for the proposed clinical use , the capability of the invention is to remove water drove the design . this is fortunate since the convection technique proposed here for removal of water is inefficient ( in terms of the solute material removed per unit of the surface area of the membrane ) in removing solute from the blood plasma . the functional kidneys of the patient are relied on to remove solutes from the plasma . the kidneys are , thus , relived of having to perform substantially all fluid removal from the blood . fig1 illustrates the treatment of a fluid overloaded patient with the present invention 100 . patient 101 can undergo treatment while in bed or sitting in a chair . patient can be conscious or asleep . to initiate treatment two relatively standard 18 g needles 102 and 103 are introduced into suitable peripheral veins ( on the same or different arms ) for the withdrawal and return of the blood . this procedure is no different from blood draw or iv therapy . needles and attached to tubing 104 and 105 and secured to skin with attachments 106 and 107 . the blood circuit that consists of the blood filter 108 , tubes , pressure sensors 109 , 110 and 111 and the ultrafiltrate collection bag 112 . the circuit is supplied in one sterile package and is never reused . it is easy to mount on the pump 113 and can be primed and prepared ready for operation within minutes by one person . during operation , the present invention requires minimal intervention from user . user sets the maximum rate at which fluid is to be removed from the patient using the control panel 114 . ultrafiltrate is collected into a graduated one - liter collection bag 112 . when the bag is full , ultrafiltration stops until the bag is emptied . information to assist the user in priming , setup and operation is displayed on the lcd display 115 . fig2 illustrates the operation and fluid path of the preferred embodiment of the present invention . the present invention consists of a microprocessor controlled console and a disposable kit . disposable kit is bonded ( with the exception of needles ) and is supplied sterile . blood is withdrawn from the patient through the 18 gage or similar withdrawal needle 201 . the needle 201 is inserted into a suitable peripheral vein in the patient &# 39 ; s arm . blood flow is controlled by the roller pump 204 . before entering the pump blood passes through approximately two meters of plastic tubing 202 . tubing is made out of medical pvc of the kind used for iv lines and has internal diameter ( id ) of 3 mm . pump is rotated by a dc motor under microprocessor control . the pump segment ( compressed by the rollers ) of the tubing has the same id as the rest of the blood circuit . the system is designed so that approximately 1 ml of blood is pumped per each full rotation of the pump , e . g . pump speed of 60 rpm corresponds to 60 ml / min . the disposable withdrawal pressure sensor 203 is a part of the blood circuit . pressure sensor 203 is a flow - through type commonly used for blood pressure measurement . there are no bubble trap or separation diaphragms in the sensor design , which reduce the accuracy . pressure sensor is designed to measure negative ( suction ) pressure down to − 400 mmhg . all pressure measurements in the fluid extraction system are referenced to atmospheric . the withdrawal pressure signal is used by the microprocessor control system to maintain the blood flow from the vein . typically , a peripheral vein can continuously supply 60 - 100 ml / min of blood . this assumption is supported by the clinical experience with plasma apheresis machines . in some cases , blood flow can be temporarily impeded by the collapse of the vein caused by the patient motion . in other cases the vein of the patient may not be sufficient to supply the maximum desired flow of 60 ml / min . the software in the present invention microprocessor is designed to control the withdrawal of blood to prevent or recover from the collapse of the vein and reestablish the blood flow based on the signal from the withdrawal pressure sensor . the same pressure signal from the sensor 203 is used to detect the disconnect of the withdrawal bloodline 202 from the needle 201 . this condition is detected by the abrupt decrease of the withdrawal pressure generated by the pump . the resistance of the 18 gage needle , which is 4 cm long with an approximately 0 . 8 mm id at a flow rate corresponding to a 60 mmhg , pressure drop is on the order of the 100 mmhg . the resistance of 2 meters of blood tubing with a 3 . 