms of quality of life and patient life. An additional number

534 780 researchers believe that continuous dialysis for twenty-four hours a day, seven days a week would provide the best results for a patient in terms of quality of life and longevity. The actual realization of substantially continuous dialysis has been impossible until today due to technology and cost restrictions. Regardless, continuous renal replacement therapy (CRRT) is thought to be a huge improvement over intermittent dialysis as many more toxins can be removed from the blood using a CRRT device seven days a week and nearly twenty-four hours a day.

Some benefits of providing CRRT include an expected reduction in morbidity and mortality, a reduction in the amount of drug required, a decrease in fluid intake, a reduction in dietary restrictions, and several improvements in quality of life for patients with terminal phase renal failure (ESRD). Currently, CRRT machines are stationary, large, heavy machines adapted to provide dialysis, hemofiltration or a combination of both to individual patients. The existing CRRT is cumbersome and requires electrical connection to 120-140 volt AC power outlets as well as several feet of pipeline to connect the machine to the patient. In addition, these machines require a continuous supply of gallons (1 gallon = 3,785 liters) of fresh filtrate water to create the dialysate liquid. Furthermore, a patient must remain connected to the existing heavy and cumbersome CRRT machine for many hours each day, limiting his or her ability to perform normal everyday activities.

A further problem with existing dialysis machines is that frequent reconnection and disconnection to the machine requires access to blood flow in a patient by piercing an artiovenous shunt. These shunts only last for limited periods of time and expose the patient to infection, coagulation and other complications that result in multiple hospital stays and repeated surgical interventions. Another problem with existing dialysis machines is when these machines become smaller and a bit more portable, smaller hemofilters or dialyzer filters must be used that do not clog or coagulate too quickly so that extended or continuous dialysis can be performed. A common type of dialyzer includes nine hundred or more cylindrical hollow fibers through which blood flow is provided. The hundreds of cylindrical hollow fibers are contained in a shell

534 780 or container in which the dialysate liquid is circulated around and past the outer walls of the hollow fibers. The outer walls of the hollow fibers or lumens are semi-porous so that impurities in the blood can be removed from the blood and into the dialysate. A problem that arises in a dialyzer is clogging or coagulation of blood flow within the individual hollow fibers. Such clogging of blood flow through the fibers reduces the efficiency of the dialyzer's filtration and blood purification properties. Furthermore, it should be understood that proteins and other compounds or substances in the blood can clog the pores of the semiporous membrane over time and decrease the efficiency of the dialyzer filter. If a dialyzer filter is to be in continuous operation twenty-four hours per day seven days a week, it is important that such a dialyzer be operable for extended periods of time at or near a continuous maximum performance without being clogged or substantially reduced during use. Furthermore, it would also be useful if a dialyzer is maintained efficient and powerful when a low power pump is used to pump blood thereby requiring a minimal amount of energy for the highest possible purification of a patient's blood at the lowest amount of energy required.

Dialysis membranes have been studied for well over half a century. The early inventors of dialysis or dialactic therapy understand the basics of diffusion and how toxins diffuse across a diaphragm's membrane from blood to a dialysate fluid. There are many factors that affect diffusion in the transfer of solute across a semipermeable membrane. Such factors have been explained in various previous articles on the work of a dialyzer. But again, there has been limited or minimal research to provide a dialysis device in which the dialyzer operates effectively over extended periods of time (more than 15 hours) to provide a fully portable or portable low power dialyzer with a dialyzer with increased efficiency over that of previous dialysis machines. with a dialyzer having a membrane having the same or similar membrane surface area.

SUMMARY

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Embodiments of the invention provide a continuous kidney replacement therapy (CRRT) device weighing between 2 and 10 pounds (between 0.9 and 4.5 kg). The CRRT device may be portable, mobile or fully supported on the patient. Blood and dialysate are each pumped in a pulsed (pulsatile) or pulsatile way through a dialyzer such that a substantial portion of the maximum pulse of blood flow and pressure coincides with a substantial portion of a low pulse flow and pressure portion of dialysate flow. . A pressure difference between the dialyzer dialysate inlet and the dialysator blood inlet changes periodically from a high pressure difference of between 70 and 120 mmHg (between 9333 and 15999 Pa) for a first period of time and a low pressure difference of between -10 and 10 mmHg (between -1333 and 1333 Pa) for a second period of time. The frequency of the high and low pressure differential cycle is between about 0.5 and 4 Hz.

In one aspect thereof, a typical embodiment of the invention provides a continuous renal replacement therapy (CRRT) device. The CRRT device includes a blood pump channel that establishes a pulsating or pulsatory blood flow. The pulsed blood flow is periodic so that each period of the blood flow comprises a portion of high blood pressure with a first duration and a portion of low blood pressure with a second duration. The first duration and the second duration have a duration ratio of about 3: 4 to 4: 3. The duration ratio requires that the duration durations of the high blood pressure portion and the low blood pressure portion be at least 75 percent (%) of each other. For example, if the high blood pressure portion of the pulse has a duration of one (1) second, then the low blood pressure portion should have a duration of between about 0.75 and 1.25 seconds. The typical embodiment further comprises a dialysate pump channel providing a pulsating or pulsatory dialysate flow. The pulsed dialysate flow is periodic so that each period comprises a portion of high dialysate pressure of the second duration and a portion of low dialysate pressure of the first duration. The high blood pressure portion and the low dialysate pressure portion occur, at least in part, during the same first periodic time frame. The typical embodiment also includes a dialyzer or hemofilter. The dialyzer comprises a blood inlet that receives the pulsating blood flow from the blood pump channel. The dialyzer of the typical embodiment also comprises a plurality of fibers. Each fiber comprises an exterior semi-permeable

534 780 membranes and a lumen extending along the length of the fiber. The lumen provides a passage through the fiber for the pulsed blood flow to flow through. The dialyser further comprises a blood outlet for the pulsed blood flow to exit through the dialyzer as well as a dialysate inlet to receive the pulsed dialysate flow pumped by the dialysate pump channel. The dialyser also has an outer tube which establishes a dialysate container or chamber, about the majority of fibers. In other words, the fibers are substantially contained within the dialyser and are within a dialysate chamber or container. The pulsed dialysate flow flows through the dialysate chamber. The semipermeable membrane of most fibers is the medium between each lumen and dialysate chamber. The dialyser includes a dialysate outlet for the pulsed dialysate flow to exit the dialyser. The combination of the blood pump channel, dialysate channel and dialyser is designed to establish a maximum transmembrane pressure (TMP) across the semipermeable membranes of the plurality of fibers and between the pulsed blood flow in the lumen and the pulsed dialysate flow in the dialysate chamber. The maximum TMP occurs during the first periodic time frame. The maximum TMP is between about 70 mmHg (9333 Pa) and 120 mmHg (15999 Pa).

Different embodiments of a typical CRRT device require that the low blood pressure portion of the blood flow and the high dialysate pressure dialysate flow portion both occur, at least in part, during a second periodic time frame. Furthermore, the blood channel, dialysate channel and dialyser are constructed to establish a minimal TMP across the semipermeable membranes of the majority of fibers during the second periodic time frame. The minimum TMP is between about 10 mmHg (1333 Pa) and -10 mmHg (-1333 Pa).

Additionally, in various embodiments of a typical CRRT device, a two-channel ventricle pulsatile pump is included which includes the blood pump channel and the dialysate pump channel in the device. The same mechanical mechanism can be utilized in the two channel pulsatory chamber pump to affect both the blood pump channel and the dialysate channel.

Furthermore, some embodiments of the invention can be fully carried on the personal user, while other embodiments of the CRRT device are of light weight.

534 780 weight (eg, between 2 and 15 pounds (between 0.9 and 6.8 kg) while in operation) and mobile so that they can be moved around a medical device or in a user's home.

Some embodiments of a typical CRRT device provide a blood pump channel which can provide a pulsating or pulsatory blood flow with a periodic flow rate of between 0.5 and 4 Hz and wherein the dialysate pump channel provides a pulsating or pulsatory dialysate flow with the same or similar periodic flow. This periodic flow rate of blood has been found to be the least detrimental to the blood cells as they pass through the blood pumping duct and also prevents clogging or coagulation of the blood in the pumping duct, the fiber lumen of the dialyzer, as well as in the semipermeable membranes between the fiber lumen and the dialysate of the dialysate chamber portion. The pulsating or pulsatory blood flow described herein is found to provide an unexpected wash or push-pull effect that helps prevent clogging for the extended operating periods of typical CRRT devices. It should be noted that the elements of typical CRRT devices interact in an unexpected manner so that the maximum TMP occurring during the first periodic time frame and the minimal TMP occurring during the second periodic time frame prevent clogging of the lumen and the semipermeable membranes.

Another example of an embodiment of a continuous renal replacement therapy (CRRT) device is provided which includes a blood pump channel to provide a pulsatory blood flow. The pulsatory blood flow is periodic so that each period comprises a portion of high blood pressure with a first duration and a portion of low blood pressure with a second duration. The first duration and the second duration have a duration ratio of between about 3: 4 to about 4: 3. Of course, this includes a 1: 1 duration ratio. These typical embodiments further include a dialysate pump channel which provides a pulsatory dialysate flow. The pulsatory dialysate flow is a periodic flow such that each period of the flow comprises a portion of high dialysate pressure of the second duration and a portion of low dialysate pressure of the first

534 780 duration; wherein the portion of high blood pressure and portion of low dialysate pressure occurs periodically, at least in part, during the same first periodic time frame. The typical embodiment further includes a dialyzer. The dialyzer includes a blood inlet to receive the pulsatory blood flow. The dialyzer also comprises a plurality of fibers wherein each fiber comprises a lumen extending along the length of the fiber and an outer semipermeable membrane. The lumen is for carrying the pulsatory blood flow through a substantial portion of the dialyzer. The dialyzer further comprises a blood outlet for the pulsatory blood flow to exit the dialyzer, a dialysate inlet for receiving the pulsatory dialysate flow, a dialysate division area which is around or about the plurality of fibers, for the pulsatory dialysate flow to flow through; and a dialysate outlet for the pulsatory dialysate flow to exit the dialyzer. The association of the blood pump duct, dialysate duct and dialyzer is designed to establish a maximum pressure differential of blood inlet-to-dialysis outlet that occurs during the first periodic time frame. The maximum pressure difference of blood inlet dialysate outlet can be in the range of between about 60 mmHg (7999 Pa) to about 150 mmHg (19998 Pa), which provides an unexpectedly elevated amount of toxins from a patient's blood to move across the semipermeable membrane of the fiber into the the dialysate contained in and pulsed through the dialyzer.