5 mm id at the same flow rate is on the order of 20 mmhg . this enables automatic reliable detection of the line disconnection . the occlusion of the withdrawal bloodline is detected in the similar fashion . the occlusion can be caused by the collapse of the vein or by the kinked blood tube . occlusion results in a rapid decrease ( more negative ) of the pump suction pressure that is detected by the microprocessor . in response to this condition , the microprocessor stops the pump and alarms the user . the present invention uses a double roller pump 204 to pump blood . as the pump 204 rotates , rollers compress the segment of pvc tubing and generate flow . pump 204 is adjusted to be fully occlusive until the pressure limit is reached . the rollers are spring loaded to limit the maximum positive and negative pressure generated by pump head . this feature is not normally used to limit pressure in the circuit and is only included as a secondary safety precaution . a direct drive stepper motor rotates the rollers , and the speed of the motor is determined by the controller microprocessor . the rpm of the pump 204 is used as a feedback signal by the controller to determine the blood flow . normal operational blood flow in the present invention is between 40 and 60 ml / min . this minimum rate of blood flow is needed to generate trans membrane pressure ( tmp ) needed for ultrafiltration and to prevent stagnation and clotting of blood in filter 207 . pump pressure sensor 205 is incorporated into the post - pump segment of the blood tubing connecting pump 204 to the blood inlet port 214 of the filter 207 . like other pressure sensors in the present invention it is a flow through device that does not create a blood - air interface and does not disturb the blood flow . the pump pressure signal is used by the microprocessor to determine tmp used to calculate the ultrafiltration rate . it is also used to detect abnormal conditions in the circuit such as occlusion or unacceptable clotting of the filter and disconnection of the blood line between the pump 204 and the filter 207 . on its way from the pump 204 to the filter 207 , blood passes through the air detector 206 . the air detector 206 is of ultrasonic type and can detect air in amounts exceeding approximately 50 microliters . the detector 206 uses technology based on the difference of the speed of sound in liquid and in gaseous media . if an air bubble is detected , the pump 204 is stopped almost instantaneously ( within few milliseconds ). the bubble detector output signal is hard wired into the motor control logic and the pump 204 is stopped independently of the microprocessor control if a bubble is detected . air can only enter the present invention circuit from the pre - pump ( negative pressure ) segment of the blood circuit 202 . all the rest of the circuit downstream of the pump 204 is always pressurized . for this reason , the bubble detector is placed before the filter . blood pressure in the post pump , pre - filter segment of the circuit is determined by the patient &# 39 ; s venous pressure , the resistance to flow generated by the return needle 210 , resistance of hollow fibers in the filter assembly 207 and the resistance of interconnecting tubing 208 . at blood flows of 40 to 60 ml / min the pump pressure is in the 300 to 500 mmhg range depending on the blood flow , condition of the filter , blood viscosity and the conditions in the patient &# 39 ; s vein . the filter 207 is a main component of the present invention . inside the filter 207 the ultrafiltration occurs . whole blood enters the bundle of hollow fibers from the connector on the top of the cap of the filter canister . there are approximately 700 hollow fibers in the bundle , and each fiber is a filter . blood flows through a channel approximately 0 . 2 mm in diameter in each fiber . the walls of the channel are made of a porous material . the pores are permeable to water and small solutes but impermeable to red blood cells , proteins and other blood components that are larger than 50 , 000 - 60 , 000 daltons . blood flow in fibers is tangential to the surface of the filter membrane . the shear rate resulting from the blood velocity is high enough such that the pores in the membrane are protected from fouling by particles , allowing the filtrate to permeate the fiber wall . filtrate ( ultrafiltrate ) leaves the fiber bundle and is collected in space between the inner wall of the canister and outer walls of the fibers . the geometry of the present invention filter is optimized to prevent clotting and fouling of the membrane . the active area of the filter membrane is approximately 0 . 1 m2 . the permeability kuf of the membrane is 30 to 33 ml / hour / m2 / mmhg . these parameters allow the desired ultrafiltration rate of approximately 500 ml / min at the tmp of 150 to 250 mmhg that is generated by the resistance to flow . the effective filter length is 22 . 5 cm and the diameter of the filter fiber bundle is 1 . 2 cm . this results in the shear rate of 1 , 200 to 1 , 800 sec - 1 at the blood flow rate of 40 to 60 ml / min . the tmp in the present invention is defined predominantly by the resistance of the return needle 210 and the resistance of the filter bundle inside the filter 207 . the properties of the filter 207 and the needle 210 are selected to assure the desired tmp of 150 to 250 mmhg at blood flow of 40 - 60 ml / min where blood has hematocrit of 35 to 50 % at 37 ° c . the quantitative clinical goal was formulated for the apparatus being developed in terms of fluid removal . applicants &# 39 ; research established that for the fluid removal device to be clinically useful it should remove water at the rate of 100 to 500 ml / hour . lower rates of fluid removal are only required in hemodynamically unstable patients that are treated in the icu and are not the targeted patient population . fluid removal rates higher than 500 ml / hour ( theoretically as high as 1 , 000 ml / hour ) may be practical in some patients but are expected to be too high risk in the majority . it is only advisable to remove water from blood at the rate at which fluid can be recruited from tissue . higher rates may lead to hypotension . blood hematocrit ( volume fraction of red blood cells ) in chf patients is expected to be in the range of 30 to 40 % of the total blood volume . it is possible to condense the filtered blood to the hematocrit range of 50 % to 60 % and still be able to return blood through a standard needle . therefore , extraction of approximately 20 % to 30 % of volume from blood as water is possible . assuming this extraction rate , the amount of blood removed from a peripheral vein is less than 2 % of the total cardiac output . in addition , at this extraction rate , the potential ultrafiltrate flow may be as much as 1 l / hour . alternatively , a lower extraction rate , e . g ., 0 . 1 liter / hour , may be selected . at the blood flow rate of 60 ml / min applicants successfully extracted up to 12 ml / min ( or 720 ml / hour ) of ultrafiltrate in the laboratory using the filter described here . therefore , it is possible to consistently extract the required 500 ml / hour of water from the blood flow withdrawn and returned into a peripheral vein . applicants established that the much higher blood flows that are used in adults by all existing renal replacement therapy machines and particularly by machines for acute cvvh treatment of chf patients are necessitated by the filter designed to remove solute and more specifically by the relatively high surface area of the filter . this large surface area is needed for solute removal . if the goal of treatment was to remove water only , high blood flow will still be needed to reduce the time of exposure of blood to the synthetic membrane and to prevent clotting . another important consideration that forces the designers of cvvh machines to use high blood flow and consequentially the central venous access is the need to maintain substantially high wall shear rate of blood flowing inside the filter capillaries ( hollow fibers ). flow of blood inside a fiber is laminar . shear rate at the wall can be calculated using the simple equation 1 : q is blood volumetric mass flow rate and “ d ” is the internal diameter of the capillary . the ultrafiltration rate is influenced by membrane fouling which is an equilibrium of wall shear rate and ultrafiltration rate per unit surface area . with the increasing surface area the wall shear rate will decrease unless the blood flow is increased to compensate . it becomes apparent from literature that the wall shear rate should be 1 , 000 sec - 1 or higher to achieve sufficient filtrate flux at high hematocrit . it is also known from literature that the high shear rate in excess of 2 , 500 sec - 1 is undesirable since it can cause hemolysis and damage to red blood cells . at the same time , it is apparent that the surface area and size of the filter should be minimized . biocompatibility is inversely proportional to the surface area exposed to blood . the likelihood of clotting increases with residence time proportional to the filling volume . also , cost of a smaller filter is lower . to minimize the cost of the filter , the use of commercially - available fibers with optimized biocompatibility and consistent filtration properties is desired . suitable filter fiber is available , for example , from minntech inc . in minnesota . each fiber has internal diameter of 0 . 2 mm . pores in the fiber walls are optimized to retain solutes of greater than 50 , 000 daltons . the permeability of this fiber is 33 ml / hour / m2 / mmhg . if a membrane with total surface area of only 0 . 1 m2 is constructed from this fiber , the resulting theoretical ultrafiltration rate will be 330 ml / hour at tmp of 100 mmhg and 660 ml / hour at tmp of 200 mmhg . these numbers are consistent with the objective of the design . to calculate the kuf of the filter , the permeability of fiber is multiplied by the surface area of the membrane . therefore kuf filter = 33 × 0 . 1 = 3 . 3 ml / hour / mmhg . it is known from literature that the blood flow is not equal between the fibers in the filter bundle . blood flow and consequentially the wall shear rate tends to be lower in the fibers closer to the periphery and higher in the central ones . accordingly , blood residence time is longer in peripheral fibers . it is known from the practice of dialysis that the peripheral fibers tend to clot first . to reduce the extracorporeal blood volume and the time that blood resides outside of the body it is desired to use blood lines that have internal diameter as small as practical without creating excessive resistance to flow . for our application , an internal diameter of around 3 . 