Further embodiments of a typical CRRT device are designed such that the low blood pressure portion and the high dialysate pressure portion both occur, at least in part, during a second periodic time frame. The blood pump channel, dialysate channel, and dialyzer are designed to establish a minimal blood inlet to dialysate outlet pressure that occurs during the second periodic time frame and establishes a pressure difference between about 10 mmHg (1333 Pa) and about -10 mmHg (1333 Pa).

Other embodiments of the invention provide a method of continuous renal replacement therapy (CRRT). An example of a method for providing CRRT includes pumping blood, through a pulsatory blood pump, to provide a pulsatory blood flow. The pulsatory blood flow is a periodic flow wherein each period comprises a portion of high blood pressure with a first duration and a portion of low blood pressure with a second duration. The first

534 The 780 time duration and the second time duration have a duration ratio which may be in the range of about 3: 4 to 4: 3. The process of providing CRRT further comprises pumping dialysates, with a pulsatory dialysate pump, to provide a pulsatory dialysate flow. The pulsatory dialysate flow, as well as the pulsatory blood flow, has a period which includes a high dialysate pressure portion with the second duration duration and a low dialysate pressure portion with the first duration duration. The high blood pressure portion and the low dialysate pressure portion occur, at least in part, during a first periodic time frame. The typical process further establishes a pressure difference of blood inlet to dialyzer outlet - between a blood inlet of a dialyzer and a dialysate outlet of the dialyzer; wherein the pressure difference of blood inlet-to-dialysate outlet oscillates between a maximum pressure difference and a minimum pressure difference at a pumping frequency of between 0.5 and 4 Hz. The maximum pressure difference has a range of between about 60 mmHg (7999 Pa) and about 150 mmHg (19998 Pa). The minimum pressure difference has a range of between about 10 mmHg (1333 Pa) to about -10 mmHg (-1333 Pa) for the pump frequency of between 0.5 and 4 Hz.

Although two separate pulsatory pumps can be used in a typical CRRT procedure, different embodiments of the pulsatory blood pump and pulsatory dialysate pump are each a separate pump channel of a two channel pulsatory pump and each is operated by the same mechanical mechanism. Embodiments of the invention utilize a pulsatory blood pump and a pulsatory dialysate pump which provide a pulsatory average flow rate of between 30 and 90 milliliters / minute (ml / min).

In yet another embodiment of the invention, there is provided a method of dialysis which comprises receiving blood into a blood circuit of a dialysis device; pumping the blood in a pulsating manner to provide an average rate of pulsatory blood flow of between 30 and 90 ml / min; pumping the dialysate, in a dialysate circuit of the dialysis device, in a pulsating manner to provide a pulsatory dialysate flow having an average velocity of between 30 and 90 ml / min; receiving the pumped blood at a blood inlet of a dialyzer; receiving the pumped dialysate at a dialysate inlet of the dialyzer, to establish a pressure difference between the blood in the blood inlet and the dialysate in the dialysate inlet which fluctuates at a

534 780 periodic rate of between 0 mmHg (0 Pa) +/- 10 mmHg (1333 Pa) and 120 mmHg (15999 Pa) +/- 20 mmHg (2666 Pa); transfer urea molecules from the blood to the dialysate while the blood and dialysate pass through the dialyzer; and purifying the dialysate through a filtering agent so that the purified dialysate can be reused and recycled in the dialysate circuit.

Process also includes purification of the dialysate which further comprises regenerating the dialysate by adding predetermined additives to the dialysate.

The periodic speed can be between 0.5 and 4 Hz.

Furthermore, the method can be provided with at least one of a fully portable and portable dialysis device.

According to the method, pumping the blood may further comprise pumping the blood with a first channel of a two channel pulsatory pump and pumping the dialysate further comprising pumping the dialysate with a second channel of a two channel pulsatory pump.

Furthermore, the two-channel pulsatory pump can pump the blood and pump the dialysate so that a maximum flow of blood occurs in an alternating pattern with a maximum flow of the dialysate.

Pumping of the blood may take place in a pulsating manner comprising providing in an alternating manner a portion having a high blood flow and a portion having a low blood flow, and wherein pumping of the dialysate is effected in a pulsating manner comprising providing in an alternating manner a portion having high dialysate flow rate and a low dialysate flow rate portion, high blood flow rate portion and low dialysate flow rate portion each occur; at least for a majority of their durations, at the same time.

Embodiments of the invention provide a method for CRRT which can be performed by either a fully portable or a fully porous dialysis device.

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Various embodiments of the invention provide one or more pumping devices which pump the blood and dialysate so that a maximum flow of blood occurs at alternating times with a maximum flow of the dialysate. And various embodiments of the invention are designed so that the pumping of the blood in a pulsating manner comprises providing, in an alternating manner, a high blood flow rate and a low blood flow rate portion, and wherein the pumping of the dialysate in a pulsating manner comprises providing, in an alternating manner, a portion having a high dialysate flow rate and a portion having a low dialysate flow rate. The high blood flow rate portion and the low dialysate flow rate portion occur each, at least for a major portion of their duration, at the same time.

As such, various embodiments of typical CRRT procedures and devices have been found to function to provide elevated purifications in an artificial kidney when a pulsating or pulsatory pump is used in combination with other elements discussed herein to provide an unexpected solution to a problem encountered by renal failure patients but which has never been effectively resolved to provide a fully portable or portable CRRT device that is operable to enhance the quality of life of such patients and unexpectedly resists the blockage or coagulation of blood in its dialyzer elements.

BRIEF DESCRIPTION OF THE DRAWINGS

For a more complete understanding, reference is now made to the following description to be read in conjunction with the following figures, in which:

Fig. 1 illustrates an example diagram of a portable or fully portable dialysis device according to an embodiment of the invention;

Fig. 2 illustrates a basic structure of a two-channel pulsatory pump in conjunction with a dialyzer according to various embodiments of the invention;

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Fig. 3 illustrates a test structure using a roller pump for pumping dialysates and a roller pump for pumping blood through the fibers of a dialyzer;

Fig. 4 illustrates a simple hollow fiber and adjacent dialysates in an exemplary dialysate chamber of a dialyzer according to an embodiment of the invention;

Figures 5A and 5B provide samples of experimental result measurements of: a) input and output dialyzer pressures and b) input and output dialyzer flow rates for a combination roller pump and centrifugal pump assembly used in an experimental dialysis device for comparison.

Figures 6A and 6B provide experimental result measurements for samples of: a) inlet and outlet pressures of a dialyzer and b) inlet and outlet flow rates of a two-channel pulsator pump used according to an embodiment of the invention;

Fig. 7 depicts a graph showing experimental result measurements of the ratio of ultrafiltration to a designed transmembrane pressure (TMP) across a fiber within a dialyzer for both roller and pulsatory (shuttle) pump structures;

Fig. 8 depicts a table providing a comparison of examples of experimental results for a roller pump structure and an example of a pulsator pump structure.

Figures 9A and 9B provide graphs illustrating experimentally calculated and estimated blood flow, dialysate flow and transmembrane pressure over the axial length of a roller pump structure and a pulsator or pulse pump structure;

Fig. 10 provides a graph illustrating experimentally calculated and estimated ratios of convective urea flow across a semipermeable membrane in a roller pump structure and in an example of a shuttle pump or pulsator pump structure;

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Fig. 11 provides a table providing experimental and numerical results for urea purification with a roller pump and an example of pulsating pump design for various blood and dialysate fluid flow rates;

Fig. 12 provides a table providing experimental and numerical results for creatinine purification with a roller pump and an example of the design of pulsating pump for different blood and dialysate fluid flow rates;

Fig. 13 provides a table providing experimental and numerical results for potassium purification with a roller pump and an example of the design of pulsating pump for different blood and dialysate fluid flow rates; and

Figures 14A and 14B provide a graph of the relative concentration distribution C / C_O along the axial length of an example dialyzer inlet and blood outlet in a roller pump structure and an example of a shuttle pump or pulsator pump structure.

Fig. 15 is a graph of an example of blood and dialysate pressure or flow over time produced by a single two-channel pulsator pump or two synchronized pulsator pumps.

DETAILED DESCRIPTION

Referring now to the figures, wherein like reference numerals are used herein to denote like elements herein, the various views and embodiments of examples of improved purification in artificial kidneys containing a pulsatory pump are illustrated and described and other possible embodiments are described. The figures are not necessarily drawn to scale, and in some cases the figures have been enlarged and / or simplified in some places for illustrative purposes only. A person skilled in the art will appreciate the many possible applications and variations based on the following examples of possible embodiments.

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A typical embodiment of a portable CRRT device is depicted in Figure 1. Portable CRRT device 100 is built into, or is part of, a patient's portable belt, belt or other portable device 102. The belt or belt 102 may include a pair of end portions 104 which are adapted to be secured together with a fastener (not specifically shown). The end portion / fastener 104 may be any number of fasteners suitable for securing the ends of the waist belt or shoulder strap together over the patient's shoulder. Such a fastener is not limited to push buttons, buttons, buckles, clamps, hooks and loops, zippers, seals, buttons, etc. An embodiment of a typical CRRT device 100 can be imagined to be in the form of an ammunition or military supply belt. It can also be in the form of a vest or waist bag. Other embodiments may be in the form of a small backpack. An example of portable CRRT device 100 is to be worn continuously by a patient either above or below the patient's clothing. Other embodiments are portable devices substantially contained in a carrying case, briefcase, handbag or a backpack style container.

One or more circuit boards or microcircuits 106 are contained in the portable CRRT device 100. A microcontroller 108 is used to control and control various aspects of the portable CRRT device 100. The microcontroller 108 is preferably a low or very low power microcontroller, but can be substantially any microcontroller adapted to operate in a typical CRRT device 100. One of the many functions of microcontroller 108 is to control the battery 110. An example of CRRT device 100 will operate continuously for at least five hours to about thirty hours using, on average, less than 10 continuous watts of power. Some embodiments of portable CRRT device 100 will use, on average, less than 3 continuous watts of power. Embodiments of the invention will weigh between two and ten pounds (between 0.9 and 4.5 kg) when in operation.

The battery 110 is removably installed in the portable CRRT device 100. The battery 110 can be rechargeable and can be recharged either while contained in the portable CRRT device 100 or while disconnected and removed from the portable CRRT device 100. The battery 110 can store

534 780 enough energy to operate a typical portable CRRT device 100 for at least five hours and in some embodiments up to thirty or more hours of continuous uninterrupted device operation. The microcontroller 108, by itself, or via additional circuitry, checks the charging status of the battery 110. If the microcontroller 108 determines that the battery 110 has low charge or has less than an estimated predetermined amount of operating time remaining (e.g. one hour remaining), the microcontroller 108 can initiate an alarm state via the alarm circuit 112. The alarm circuit 112 may provide any combination of an audio, visual or physical alarm. The physical alarm signal may include vibrations or small stabbing-style shocks to the patient. An alarm condition or warning can be displayed on a monitor 114 which can be viewed by the user. Display 114 may be a liquid crystal display, a display with LED, or some other low power display technology. An alarm condition can also turn off all or predetermined portions of examples of portable CRRT device 100. For example, a low-state battery may lower the speed of the pump, discussed herein below, to a very low volume of blood flow through the dialyzer to save energy and reduce the possibility of blood clotting. The pump can be lowered in speed to one tenth to one half of its normal pump speed to conserve energy until a battery charge or new battery is provided.

A moisture sensor 116 is electrically connected to the microcontroller 108. The moisture sensor 116 detects high humidity, condensation or liquid present within the package or coating (not specifically shown) of the portable CRRT device 100.1. For example, the moisture sensor 116 may be located near the pump (which is discussed later) or the filtration and absorption portion. In other embodiments, an additional moisture sensor may also be located near the dialyzer input and / or outputs.

The package or cover that partially or completely encloses the typical CRRT device 100 may be the combination of plastic, fabric, rubber material, polypropylene or other suitable material. The cover may cover part of the portable CRRT device 100 and allow access to various parts of the device, such as the monitor 114 and / or the user or doctor controls 118.

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High humidity, condensation or presence of fluid inside a typical portable CRRT device 100 may be indicative of a patient's blood, dialysate leakage or other fluid leakage within the CRRT device 100. Upon sensing moisture, the moisture sensor 116 provides a signal to the microcontroller 108 and an alarm is started. via the alarm circuit 112. Depending on the location of the moisture sensor in the portable CRRT device 100, different responses to the sensed moisture may occur. For example, if a moisture sensor detects moisture near the bloodstream (as will be described later), CRRT device 100 may shut off the pump completely, provide a sound alarm to the patient carrier, and utilize wireless communication in circuit system 120 to call an emergency number for assistance. Conversely, if moisture is sensed near the filtration and pad absorption portion 122 of the portable CRRT device 100, then in some embodiments, the alarm circuit 112 may sound an audible or patient-sensed alarm, lower the pumping rate of the blood and dialysate, and make additional checks via a gas or pressure sensor for possible air pollution or bubbles in the dialysate loop of CRRT device 100. It is noted that the wireless communication circuit system 120 may also be capable of providing the geographical location of the typical portable CRRT device 100 via a wireless communication 120 to a third party capable of providing assistance or service to the patient.

Pump 124 is an electric pulsatory pump. Embodiments of the present invention may have a one-channel pulsatory pump or a two-channel pulsatory pump. Fig. 1 shows a two-channel pulsatory pump 124. Each duct of the two-channel pulsatory pump contains a rubber tube portion 126 and 128. Each rubber tube portion includes valves at each end to direct a fluid flow through the rubber tube portion in a principal direction (depicted by an arrow on the tube portion 126 and 128). An engine and transmission (not specifically shown) in the pulsatory pump provides a sliding element which oscillates back and forth between the rubber tubes 126 and 128, thereby alternately pressing against the elastic rubber tubes and creating a pulsating flow into and out of each of the established created the chambers in the rubber tubes 126 and 128. There are different types of pulsatory pumps. Some such pumps use pneumatic pressure to cause bending of the valve chambers in the pump.

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In some embodiments, a pulsatory electromagnetic pump can be used to minimize moving parts of the pump. Typical two-channel pulsatory pumps would use the same motor and transmission to ultimately pump both channels into the pump in a pulsating way. This saves energy and minimizes the size of the pump system in the exemplary portable CRRT device 100. Microcontroller 108 can be used to control various pump variables. Any adjustable pump variables include, but are not limited to, adjustment of pump stroke, volume per stroke, speed or number of strokes per minute in each chamber, torque of the pump motor and transmission combination, pump speed (ie the number of pump cycles per minute), pump pressure, pump pressure difference and input pressure. output of the pump, and pump break and cycle times. In some embodiments, the slide element 130 may be adjusted to press against each of the flexible tube portions 126 and 128 of longer or shorter stroke so that the volume of fluid pumped through one of the channels is greater than the volume of fluid pumped into each pulse of the other. pump channel. The slide element 130 is part of a mechanical mechanism which can pump both channels (blood and dialysate) into a typical CRRT device.

An example of a portable CRRT device 100 has two fluid circuits: a blood circuit indicated by arrows 132 and a dialysate circuit indicated by arrows 134. Two-channel pulsatory pump 124 is used in the typical embodiment. The pulsatory pump may have flexible rubber cartridges 126 and 128 which provide flow in a single direction or in opposite directions. The shown pulsatory pump 124 provides flow across the pump or through the pump in opposite directions. Each cartridge or pump chamber 126 has an inlet valve at an inlet side 134 of the chamber and an outlet valve at the outlet side 136 of the chamber. The input and output valves are ball valves that minimize damage to the blood cells in the bloodstream. Flap valves can also be used in various embodiments.

In the bloodstream 132 of a typical CRRT device 100, the blood from a patient is introduced into the bloodstream at the blood entry 138 and the bloodstream through the bloodstream 132 where anticoagulant fluid (e.g., heparin and other appropriate equivalents) can be mixed with the blood via an anticoagulant reservoir 140 and an anticoagulant micropump.

534 780

142nd The anticoagulant micropump may be a piezo or diaphragm MRRO pump or other type of low energy pump to provide small amounts of anticoagulant fluid to the blood flow.

The blood flows past a first pressure sensor 144 before being introduced into the inlet valve 134a of the two-channel pulsatory pump. As the pressure member 130 presses against the flexible tubular rubber cartridge portion 126, the inlet valve 134a and the outlet valve 136a open and allow a pulse of blood to pass through the first pulsatory chamber [126], additional fluid reservoirs 146 and micropumps 148 can add additional substances to the blood flow, anticoagulant fluids or drugs. The blood then passes by a second pressure sensor 150 which senses the pressure of the output shock of the first channel of the pulsatory pump before the blood is introduced into the dialyzer 152. An example of a blood channel pulsatory pump chamber can provide a blood flow rate of between about 15 to 100 ml per minute. (pulsatile). A typical pulsatory pump 124 may have dimensions of 9.7 x 7.1 x 4.6 centimeters with a weight of less than 400 grams when not in operation. An example pump uses less than 10 watts of energy to pump two channels of fluid using a simple motor, transmission and pressure element which alternatively pumps each channel in a pulsatory manner at or about 180 degrees phase offset. If the pressure means of a typical pulsatory pump are difficult to compress or require very little energy to compress one of the flexible tubular rubber chamber portions, the pulsatory pump may provide an alarm to the microcontroller 108 indicating a possible occlusion in the blood or dialysate circuit's input or output valves. a pulsatory pump channel. A low power pulsatory pump using from 2 to about 7 watts of steady state can also be used successfully in typical embodiments of the invention.

The pulsatory pump can be fine-tuned by adjusting the motor transmission or reorientation of the pressure elements a small distance from centering between the tubular members 126, 128 so that the pulses or cycles of the two pulse chambers are in phase, 180 degrees offset, or alternatively so that each chamber pumps at a predetermined number of degree of phase shift to utilize

534 780 pump pulses to assist in maximizing the dialysis process that occurs in dialyzer 152. Embodiments of typical pulsatory pumps can be slowed so that the blood flow or fluid flow through a chamber of the pump is between 1 and 50 ml per minute during an alarm condition in portable CRRT device 100 .

Other types of pumps can also be used successfully or introduced into embodiments of the CRRT device 100. For example, two separate pulsatory pumps can be used provided that the pumps can be timed so that the pumps continuously pulse at predetermined alternating output rates. This can be done with stepper motors and microprocessor control. Other pumps which can provide a pulsed output, such as a double-sided piston pump having an intake stroke through a valve and a pump stroke through a second valve, may also be used. Such a pump may be magnetically or pneumatically driven.

With continued discussion of the blood circuit 132, the blood passes through the second pressure sensor 150 and then through the hollow lumen or fibers of the dialyzer 152 (not specifically shown). The dialysis procedure takes place in the dialyzer wherein impurities in a patient's blood move via a mass transfer through the outer membranes of the porous lumen through which the blood is transferred. The impurities are deposited in the dialysate as they exit through the outer surface of a lumen's outer membrane into the dialysate flowing through the dialyzer in dialysate circuit 134.

The purified blood exits the dialyzer and moves toward the bloodstream exit 154 to return to the patient. As the blood flows from the dialyzer 152 to the bloodstream's outlet 154, additional drugs, vitamins and other fluids can be added to the blood via reservoirs 156 and micropumps 158.

The blood pressure as it exits the dialyzer 152 is sensed by a third pressure sensor or converter 160.

The microcontroller 108 can display pump status or other pump / flow related information such as pressure, flow rate or phase difference of two channel pump on the screen 114. User Controls 118, such as buttons, switches,

534 780 slider controls, knobs, clutch details or pressure sensitive switches (not specifically shown) can be used to enable the patient, physician, nurse, technician or computer-based device to adjust various settings and controls on the exemplary CRRT device 100. Furthermore, communication device 120 can be utilized to transmit and receive control settings via a paging or other telecom or wireless communication channels or network for fine tuning or drug adjustment of the typical CRRT device. For example, adjustments for pump speed 124, flow rate, pump RPM, pulse rate or any of the micropump flow rates can all be monitored or controlled via user interface 118 and 114 or wireless communication circuit 120. A doctor may have the ability to monitor the CRRT device from a satellite location and have the ability to adjust CRRT settings at long range.