0 mm is well suited . when blood exits the tubing and flows into the fiber bundle the diameter of the channel through which blood flows increases substantially . this creates turbulence and stagnant zones at the entrance into the bundle . these factors increase the probability of clotting . it is therefore beneficial to design a filter that has a minimal but still practical diameter of the fiber bundle . this is achieved by reducing the number of fibers and increasing the length of the bundle . this approach is limited by two constraints . resistance of the bundle to flow increases in proportion to the bundle length . also , long filters substantially in access of 20 - 25 cm are difficult to manufacture and cumbersome to use . applicants chose to use a maximum length of the filter that is practical from the manufacturing standpoint . the resulting working length of the bundle is 22 cm ( centimeters ). to ensure the required surface area of the membrane , approximately 620 to 720 fibers of this length are need to be bundled in parallel . assuming the fiber density of approximately 630 capillaries per cm2 the diameter of the bundle is 1 . 2 cm . such filter can be easily manufactured using existing methods and equipment . at the blood flow of 40 to 60 ml / min and blood hematocrit of approximately 40 %, the resistance of this filter to blood flow is on the order of 100 to 200 mmhg . this pressure level is acceptable for the design of a circuit with a standard peristaltic pump and an 18 to 20 gage ( internal diameter of 0 . 8 - 1 . 0 mm ) return needle . applicants overcame the perceived impossibility of clinical peripheral vein ultrafiltration that limited the use of mechanical fluid removal in chf patients outside of the icu environment . applicants did this by drastically reducing the filter membrane surface area compared to common dialysis or cvvh filters to maintain high shear rate and low blood residence time . specifically , a filter with the membrane surface of less than 0 . 2 m2 and preferably 0 . 05 to 0 . 15 m2 can remove the desired 100 to 700 ml / hour of water from the extracorporeal blood flow of less than 100 ml / min or more specifically of 40 to 60 ml / min with an average blood cell residence time outside the body of less than 2 minutes , and may be less than 1 minute . although the filter is made of high permeability fiber due to the small surface area the kuf of the filter is less than 5 ml / hour / mmhg or preferably 2 to 4 ml / hour / mmhg . typical filters used in adult patients have kuf of 30 to 50 ml / hour / mmhg . the much lower kuf gives the present invention device design an advantage of inherently safer operation . food and drug administration ( fda ) classifies all filters with kuf greater than 8 ml / hour / mmhg as “ high permeability dialyzers ”. according to current fda safety standards these devices have to be labeled for use only with ultrafiltration controllers that are independent of tmp based ultrafiltrate rate calculation . a small error of tmp measurement or a deviation of membrane permeability from the specification can result in substantial over or under filtration . the use of a low kuf filter enables , if desired , the present invention to avoid using a cumbersome and expensive ultrafiltration controller that typically involves a scale balance and an ultrafiltration pump . a filter that is relatively long and narrow may optimize the blood flow inside the filter , maintain the desired wall shear rate and minimize membrane fouling and filter clotting . a filter with a fiber bundle that is approximately 20 cm long and 1 . 5 cm in diameter is particularly well suited for the application and is practical for manufacturing . filters for ultrafiltration of blood with small surface area of less than 0 . 2 m2 are known . example of such filter is miniflow ™ 10 from hospal . miniflow has surface area of 0 . 042 m2 and kuf of 0 . 87 ml / hour / mmhg . all such filters without exception are used for hemofiltration therapy in neonatal patients and infants . the clinically used amount of blood flow through these filters is within the range that we targeted or 10 to 60 ml / min . nevertheless , this amount of flow — if expressed as a fraction of the cardiac output for infants — is the same as the blood flow used in adult hemofiltration . consequently , these infant filters are used with the central and not peripheral venous access . to minimize clotting and fouling of the membrane it is desired to maintain substantially high blood flow through the filter even if the desired ultrafiltration rate is low . in traditional machines for renal replacement therapy it is typically achieved by reducing the tmp . flow of ultrafiltrate is actively controlled by the roller pump in the ultrafiltrate removal line between the filter and the ultrafiltrate collection bag . when the pump is slowed down ultrafiltrate flow is retarded , pressure gradient across the membrane is reduced and ultrafiltration is slowed to the desired level . alternatively , if the pump rpm is increased , the flow of ultrafiltrate is accelerated . negative pressure can be developed by the pump to actively suck the ultrafiltrate across the membrane . for reasons of safety and simplicity , it was desired to have a machine that can reduce the ultrafiltration rate at the user command without an ultrafiltrate pump . for our preferred embodiment we used the duty cycle controlled ultrafiltration . a simple pinch valve is placed in the ultrafiltrate line . when the valve is closed , pressure across the membrane quickly equilibrates , and no ultrafiltration occurs . when the valve is opened , ultrafiltration occurs at the rate determined by the tmp and the kuf of the membrane . this rate can be calculated by the controller . valve is cycled approximately every minute . the fraction of the cycle during which the valve remains opened determines the average rate at which fluid is removed . since the system embodying the present invention does not employ an ultrafiltrate pump that can create sub - atmospheric pressure on the ultrafiltrate side of the membrane , a simple and reliable method of controlling the total amount of fluid removed in one treatment iteration is possible . the ultrafiltrate is collected into a sealed bag that is connected by a tube to the ultrafiltrate collection chamber of the filter casing . during the treatment the bag is gradually filled up with fluid . it is desired to have a bag that has a relatively small volume and specifically volume of 0 . 5 to 1 . 5 liters . when the bag is full and its walls are fully distended , the pressure in the bag will start to rise until it is equal to the average pressure of blood inside the filter capillaries . although some circulation of fluid is still possible in and out of fibers the net loss of fluid is zero . until a nurse empties the bag , no removal of fluid is possible . pressure sensors are used in the blood circuit to alarm the disconnection and occlusion of blood lines . the pre - and post - filter pressure signals are also used to calculate tmp and ultrafiltration rate . two types of pressure measurement devices are typically used in machines for renal replacement therapy . machines such as bp11 from baxter use disposable air filled separation or drip chambers that are connected to permanently installed pressure sensors that are the part of the machine . this design introduces potentially hazardous air into the circuit . air can cause embolism and accelerated clotting . also , this type of measurement is affected by gravity . machines such as prisma from gambro use flexible silicone diaphragms to transmit blood pressure to sensors once again mounted on the apparatus itself . this method overcomes the deficiencies of drip chambers . separation diaphragms are subject to error when the travel of a diaphragm is restricted . this necessitated a complicated diaphragm positioning system if the system is used for a substantial duration of time . also , a substantial area of a diaphragm ( typically 2 - 3 cm in diameter ) is required to ensure reliable transmission of pressure . at a low blood flow it is likely that a stagnant zone will form inside the diaphragm chamber that will eventually lead to clotting . the present invention utilizes flow through disposable pressure sensors . this sensors are a part of the disposable blood circuit . they do not disturb the laminar blood flow inside the blood line since the internal diameter of the sensor element is the same as of the blood tubing ( 3 mm ). the sensing element is less than 2 mm in diameter and is embedded flush in the wall of the sensor housing . the housing is bonded flush with the internal wall of the blood line tube to form a continuous channel . although similar disposable blood pressure sensors ( such as ones made by merit medical of utah ) are used widely for invasive blood pressure measurement this design has never been previously used in an apparatus for fluid removal . the present invention is intended to provide safe , controlled fluid removal in patients with fluid overload for up to 8 to 24 hours . these patients all suffer from decompensated chronic chf and are admitted or on the verge of admission to a hospital . regardless of the exact nature of their disease theses patients present at the hospital with a number of symptoms that manifest fluid overload and result in difficulty of breathing and pulmonary edema require immediate treatment . these patients are typically 5 to 20 kg over their dry weight and , if treated with diuretics , can tolerate fluid removal rates of up to 0 . 