Dialyzer 152 is shown as a single dialyzer but may be a single or multiple dialyzers connected in series or in parallel. Dialyzer 152 may take the form of a cartridge which can be buttoned or quickly connected and removed into and out of the blood / dialysate circuits by a physician, nurse or technician. The dialyzer 152 used in a typical CRRT device 100 may comprise from 0.2 to about 1 square meter of dialysis surface area (e.g., semipermeable membrane) on the lumen fibers therein. During dialysis, blood circuit 132 generally has blood flow in a direction opposite to that of dialysate circuit flow through dialyzer 152. Furthermore, it has been determined by experimentation that the pulsing of the phase-shifted pulsatory pumps helps maximize the dialysis process. Particularly embodiments of the invention which contain a phase-shifted pulsatory blood flow and dialysate flow through a dialyzer provide improved purification with respect to removal of toxins from the blood compared to similar systems which do not contain a pulsed or pulsatory dialysate in the blood flow through dialyzer. Such results have been found with typical embodiments even when the dialysate and blood flow through the dialyzer are in the same direction.

The combination of the first, second and third pressure sensors / converters 144,150 and 160 provides differential measurements which can be analyzed with the microcontroller 108. For example, the pressure difference across the dialyzer 152 is

534 780 too high (from blood input to output), it may mean, inter alia, that dialyzer 152 has multiple clots or occlusions in the fiber lumen therein. A high pressure difference can also mean that the dialyzer is operated at an excessively high blood flow. As a result, an alarm situation can be initiated by the microcontroller 108 or the blood pump 124 can be automatically adjusted by the microcontroller to operate at a higher or lower torque, pump speed or pulse volume so that the flow rate is reduced in predetermined turns in an attempt to stabilize or reduce the differential pressure. If a pressure at a pressure transducer 144,150,160 falls below a predetermined low pressure, it may be an indication that a fluid leak has occurred in the bloodstream or that air has been pushed into the bloodstream 132. Microcontroller 108 may shut down or lower the predetermined portions of the portable. CRRT device 100 in response to blood circuit pressure measured below a predetermined level. Further, the microcontroller 108 can initiate an alarm state in connection with the alarm circuit 112 and the communication circuit 120.

Still referring to Fig. 1, typical dialysate circuit 134 will now be discussed. A fourth pressure transducer 162 measures the dialysate pressure at the input side of the dialysate pump cartridge 128. The fourth pressure transducer or sensor 162 provides a pressure read at the input of the dialysate pump cartridge to the microcontroller 108. . It is understood that the pulsatory pump portion of the dialysate may also be a separate pulsating pump device in various embodiments.

Purified fresh dialysate from the sorbent filters 122 flows through dialysate circuit 134 past the fourth pressure transducer 162 and is pumped through the tubular rubber portion of the dialysate pump 128. During normal pump operations, the operating flow rate of the dialysate through the dialysate pump is between about 40 to about 100 ml per minute (pulsatory).

Embodiments of portable CRRT device 100 are designed to operate using less than about 1 liter of dialysate. Other embodiments may

534 780 function properly and only require between 300 ml and 400 ml of dialysate in the closed dialysate liquid circuit 134.1 Embodiments designed for young adults or children, the amount of dialysate needed for operation may be between about 100 to 300 ml of dialysate. Minimizing the amount of dialysate reduces the overall weight of a typical device and reduces the operating cost by minimizing the amount of medical waste produced by the dialyzer. The combination of dialysates and filters 122 allows the embodiment to circulate the dialysate liquid for at least 24 hours before a filter requires replacement. In various embodiments, filter replacement may be required at intervals between 24 and 48 hours. Furthermore, since less than one liter of dialysate is all that is needed in closed dialysate circuit 134, sterile or ultra-pure dialysate can be used economically in typical embodiments of portable CRRT device 100. It is further understood that in embodiments providing a portable and / or partially portable CRRT device, the amount of dialysate in dialysate circuit 134 may be more than one liter. In fact, the amount of dialysate or ultra-pure dialysate in a portable or mobile or partially portable device can range from 300 ml to about 5 or 6 liters of dialysate.

In existing large, essentially stationary dialysis machines, it is common to use about 90 liters of dialysates per patient per run. Depending on the amount of water required to create the dialysate, filtered water, rather than ultra-pure water, is generally used in such large dialysis machines. Filtered water is much less expensive than ultra-pure or sterile water. Filtered water used in large modern dialysis machines is allowed to have some bacteria in it. Bacteria are larger than the size of the pores in the membranes or fiber lumen used in a typical dialyzer 152. Because the bacteria are larger than the pore size of the semipermeable membranes of the lumen of the dialyzer 152, the bacteria cannot pass through the membrane and enter the patient's blood into the bloodstream 132.

Conversely, medical research has provided some results that are cumbersome with the use of non-sterile dialysate (ie, dialysates containing filtered water, bacteria, toxins or microorganisms). Medical research has shown that microorganisms and bacteria in non-sterile dialysate create waste products, toxins or toxins in the dialysate. Waste products from the bacteria can pass through the porous membrane of the dialysate and enter the patient's blood while the bacteria in question

534 780 themselves cannot. Such toxins are sometimes referred to as endotoxins. It has been found that endotoxins passing from the dialysate through the dialyzer membrane to a patient's blood can adversely affect a patient's health. The endotoxins can result in making a patient sick.

Since typical embodiments of portable CRRT device 100 require less than one liter of dialysate, it is economically appropriate to use ultra-pure or sterile water as the main component in preparing the dialysate.

The dialysate exits through the pulsatory dialysate pump portion 128 via the output valve 136b in a pulsating pulsatory manner and passes a fifth pressure transducer 164, which measures the dialysate pressure on the dialyzer input side 222 of dialyzer 152. of the blood in the lumen of the membrane fibers in the dialyzer. Furthermore, fluids are relatively non-compressible, and the peaks of the dialysate flow pulses alternate over time with the peaks of the blood flow pulses through the blood circuit 132 of the dialyzer. In other words, the pulsed flow of the dialysate through the dialyzer alternates with the pulsed flow stops of the blood through the dialyzer. In some embodiments of the invention, the alternating peak flows of the dialysate fluid and blood through the dialyzer are approximately 180 degrees offset.

While the dialysate is in the dialyzer 152, waste products and toxins in the blood pass through the lumen membrane of the dialyzer and into the dialysate due to diffusion and further in embodiments of the present invention in addition to convection and osmos forces. The improved purification of blood toxins that occurs in embodiments of the present invention is partly caused by transmembrane pressure (TMP) caused by alternating pulsatory pump action and pressure differential between the blood in the lumen membranes and the dialysate outside the fiber lumen membrane.

The dialysate flow exits through the dialyzer 152 and flows through a sixth pressure sensor or transducer 166. The pressure transducer 166 sends a signal to the microcontroller 108 indicating the pressure of the dialysate exiting through

534 780 dialyzer 152. The sensed pressure may help indicate a clogged dialyzer, dialyzer leak or other emergency.

The dialysate circuit 134 now moves the used dialysate from the output of dialyzer 152 past the pressure transducer 166 and toward dialysate filter section 122. The dialysate used in this portion of dialysate circuit 134 contains toxins, contaminants and other unwanted substances that have been removed from the patient's blood. The filtration section 122 filters or reacts with the predetermined substances in the dialysate used to recover the dialysate for continued recycling and use in the dialysate circuit. Any compounds or substances considered toxins or impurities in the dialysate used^sis urea, creatinine and ammonia, along with other substances that are usually removed by a patient's kidney.

In a typical embodiment, a first filter 168 in the filtration section 122 contains urease. The urea is used to filter the dialysate used and further functions to break down the urea that was removed from the blood in the dialyzer 152. When the urea is broken down, urea, at least two unwanted by-products are created. In general, the two by-products of degraded urea are ammonium (ammonia) and carbon dioxide.

The dialysate with ammonia and carbon dioxide passes through the first filter 168. Urea is essentially removed from the dialysate, but ammonia and carbon dioxide also need to be removed from the dialysate. The dialysate, ammonia and carbon dioxide enter the second filter 170. The second filter 170 contains a compound zirconium or zirconium phosphate (ie ZrPx). The zirconium in the second filter 170 captures the ammonia. It is understood by those skilled in the art of dialysis chemistry that different chemicals and derivatives thereof can be utilized to produce the same or similar results.

The zirconium or second filter 170 will possibly be saturated with ammonia. When saturation or near saturation, the zirconium filter 170 will be less effective for removing ammonia from the dialysate. It is not advantageous to allow ammonia or ammonium to circulate through the dialysate circuit and back to the dialyser. Thus, in a typical portable or portable CRRT device 100, a sensor 176 is placed in the dialysate circuit after

534 780 zirconium filter 170 to detect the presence of ammonia in the dialysate. The sensor 176 may be a pH sensor, an ammonia-specific sensor or a conductivity sensor. Various probe and electromagnetic-style sensors may be used herein to sense the conductivity pH or the amount of ammonia in the dialysate after flowing through the zirconium filter. If an ammonia sensor is used, it will detect if a predetermined amount of ammonia is present in the dialysate. If a pH sensor is used, it would detect whether the pH of the dialysate has become a predetermined amount more alkaline than an acceptable amount, i.e., the dialysate's alkalinity is outside an acceptable range of alkalinity. As more ammonia is present in the dialysate, the dialysate becomes more alkaline. It is noted that depending on the actual chemicals and absorbents used in the filters, the dialysate may become more acidic and as such a sensor would be used to sense it. If a conductivity sensor is used, it would sense the conductivity changes of the dialysate. If the conductivity of the dialysate is outside a predetermined range, it could be determined that ammonia increases in the dialysate. An electromagnetic sensor can also be used to sense the conductivity across the dialysate in a manner that would detect a change in conductivity (i.e., ammonia) within the dialysate.

The sensor 176 is electrically connected to the microcontroller 108. If the signal provided by the sensor 176 to the microcontroller 108 indicates that ammonia in the dialysate exceeds a predetermined amount (i.e., the pH or conductivity is outside a predetermined range), the microcontroller can conclude that the zirconium filter 170 a portion or a predetermined percentage of the ammonia in the dialysate stream therethrough. As a result, an alarm condition can be started by the microcontroller 108. The alarm condition may instruct the user that one or more of the filter cartridges for the filter sections in the filter section 122 require replacement. The alarm condition can also decrease the pump speed or flow rate of one or both channels of the pulsatory pump 124. Decreasing the pumping rate of the two-channel pulsatory pump 124 may increase the amount of ammonia absorbed by the zirconium filter 170 due to a reduced flow rate of the dialysate therethrough. Furthermore, the alarm condition may decrease the dialysate flow rate and increase or decrease the blood flow rate through a typical portable or

534 780 portable CRRT device 100 by adjusting the position or movement of the pressure means 130.

The sensor 176 used to detect the presence of ammonia in the dialysate should be placed after the second filter 170 containing zirconium phosphate. In other embodiments, the sensor 176 may be positioned after the third filter 174 containing aqueous zirconia or after the fourth filter 178 containing carbon. In other embodiments of the invention, a simple filter may be used containing layers or a mixture of the contents of the four filters discussed above. Furthermore, a parallel filter 172 can be placed parallel to another filter, for example filter 168, to reduce the pressure drop across one or more filters in the filter section 122. Both filters which are parallel (i.e. 172 and 176) can each be layered filters having various filtration and adsorbent substances therein. Further, the layers can be mixed, for example, carbon particles can be mixed through the entire filter among the layers of urease and zirconium or zirconium phosphate. Furthermore, the zirconium and zirconium phosphate can be mixed together with one or more filters. All in all, the filter section 122 of a typical embodiment may comprise one or more filter cartridges or filters that react with or adsorb substances found in the dialysate used so that the dialysate can be renewed for continued circulation and use in the dialysate circuit 134.

The third typical filter 174 includes aqueous zirconia, which can further remove impurities and ammonia from the dialysate. A bubble device, exhaust gas, valve device or hydrophilic membrane (exhaust gas) 180 may be part of each of the filters or filtered parts or be a separate element therefrom. An exhaust gas 180 is used in a typical embodiment to remove air, carbon dioxide and other gas bubbles which may be formed or found in the dialysate circuit 134. It is important that a very limited amount of gas bubbles pass through the dialyzer 152. As such, an exhaust gas 180 should be placed before the dialysate pump portion of the pulsatory pump 124 and after one or more of the filtering filters in the filtration section 122 with respect to the dialysate flow direction of the dialysate circuit 134.

534 780

A fourth typical filter 178 contains carbon and is used to further purify the dialysate from contaminants via adsorption. The first, second, third and fourth filter sections 168, 170, 174, 178 may be formed as combination or separate filter sections. Each cartridge can be inserted and removed from a typical portable portable CRRT device 100 by the patient, physician, technician or nurse. Each filter cartridge may contain layers or combinations of chemicals or adsorbents. In fact, a typical embodiment may have a simple cartridge filter-containing layer of required substances to clean and refresh the dialysate after passing through the dialyzer 152. The cartridges can be installed in a series, a parallel or a combination of series and parallel propagation. Further, the filter cartridge or cartridges may contain an exhaust gas 180 thereon or such an exhaust gas may be a separate element downstream from one or more of the dialysate filters or dialysate filter section 122.

The filter section or individual filter cartridges require replacement at intervals of about 12 hours to 48 hours. By experimentation, if the total volume of all of the combinations of all sorbent material needed, in necessary amounts, combined, will be between 400 cm³ and 2,500 cm³ and will not require change or replacement for 24 to 48 hours while the embodiment pumps dialysate at a flow rate of about 20 to 70 ml / min.

An auxiliary reservoir 182 and micropump 184 may be connected to the dialysate circuit 134 after the filtration section 122, but before the dialysis input side 133b. Although not specifically shown in Fig. 1, several reservoirs 182 and micropumps 184 may be connected to dialysate circuit 134. Micropump 184 may be any of the micropumps discussed above with respect to micropump 142. Here, micropump 184 and reservoir 182 can add chemicals and additives that refresh the dialysate and extend its ability to act as a dialysate. An example of portable CRRT device 100 may have as little as 300 ml to as much as about 1 liter of dialysate in the dialysate circuit 134. An example of a portable CRRT device can have from 300 ml to about 5 liters of dialysate. It is important for the sorbent section or filtration portion 122 to have the ability to purify and refresh the dialysate continuously as the dialysate circulates around the dialysate circuit 134.

534 780

In various examples of portable or portable CRRT devices 100, ultrafiltrate or other fluids may be removed from the patient's blood in addition to the dialysis process. If a patient's kidneys are not functioning properly, it may be important to remove excess fluids from the patient's blood as well as blood contaminants such as urea and creatinine.

In the dialysate circuit between which the dialysate exits the dialyzer 152 and enters the filtration section 122 of various typical embodiments, the ultrafiltrate / dialysate, together with other impurities and liquids obtained via the dialyzer 152, may be dropped or removed from the dialysate circuit 134. The ultrafiltrate 188 may be deposited. in a bladder or reservoir 190. The bladder or reservoir 190 may be contained or suspended under a typical portable or portable CRRT device 100 and have the capacity to store from about 0.1 to maybe 2 liters of ultrafiltrate liquid. A fill rate sensor connected to the liquid bladder 190 is in electrical communication with the microcontroller 108 to enable an alarm condition or liquid fill rate read in the ultrafiltrate blower 190 when the bin fill rate reaches a certain level or volume. The liquid bladder 190 may also be inserted into portable or portable CRRT device 100 as an empty cartridge filled via a micropump and valve combination (not specifically shown). A fill rate sensor 200 can help the microcontroller determine the fill rate of cartridge bladder 190 and can turn off the ultrafiltrate supplying the pump and provide a signal to the user that the cartridge needs to be emptied. In various embodiments, the liquid bladder or cartridge 190 may contain an absorbent material (not specifically shown) to absorb liquid present in the bladder and prevent splashing. The absorbent material may be cotton, polymer, sponge, a compressed material, powder, a gel or other material which absorbs liquid and / or restricts the blistering of the bladder or cartridge. The bladder may be designed to expand as it is filled. The bladder can press against the microswitch (not specifically shown) when full or expanded, thereby providing a fill rate signal to the microprocessor 108.

In some embodiments, the ultrafiltrate cartridge or bladder 190 may contain a means for emptying the liquid bladder 202 thereon in the form of a cap, plug, valve, removable inner bias tube or otherwise.

534 780

Typical embodiments of a portable or portable CRRT device 100 can provide therapy to a patient containing basic dialysis, a complex dialysis method, ultrafiltration and various medical therapies to and for the patient. As discussed, a growing body of literature continues to indicate that increased dialysis time, which includes both longer and more frequent dialysis treatments, can be associated with improved outcomes and treatment of patients with terminal phase renal failure (ESRD), both in terms of probable longevity as well as expected morbidity. and mortality.

While experimenting with embodiments of the present invention, there were interests related to continuous and consistent transport of solute across the membranes of the hollow fibers of dialyzer 152 for the extended 24 hour time periods. Further, further interests were concerned with potential coagulation or clogging of the individual fiber membranes of dialyzer 152 which would limit the life and usefulness of dialyzer 152 for less than 24 to 48 hours. Various types of pumps including roller pumps, centrifugal pumps and pulsatory pumps are used as a pumping mechanism for the blood circuits and dialysate circuits of various embodiments. The various types of pumps all worked and produced a functioning portable or portable CRRT device, but the embodiments containing a two-channel pulsatory pump wherein the dialysate circuit 134 and the blood circuit 132 alternatively were pumped in a pulsatory manner at or about 180 degrees phase shift provided something unexpected; positive results. It was found that the purification levels of the blood toxins across the hollow fiber membrane dialyzer are higher when an alternating, two-channel phase-shift pulsatory pump is used with respect to in-phase pulse or pumps such as roller pumps and centrifugal pumps which are not supplied as much as are provided. membrane between the dialysate circuit and the blood circuit. Furthermore, unexpectedly, the use of the pulsatory pump decreased clogging of the hollow fiber membrane of the dialyzer due to protein build-up or up, coagulation or other fluid flow inhibitors in experimental devices over experimental configurations that did not contain the pulsating pump provided by both the pulsatory pump's blood and both.

534 780

The mechanisms of solute transport across the membrane of a hollow fiber of a dialyzer have been studied for more than half a century. Historically, it seems that precursors of dialytic therapy were well aware of the diffusion phenomenon when designing dialysers based on countercurrent mass exchangers. It was soon realized that convection also played a role as a mechanism in ultrafiltration and solvent drag (solvent transport) phenomenon. The factors affecting solute transport across semipermeable membranes have been summarized in various papers including Ronco et al., Evolution of Synthetic Membranes for Blood Purification, Nephrol Dial Transplant (2003) 18 the purification of small solutes such as urea, but larger solutes are mainly affected by ultrafiltration rates.

More recently, a push-pull hemodia filtration device (HDF) provided fast forward and reverse filtering through a dialyzer. This HDF mechanism led to alternating the flow of body fluid and dialysate across a high flow hollow fiber membrane. Although this technique worked, it had various drawbacks. One of the major drawbacks of conventional push-pull HDF devices is the necessity of a disposable blood reservoir bag required to prevent flow variation, as well as the difficulty of maintaining a transmembrane pressure (TMP) across the hollow fiber membranes that would not result in a collapse. the hollow fibers during back filtration. An aid to these push-pull HDF device problems has been to use volume controls for ultrafiltrate removal and rigid synthetic hollow fibers made of polyacrylonitrile, polysulfone and polyamide.

An example of portable or portable artificial kidney or CRRT device 100 shown in Figure 1 provides a lightweight, belt type or small portable device which is battery operated and utilizes unexpected benefits at a low pulsed flow rate of 40 to 80 ml per minute. of blood and dialysates through a dialyzer filter. Such a device can be used to provide continuous dialysis treatment to patients with terminal phase renal failure (ESRD) in a manner that may be continuous or substantially continuous for 18 to about 48 hours.

534 780

The base elements of a typical portable or portable CRRT device shown in Figure 1 can be summarized and described as comprising the following four main sections or sections. The first section is a dual-channel or pulsed dual-pump (or pumps) 124 which drives blood through the device in one channel and dialysates through the device in another channel. Both blood and dialysate can be pumped at different flow rates. Furthermore, the blood and dialysate are pumped in a phase-shifted or counter-phase pulsatory manner so that the maximum pumping pressure of the dialysate and blood channels occurs at alternating times. A second portion is a high flow hollow dialyser fiber membrane 152 comprising from 0.2 to about 1 square meter (m²) of membrane surface. In tests performed in a typical device, an AN-69 dialyzer of about 0.6 m was used² of fiber membrane surface. A third part or section of a typical embodiment comprises a dialysate regenerating system 122 comprising one to four specially designed powder or particle filled containers or cartridges containing the same or substantially similar sorbents used in the REDY system used in very large clinical size dialysis machines. The dialysate regeneration system may also contain reservoirs 182 of electrolyte additives including magnesium, calcium, potassium and sodium bicarbonate to the ultrafiltrate pathway, and a pH control circuit 176 for monitoring and extending the usefulness of the dialysate flowing through a typical device dialysate circuit. A fourth section or portion of a typical embodiment may include auxiliary pumps, generally small or micropumps 140, to deliver heparin 142 and perhaps other drugs 156 or substances to the bloodstream. An auxiliary pump or valve 186 may also be included to assist in the removal of the ultrafiltrate from the dialysate circuit. It is to be understood that each of these auxiliary or micropumps are all operated at microprocessor controlled 108 or pre-specified flow rates.

It is well known that conventional hemodialysis devices use quasi-stationary counterflows of blood and dialysates in a dialyzer. It was believed that such a stationary counterflow of liquids across the hollow fiber membrane of a dialyzer could contribute or be responsible for higher or increased performance in various dialyzers. Using embodiments of the present invention, a study was conducted to compare the effect of a pulsatory flow in the transport of solutes in a dialyzer with

534 780 high flow compared to stationary or quasi-stationary counterflows in a substantially identical high flow dialyzer. Comparative experimental studies and numerical studies were developed to clarify and verify the role of parameters affecting the transport phenomenon across a diaphragm's diaphragm when counter-phase pulsatory flow is used in both blood and dialysate circuits of a typical portable or portable CRRT device compared to using the same CRRT device with power demanding standard roller pump commercially used in dialysis equipment. It should be noted that the flow generated by a commercial roller pump is somewhat reminiscent of pulsation of either blood or dialysate provided by a pulsatory pump but with a substantially lower amplitude. The following equations have been adjusted to determine the purification and each week's standard urea Kt / V as measurements to evaluate the efficiency of dissolution removal of a high flow dialyzer in comparative studies.

Equation 1a:

purification = flow [21 solute] [/ ösf blank /]

Equation 1b: weekly standard

Kt (effective purification) x (time) V (total body water)

Equation 1a can be applied to both the blood side and the dialysate compartments of a typical dialyzer. Equation 1b reports on the effects of the distribution of solutes volume in blood, in the gaps and recirculation.

The comparison study was conducted to provide data explaining the improved purifications provided by a two-channel pulsatory pump in a typical portable or portable CRRT device. The experiments and comparative study provided support for the hypotheses that the improved purification of typical pulsatory pumping devices is due to pulsatory blood and dialysate flows in the counter phase generated by the pulsatory pumps together with pulse-generated washout effects established by the time dependent flow within the hollow dials. Such washout effects prevent the accumulation of large molecules near the wall. The large molecules can include proteins and other large substances that can be found in a patient's bloodstream but which do not readily pass through the fiber membranes into the dialysate.

534 780

Referring now to Fig. 2, a typical structure of a portable or portable CRRT device according to embodiments of the present invention is shown. This portion of CRRT device 100 mainly depicts the two-channel pulsatory pump 124 and dialyzer 152. This figure is provided to help explain the comparative results of experiments and studies due to the effect of an alternating blood pulsatory flow in the blood circuit 132 and dialysates in the dialysate circuit 134. Pressure sensors, reservoirs and micropumps and other elements depicted in Figure 1 are not included in Figure 2 to eliminate any figure confusion. Furthermore, Fig. 3 depicts a similar part of a portable or portable CRRT device as in Figs. 2, except that large, heavy, power-inefficient roller pumps or centrifugal pumps (standard pumps) are used to pump blood through blood circuit 302 and dialyzed through dialysate circuit 304. This dual standard pump structure 300 is, as Figure 2, designed without including pressure sensors, micropumps, reservoirs and various other details depicted in Fig. 1.

Referring to both Figures 2 and 3, dialyzer 152 is the same in both configurations. The particular dialyzer 152 is shown with only a hollow membrane lumen 212 extending the length L 210 of the dialyzer 152. It should be understood that in a real example or real dialyzer there would be hundreds of hollow lumen (fibers) extending the length L of the blood dialyzer. to pass through into the bloodstream 132. The membrane 214 on the outside of the membrane lumen 212 is the membrane through which the mass transport of solutes, including toxins, crosses or passes through the membrane 241 from the blood circuit 132 to the dialysate circuit 134. via the blood input 218 (Bi), travels through a plurality of lumen membranes 212 and exits the dialyzer via the blood output 220 (B_O). Similarly, dialysates in the dialysate circuit flow into the dialysate compartment 216 via the dialysate input or inlet 222 wherein the dialysate travels the length 210 of the dialyzer and exits the dialyzer compartment 216 at the dialysate output or outlet 224 and continues through the remaining portion of dialysate circuit 134.

534 780

With reference to Figure 2 specifically, the two-channel pulsatory pump 124 uses a 2-5 watt DC motor to move the pressure element 130 in an oscillating manner back and forth towards the flexible blood compartment 126 and dialysate compartment 128.1. 2 For example, the pressure means 130 are shown to push against the flexible chamber 128 of the dialysate and thereby push the dialysate in the direction of arrow 125 in the chamber through the output valve 136b while holding the input valve 133b of the dialysate chamber in a closed position. The pressure element is then moved toward the flexible blood chamber 126 and decompress the flexible dialysate chamber 128. Decompression of dialysis chamber 128 closes dialysate output valve 136b and opens dialysis input valve 133b. Meanwhile, as the pressure mechanism 130 presses against the flexible blood chamber 126 of the two-channel pump, the blood in the chamber is forced to exit the blood output valve 136a while the blood inlet valve 133a remains closed. In this way, a typical two-channel pulsatory pump provides counter-phase or alternating-phase or 180-degree phase-shifted pulsatory flows of the appropriate fluids through the blood and dialysate circuits. It is to be understood that a pulsatory pump can be produced in a variety of ways so that two chambers of liquids can be made to pump at alternating times. Such counterflow or alternating pump peak torque can be 180 degrees +/- about 90 degrees phase shift. Typical embodiments do not require the dialysate flow and the blood flow to have the same rate in ml per minute because the two liquids have different viscosities. Furthermore, the two-channel pulsatory pump can pump fluids in the same or opposite directions. Furthermore, the pumps may be two separate pumps as long as the alternating aspect of the pulsatory pumping is maintained. In other words, as a chamber of the typical pulsatory pump is filled with liquid, the second chamber of the pulsatory pump propels fluid out of its chamber and into its designated fluid circuit. Such an alternating pumping mechanism allows for a maximum pressure to be obtained in a circuit close to the same time that a minimum fluid pressure is obtained in the second liquid circuit.

The dialyzer 152 may comprise fibers having from 0.2 to about 1 m² in the membrane surface area around the hollow membrane lumen.

Referring now to Fig. 3, a standard roller pump 306 is used to pump blood through the blood circuit 302 while a similar standard dialysate pump 308 is used.

534 780 to propel the dialysate to dialysate circuit 304. Blood pump 306 and dialysate pump 308 are similar in structure and design to roller pumps used in prior art dialysis machines. During comparisons of the two configurations (Fig. 2 and Fig. 3), ultrasonic equipment was used to illustrate blood and dialysate flows through their respective circuits. Furthermore, pressure sensor and measuring equipment were used to measure blood input pressure at blood input B, by dialyzer 218 and blood output pressure at blood output B_O of the dialyzer 220 as well as dialyser input pressure at the input of dialyzer 222 and dialysate output pressure at the output of dialyzer 224.1 in the comparative experiments, dialysate and blood were pumped via the pulsatory double pump at flow rates of about 40, 50, 60 and 80 ml per minute (pulsatory). Furthermore, the pH of the dialysate was detected via a pH probe placed in a position after the filtration section, but before the dialysate was pumped into the dialyzer. Suitable additives such as sodium bicarbonate were added to the dialysate to maintain the pH of the dialysate between 7.3 and 7.5.

The dialyzer 152 contained approximately 4,500 lumen membranes each with a length of L 210 of about 15 cm. Fig. 4 shows the length L 210 by 15 cm and the inner radius of each lumen membrane R1 is about 120 microns. The outer radius of sample lumen R2 was about 170 microns. Thus, the thickness of the membrane 214 about lumen is equal to R2-R1 or about 50 microns thick. The diameter of the hemofilter was about 33 mm. FIG. 4 is obviously not drawn to scale but is provided to show a simple hollow fiber 212 extending a length 210 of the blood flowing dial within its inner radius R1 and dialysate flowing outside the outer radius R2 of the fiber 212 in the dialysate chamber or compartment 216. Solved substances will be transported from the blood into the center of the fiber lumen 212 through the membrane 214 and into the dialysate compartment 216.

For standard or roller pump configurations, pump speeds similar to the pulsatory pump speeds provided in the pulsatory dual pump embodiment were made so that the blood and dialysate flow rates in the configuration of Fig. 3 with the standard pumps 306 and 308 were also about 40, 50.60 and 80 ml per minute.

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Comparisons of the input and output pressures of the dialyser dialys inputs and outputs as well as blood inputs and outputs were compared in both experimental configurations. Furthermore, the purification of toxins and solutes was carefully studied based on samples of dialysate entering and exiting the dialyzer as well as samples of blood entering and exiting the dialyzer.

Unexpected purification results arose and were measured in comparative dialysis configurations using different pumps. In particular, the purification of the pulsatory pump configuration in Fig. 2 was substantially higher than the purification of the standard roller pump configurations. Purification is referred to as total removal or purification of toxins and dissolved substances from the blood as it passes through the dialyzer. Furthermore, ultrafiltration of the blood with respect to the energy required occurred more efficiently using the typical two-channel pulsatory pump configuration than using a standard roller or centrifugal pump configuration. Although measurements could only be taken at the inputs and outputs of the dialyzer, speculation can be made to try to explain what is happening in the dialyzer between the dialysate circuit and the bloodstream and across the fiber membranes. The results of the comparison are described in more detail below.

The results of test configurations using an example two-channel pulsatory pump in comparison to two standard roller pumps, one for blood 306 and one for dialysate circuit 308. off at dialysis input 222 and dialysate output 224. Flow rates were measured using high resolution Doppler ultrasonography and pressure was detected using microprocessor sensors. Results of the tests were divided into two categories; a first category was the ultrafiltration results and the second category was the purification results. It was determined that the ultrafiltration output of the dialyzer was found to be mathematically proportional to the pressure difference between the blood within the lumen of the dialyzer and the pressure of the dialysate in dialysate compartment 216. The measurements were taken during a steady state operation of a typical embodiment; thus there was a preloaded pressure in the inlet to a afterload pressure in the outlet from the pumps. The preload was induced by

534 The 780 pump head pressure and the afterload were created by the resistance of the dialysate filtration and absorption section 122 of a typical embodiment.

The ultrafiltration properties are depicted with further experimental results in Figures 5A, 5B and Figures 6A, 6B. Figures 5A and 5B show experimental results providing pressure and flow rates at the dialys inputs (inlets) and outputs (outlets) together with an ultrafiltration output for the roller pump configuration. In FIG. 6A and 6B, pressure and flow rates are shown for the inputs and outputs of the dialyzers together with the ultrafiltration output for the exemplary pulsatory pump configuration. In each situation, pressure and flow rates at the inlets and outlets of the dialyzers are shown together with the ultrafiltration rate. The pulsatory pump configuration of Figure 6A generates a time-dependent flow of blood and dialysate through the dialyzer. In particular, the input and output pulses 402 and 404 for peak or maximum blood pressure are generally about 180 degrees phase shifted with the peak or maximum pressure of the dialysate input and output 406 and 408. The duration or length of the high blood pressure portion 405 of the periodic blood pulse feed may be longer or shorter than the length of the low pressure portion of the periodic blood pulse feed. The duration of the high and low blood pressure portions may have a ratio of from about 4: 3 to about 3: 4. Ideally, the ratio would be 1: 1, but it is not specifically necessary for good purification results. The blood pressure at the input 402 and the blood pressure at the output 404 of the pulsatory configuration dialyzer is in the range of from about 0 to + 60 mmHg (from about 0 to + 7999 Pa). Meanwhile, the dialysate pressure at the input 406 and the dialysate pressure at the output 408 are at or near their pressure minimum of between about -90 to -60 mmHg (between about -11999 Pa to -7999 Pa) when the blood pressure is near its peak of between 55 and 60 mmHg. The duration duration of the minimum 409 of dialysate pressure is similar in length and overlaps fifty percent (50%) or more of the time duration of the maximum duration of blood pressure 405. Further, the pressure of the dialysate input and output 406, 408 is at their maximum of about 0 mmHg (0 Pa) while the pressure of the pulsating blood flow at the input 402 and the output of the dialyzer 404 is at their minimum pressure of about 0 to 10 mmHg (0 to 1333 Pa) shown in Fig. 6B. The duration of the pressure maximum of the dialysate input and output 411

534 780 overlaps at least about fifty percent (50%) of the pulsed blood pressure during its minimum duration 407.1 Fig. 6b, it can be seen that when the difference between the blood pressure at the input and the output 402,404 is greatest with respect to the pressure of the dialysate at the input and the output 406,408 the flow of dialysate 410 at or near its maximum. It is also important to note that the difference in pressure between the dialysate flow through the dialyzer and the blood flow through the dialyzer appears, at the input and output of the dialyzer, to be in the range of about 120 mmHg (15999 Pa) +/- (2666 Pa) for a maximum difference. and 10 +/- 10 mmHg (1333 Pa) for a minimal pressure difference.

In the study of Fig. 15, an ideal, even easier to understand, graph is shown of the output pressures of the blood channel of the pulsatory pump and the dialysate channel of the pulsatory pump. The periodic dialysate flow 500 has a maximum pressure duration period that overlaps the minimum pressure duration period of the periodic blood flow 502. This overlap 504 coincides with the overlap of the duration portion 411 of the dialysate with the maximum and the blood duration period 407 at the minimum of the actual experimental results in Figure 6A. Further back in Fig. 15, the duration portion with minimum pressure of dialysate flow 500 overlaps or coincides with the duration portions of maximum pressure of periodic blood flow 502. This overlap 506 coincides with the overlap of the duration portion of the dialysate 407 with minimal pressure and the period of blood at maximum pressure of the periodic blood flow 405 of the actual experimental results of Figure 6A. The main pressures caused by the filtration and adsorption portion 122 and the blood and dialysate inlet and outlet ports appear to have established the unexpectedly increased TMP across the fiber membranes resulting in an elevated purification of blood toxins flowing through the dialyzer. The enhanced purification is achieved using very little power into the pulsatory pump or pumps which can be introduced into the embodiments of the invention.

In Fig. 15, it should also be noted that the lengths or durations of the high pressure (and high flow rates) pulses for dialysate fluid flow 500 should be substantially the same as the lengths or durations of the low pressure pulses (or low flow rates) of the blood flow 502.

534 The high pressure pulse duration of the 780 dialysate to the low pressure pulse duration should have a range of ratio between 3: 4 and 4: 3. It follows that the lengths of the low pressure (and low flow rates) of the dialysate pulses should be substantially the same as the length or duration of the high pressure (or high flow rates) blood pulses. This is useful in embodiments of the invention to synchronize two separate pulsatory pumps. Such dual separate pulsatory pumps may be synchronized using digital stepper motors or by having digital sensors that detect rotational positions of a motor and adjust motor speeds with a motor control circuit.

In the study of Fig. 5A, the roller pump configuration, it is immediately noted that the pressure difference between the blood input and output and the dialysis input and output has a small pressure difference range. The difference between the average of the input and output blood pressure, at some point, with the average of the input and output dialysate pressure, is in the range of about 0 +/- 5 mmHg (0 +/- 667 Pa) and about 35 +/- 7 mmHg (4666 +/- 933 Pa) of pressure difference. FIG. Fig. 5B shows a substantially lower amount of ultrafiltrated 510 produced with the roller pump configuration over time than produced with the pulsatory pump configuration of the exemplary embodiment shown in Fig. 6B. Also note that although the roller pump provides some element of pulsation in the blood flow, the reduced pressure portion of the pulse pump pulse is much shorter than the reduced pressure portion of the blood pressure pulse in the pulsating pump configuration of Fig. 2. Thus, the maximum pulse / min pulse duration of the pulsatory pump configuration is substantially similar in time (ie +/- 10%). Further, the maximum amount of pressure provided by the roller pump configuration, at similar blood flow rates, is that in the pulsatory pump configuration a pulse longer than the pulsating pulse configuration. The blood pressure at the input and output of the dialyzer provided by the roller pump configuration is generally about 20 mmHg (2666 Pa) or even lower than the peak pressure of the blood pressure sensed at the input and output of the dialyzer in the pulsatory pump configuration. This was an unexpected result as the pulsatory pump uses only about 3 watts or less of energy to produce flow rates provided in which the roller pumps of the roller pump configuration require 20 to 50 watts

534 780 of energy to provide similar blood flow rates and pressures as shown in Fig. 5.

Referring back to Fig. 6B, it can be seen that the flow rate through the dialyzer with the pulsatory pump configuration varies in a somewhat or near sinusoidal pattern for both the dialysate and blood flow. The maximum and minimum flow rates of dialysate and blood flow (peak to peak) are in the range of from about 5 ml per minute to about 20 ml per minute. As such, the fluids transported through the dialyzer are accelerated, reducing the speed of a regular pulsatory pattern. Conversely, the flow rates shown in Fig. 5B in the roller pump configuration tend to maintain a steady state for about 4 seconds and then decrease speed and accelerate rather rapidly in a negative pulse for a period of about 1 second. As such, the acceleration and speed decrease in the roller pump configuration is very short-lived compared to the steady state pulse of the same.

Fig. 7 provides the ultrafiltration rate when compared to peak transmembrane pressure (TMP). TMP is the difference in pressure between the blood in the lumen of the dialyzer 212 and the dialysate in the dialysate compartment 216 of the dialyzer filter. As shown, the ultrafiltration rate is highly dependent on peak TMP; however, the ultrafiltration dependence on TMP appears to be non-linear. This discovery is also unexpected as the current application and understanding of ultrafiltration in a dialyzer is assumed to be linear relationship with respect to TMP. This unexpected result of a nonlinear ultrafiltration dependence on the TMP was shown with a best fit of a second order polynomial to the experimental data points of the roller pump configuration in Figure 3 and separately to the exemplary dual pulsatory pump configuration in Figure 2. The slopes obtained for the pulsatory pump configuration were larger than the slopes obtained for the roller pump configuration but only by a small margin. The small margin of improved ultrafiltration capability may be significant when an exemplary portable or portable ultrafiltration device is worn or used on a patient for extended periods of time. Significance is that additional ultrafiltrate can be removed from a patient's blood using a pulsatory pump configuration, which requires much less energy than a roller pump or

534 780 centrifugal pump configuration and provides substantially similar blood and dialysate flow rates through a dialyzer. Furthermore, the pressure-like movement created by the pulsatory pumps and the combined washing effect on the inner walls of the fibers may also play a role in the further effect of exemplary embodiments of the invention.

The experimental results of the purification levels provided similar unexpected results via experimentation. In the experimental set of the typical embodiment using a two-channel pulsatory pump, the two-channel pulsator pump operated at a pulsatory speed of about 2 hertz. It is to be understood that a larger tubular pump chamber with pulse chamber or smaller pump chamber would enable a pulsatory pump to operate at an oscillating frequency of between about half a hertz to about 4 hertz. Pulsatory pumping at a rate higher than 4 hertz may be detrimental to the blood cells in the blood flow, but would probably be acceptable to the dialysate circuit. It is believed that the best pulsatory flow rate for the blood and dialysate should be between 1 and 3 hertz so that the fluids flow at rates between about 20 and 100 ml / min.

In the dialysate compartment, the two-channel pulsatory pump establishes a relatively high pulsatile negative pressure leading to a high transmembrane pressure (TMP) with a maximum of about 140 mmHg (18665 Pa), but generally in the range of about 70 to 120 mmHg (between about 9333 and 15999 Pa). A TMP caused by the pulsatory or shuttle pump is more of a general approximation due to the incorrect phase shift or counter phase pulses of fluids (i.e. blood and dialysate) flowing through the dialyzer and in opposite directions. Furthermore, the pressure measurements were made only at the inputs and outputs of the dialyzer because the actual TMP across a membrane inside the dialyzer could only be computer modeled.

The ultrafiltration rate in the pulsatory pump embodiment is unexpectedly large relative to the blood flow rate (30-50%). It was determined that in a pulsatory pump configuration, the convection urea transfer is about 31% of the total urea transfer whereas in the roller pump configuration,

534 780 convection urea transfer only about 17%. It is believed that the larger convection contribution to the urea transfer in the pulsatory configuration is an essential part of the reason why the typical pulsatory configuration is an excellent configuration for a portable or portable CRRT or artificial kidney design. Fig. 10 depicts the difference in the ratios of convective urea flow to the total urea flow across a membrane. Here, it is easy to see the additional convective urea flow provided in the pulsatory pump configuration across the axial length of the dialyzer.

With reference to Figure 8, a quick comparison of a roller pump configuration and a typical pulsatory pump configuration is shown. The blood flow (Qb.in) into both the dial pump of the roller pump and the pulsatory configuration as well as the dialysate flow (Qdjn) are both similar flow rates. The test results show that the ultrafiltration from the pulsatory pump provided more than twice the amount of ultrafiltrate (UF) for similar flow rates of the blood and dialysate into the dialyzer. Furthermore, the pressure difference between the pressures of blood coming out of the dialyzer with respect to the pressure of the dialysate coming out of the dialyzer produced a much higher maximum TMP in the dialyzer using the typical pulsatory pump configuration.

Figures 9A and 9B show graphs with a calculated pressure distribution along the axial distance L210 of the dialyzer in both the roller pump configuration and a typical pulsatory or shuttle pump configuration. The maximum pressure distribution is displayed. The pulsatory pump configuration shows that on the dialysate side, the pressure increases from about -32 mmHg (-4266 Pa) at the dialysate output side of the dialyzer and it increases to about -28 mmHg (-3733 Pa) at the dialysate input side of the pulsatory pump at about 15 cm. The blood pressure at the blood entry side of the dialyzer decreases slightly from about 23 mmHg (3066 Pa) to about 20 mmHg (2666 Pa) at the outlet side of the dialyzer. Thus, an unexpectedly high maximum TMP is produced in the range from about 48 to about 55 mmHg (from about 6399 to about 7333 Pa) across the axial distance between the dialysate and the blood flows of the dialyser. Fig. 9A indicates that the maximum TMP of a dialyzer using the roller pump configuration is much lower;

534 780 about 50% or more than the pulsating maximum TMP established in the typical pulsatory pump configuration.

Figs. 11, 12 and 13 provide comparisons of the urine, creatine purification and potassium purification achieved in the roller pump configuration and a typical pulsatory pump configuration of a portable or portable CRRT device or artificial kidney device. It is noted from FIG. 11,12 and 13 that for a similar dialyser input flow rate at the blood input 218 and dialysing input 222, TMP and the resulting purification of urea, creatine and potassium are always greater in the typical pulsatory pump or pulsed flow configuration of the typical embodiment.

Based on detailed studies and numerical studies, it is determined that the flow of solutes across a dialyzer membrane varies depending on the pumping mechanism used to move the dialysate and blood through the dialyzer. In particular, it has been determined that the use of a two-channel pulsatory pump configuration wherein a pumping channel is for the dialysate and a second pumping channel is for the blood and where the channels are pumped alternately, counter-phase, or 180 +/- 90 degrees phase offset provides excellent ultrafiltration function and purification of dissolved substances from blood for previously used roller pump and centrifugal pump configurations. Using experimental and numerical methods, the parameters affecting the transport phenomenon across the dialyzer membrane, when counter-phase of pulsatory flows are used in both the blood and dialysate compartments of the dialyzer, exemplary embodiments unexpectedly excellent ultrafiltration and purification of toxic solutes, heavy and power inefficient roller and centrifugal pump configurations. The results of the typical two-channel pump configuration have been compared to a configuration using standard roller pumps or centrifugal pumps used in conventional dialysis equipment. Although the flow of the dialysate and the blood generated by a roller pump flow is similar to a pulsatory flow, a lower amplitude flow and pressure of a roller pump is provided and its flow pulse is much longer than its non-flow or zero pressure pulse on the roller pump. Conversely, the pulsatory pump provides a substantially equal flow / high pressure pulse and non-flow / low pressure pulse.

534 780

The use of a typical two-channel pulsatory pump utilizes important new revealing discoveries regarding the flow and pressure behavior of some fluids, such as blood and dialysates as they pass through a dialyzer. Furthermore, improved purification is established for urea creatine and potassium thereby allowing the typical pumping system to operate at a slower rate for longer periods of time and using only 3 to 5 watts of energy to power the pumping. Further, using a typical pulsatory pump configuration with the filtration and absorption section of a typical portable or portable CRRT device, substantially twice the TMP (100-140 mmHg (13332-18665 Pa)) was established across the membrane when compared to the TMP established with a roller pump. or centrifugal pump configuration. The reason for the improved TMP in the typical embodiments is due to the pulsatory counter-phase flow of the blood and the dialysate through the dialyzer which establishes the difference in pressure between the blood in the fiber lumen and the dialysate chamber of the dialyzer.

Further advantages of the typical two-channel pulsatory configuration for a portable or portable CRRT device are that the ratio of the amount of ultrafiltration to the TMP is not linear, as previously understood roller and centrifugal pump configurations. Instead, the ratio of ultrafiltration to TMP is based on a non-linear second order differential equation.

It is important to note that one of the important aspects of the two-channel pulsatory pump configuration is its high efficiency in removing urea, creatine and other dissolved substances from the blood. Figures 14A and 14B show how the relevant concentration of the dissolved substances is transmitted across the membrane over the axial length of a blood dialyser and the dialysate flowing in opposite directions therethrough. In FIG. 14A, the roller pump configuration is shown wherein the relative concentration falls for the dialysate indicating that less solute is transmitted across the membrane at the blood output side of the dialyzer with respect to the blood input side of the dialyzer. On the contrary, in Fig. 14B, the relative concentration distribution or transfer of solutes from the blood to the dialysate is more linear and does not fall off. It is believed that the reason for this more evenly distributed distribution of solute transfer through the membrane and along the

534 The length of a lumen fiber in a dialyzer depends on the increased ultrafiltration across the membrane. The higher ultrafiltration rate caused by the higher TMP across the membrane (due to the alternating pulsatory pumping of the dialysate and blood) results in improved convective forces, thereby resulting in a better mass transfer of solute across the membrane along its entire axial length across the membrane. . With the typical two-channel pulsatory pump configuration, the convection was attributed to 31% purification of urea, while in the roller pump configuration, the absence of the additional convection urea transfer was limited to only 17%. The experimentation showed that the typical pulsatory pump configuration provided a 31% increase in convective mass transfer over the roller pump configuration. The typical high TMP created by the two-channel pulsatory pump configuration of a typical embodiment provides improved purification and ultrafiltration with a low approximately 2-5 watts of power required for the pump motor. This is an unexpected result when compared to the amount of energy required for a roller pump or centrifugal pump to pump dialysate or blood at similar flow rates through a dialyzer in a similar configuration. This is essential in that a small dual-channel, battery-powered, and rechargeable portable or portable dialysis device can be efficiently produced and provide better performance than a large, heavy-duty 120-volt AC or roller or centrifugal pump driven at similar flow rates.

A numerical simulation performed in addition to actual experimentation showed that the increased transmembrane transport of a typical two-channel pulsatory pump configuration is mainly controlled by diffusion across the fiber membranes, but the transmembrane transport is substantially increased over other pump configurations due to the additional convection created by TM. the membrane, which was prepared by counter-phase pulsatory pumping of the dialysate and blood circuits. Furthermore, it is believed that the improved purification provided by the typical embodiments can be attributed to a washout or push-pull-like effect associated with the pulsatory flows through and along the boundaries of the fiber membranes. Referring back to Fig. 6B, it is shown that the flows of the dialysate and the blood in the dialyzer are each equally spaced pulses with flow rate maximum and minimum located substantially at equal distances.

534 780 from each other over predetermined time periods. A predetermined time period for the typical embodiments was about 2 pulses per second, but it should be understood that the various embodiments of the invention may have a pulsatory pulse flow rate of between about 0.5 and 4 pulses per second, with flow rates ranging from about 20 to 100. ml / min during stationary operation.

It will be appreciated by those skilled in the art who have the advantage of this specification that this enhanced purification in an artificial kidney containing a pulsatory pump provides an energy efficient device for performing dialysis. Such a typical device may be carried in its entirety on a patient as a fully portable dialysis device or it may be inserted into a portable device which may be carried or pushed (on a trolley) by the patient or medical personnel. It should be understood that the figures and detailed description herein are intended to be illustrative rather than a limitation, and are not intended to be limiting to the particular shapes and examples described. On the contrary, some additional modifications, alterations, rearrangements, substitutions, alternatives, embodiments and embodiments apparent to those of ordinary skill in the art are included, without departing from the scope thereof, as defined by the following claims. Thus, it is intended that the following requirements be interpreted to include all such additional modifications, modifications, rearrangements, substitutions, alternatives, embodiments and embodiments.

534 780