5 l / hour for until symptoms are relieved . the intended use of the present invention is to assist in the initial removal of 2 to 4 liters of fluid that should result in the relief of symptoms and much improved responsiveness to medication . the present invention can be performed by a physician or nurse trained in the use of the device . treatment can be performed in the setting of a monitored hospital floor , outpatient clinic or emergency room . the present invention is prescribed by a cardiologist . the main idea behind the present invention is to remove excess water from the patient &# 39 ; s blood using a well - accepted ultrafiltration technique at the same rate at which the surplus fluid can be recruited from the tissue . the intended use of the present invention is slow continuous removal of fluid by ultrafiltration of blood . excessive removal of fluid can lead to hypotension and serious risks to health . if the fluid is removed from vascular space too fast , it is equally dangerous and can lead to hypotension . the principle method of treatment with the present invention is to remove fluid at a rate that allows vascular blood volume to be replenished with water that has accumulated in the interstitial space as a result of the patient &# 39 ; s condition . patients that should be treated by the present invention are typically 10 to 20 kg over their dry weight due to this excess water . potential excessive water loss or gain is a recognized hazard associated with rrt . modern crrt machines used in scuf or cvvh mode can potentially remove and replace tens of liters of fluid from patient in a space of several hours . as a result , even a small error in fluid balance can result in severe risks to a patient . to prevent this from happening accurate ultrafiltration controllers are used that are based on continuous measurement of the weight of extracted and infused fluids . ultrafiltration rate is adjusted accordingly by controlling the speed of an ultrafiltration pump that can apply negative or positive pressure to the ultrafiltrate side of the filter membrane . in the case of the present invention , fluid gain is not a risk . the present invention is designed for fluid removal only . to prevent excessive removal of fluid from the patient , the present invention relies on a number of inherently safe features and materials properties rather than the ultrafiltrate pump controller . ultrafiltrate rate ( uf ) is a function of tmp , permeability of the membrane and the membrane surface . membrane surface is a constant and in the case of the present invention is 0 . 1 m2 . permeability of a filter membrane is expressed as ml of ultrafiltrate per hour per unit of the membrane surface area , and may be 30 to 33 ml / hour / m2 / mmhg . the resulting kuf of the filter may be 3 ml / hr / mmhg . kuf of a filter can decrease during the operation owing to the fouling of the membrane but can not increase unless the membrane is broken . equation 2 takes into account the affects of oncotic pressure on the ultrafiltration rate . the kuf of the membrane is determined using animal blood at standard conditions such as hematocrit of 27 %, temperature of 37 ° c . and appropriate concentration of protein and electrolytes . although these conditions do not perfectly reflect clinical conditions in all patients it is a useful engineering approximation . in the present invention , tmp is a function of blood flow and the resistance of circuit elements downstream , including the filter . tmp can be calculated in real time by the microprocessor using equation 3 from the readings of pressure transducers . where pp is the pump ( pre filter ) pressure , pr is the return ( post filter ) pressure and pg is a pressure generated by the weight of the column of ultrafiltrate . given the unadjustable design of the ultrafiltrate circuit , pg is a constant . for the 20 cm level difference between the filter and the level of fluid in the bag pg is 17 mmhg . substitution of the calculated tmp into equation 2 gives a reasonable estimate of the ultrafiltration rate . during the use of the present invention the operator sets the maximum allowed rate of ultrafiltration in ml / hour . values between 100 and 500 ml / hour are allowed . the present invention microprocessor establishes the rate of withdrawal of blood in the range of 40 to 60 ml / min . this flow rate is determined by the quality of access . it is advantageous to establish and maintain blood flow constant . based on the pressure sensor readings the tmp is calculated . this allows the calculation of ultrafiltrate rate for known kuf of the filter . if the ultrafiltration rate is higher than desired it is reduced using the solenoid ultrafiltrate pinch valve 213 in fig2 . when the valve 213 is closed the pressure inside the ultrafiltrate compartment of the filter 6 rises rapidly until it is equal to the pressure in the blood compartment ( fibers ). when the system is in equilibrium , no ultrafiltration occurs . the pinch valve 213 is cycled approximately once per minute . the duty cycle ( ratio of open to closed state ) is calculated ratiometrically from the actual and desired ultrafiltration rate . traces on fig3 illustrate how the desired average ultrafiltration rate of 250 ml / hour is achieved with a 50 % duty cycle control . in fig3 , graph 300 shows time 301 versus ultrafiltrate flow rate 302 . times t 1 306 , t 2 307 , t 3 308 , t 4 309 , t 5 310 and t 6 311 are evenly spaced with equal periods between them . the valve position 303 shows during which periods the valve 213 is open 304 and closed 305 . for example , between time t 1 306 and time t 2 307 , valve 213 is open , but between time t 2 307 and time t 3 308 , valve 213 is closed . for example , when the valve is opened ultrafiltrate flow is 500 ml / min , shown by lines 312 . when the valve is closed it is zero , shown by lines 313 . thus , using equal periods of opened valve 213 and closed valve 213 will effectively achieve an average flow rate equal to half of the open valve flow rate . other factors limiting the ultrafiltration rate are the amount of blood available for filtration and the hematocrit of the blood entering the filter . equation 4 can be used to calculate the return blood hemotocrit hr from the blood flow bf , ultrafiltrate flow uf and withdrawal blood hemotocrit hw . the safe pump 204 can deliver up to 60 ml / min of blood . the maximum amount of water that can be extracted from the blood is determined by the hematocrit of the return blood . as the water is extracted , hematocrit increases . this in turn increases the viscosity of blood . blood entering the filter typically has hemotocrit of 35 to 45 %. the corresponding viscosity is approximately 3 to 4 mpa - s . viscosity of blood with the hematocrit of 60 % is 6 mpa - s . at 70 % hematocrit blood becomes too viscous to pass through the filter or the return needle . the volume of the ultrafiltrate collection bag 215 determines total amount of fluid removed from the patient . the bag 215 is connected to the fluid removal port 211 of the filter 207 by the bonded pvc tube 212 . the ultrafiltrate collection system is sealed . when the bag is full , the pressure in the bag 215 rises until it is equal to the average pressure of blood in the filter fibers . ultrafiltration is stopped until the bag is emptied by the operator by opening the ultrafiltrate valve 216 . the volume of the ultrafiltrate bag 215 is one liter . this automatically limits the maximum amount of water that can be removed from the patient without an operator &# 39 ; s interaction . bag 215 is transparent and has volume marks that allow user to read the actual amount of fluid removed in preferably 100 ml increments . blood exiting the filter 207 through the connector on the bottom of the filter casing is continuously returned to the patient through the return needle 210 . blood flow leaving the filter is the same as the blood flow entering the filter if the ultrafiltrate clamp 213 is closed . if the clamp 213 is open ultrafiltration occurs and the blood is continuously fractured into the ultrafiltrate and more concentrated blood . the hematocrit and the volume of returned blood are determined by the ultrafiltration fraction , which is the volumetric fraction of the ultrafiltrate relative to the volume of whole blood entering the filter . blood return circuit pressure sensor 209 serves several functions . the return pressure is used in the tmp calculation that is in turn used to calculate input data for the control of the ultrafiltration rate . it is also used to detect a disconnected or occluded circuit . excessive pressure signals the occlusion that can be caused by a kinked tube or a clotted needle . since the resistance of the needle 210 is much higher than the resistance of the blood line 208 , disconnection of the needle from the tubing is easy to detect from an abrupt drop of the return pressure . fig4 shows another embodiment of the invention . this embodiment is preferred when more precise removal of fluid is desired . an ultrafiltrate pump 17 is used to pump and meter ultrafiltrate into the collection bag 13 . weight scale 18 allows additional control and monitoring of the amount of the removed fluid . ultrafiltrate pressure sensor 15 is used to monitor the tmp and to detect clotting or fouling of the filter . blood leak sensor 16 is used to detect the leakage of red blood cells across the filter membrane if the membrane is damaged . the blood sensor is of photometric type and responded to the change of color of the ultrafiltrate . fig5 illustrates the effect of the reduction of the filter membrane area on the transient time of blood in contact with plastic . experiments were performed using the thermodilution method . at the time mark 0 iced water flowing through the system at 60 ml / min replaced the room temperature water . three curves show the time delay before the cold water reached the exit of the circuit ( simulated return of blood to the patient ). curve 510 corresponds to a continuous 2 meter long tube with internal diameter of 3 mm with no filter at all . curve 520 is the filter described in this invention . curve 530 is the standard f40 filter by frescenious . traces show that the time of contact of blood with plastic during the treatment with proposed invention will be significantly shorter than that with a conventional dialysis machine . fig6 proves that the desired amount of ultrafiltrate can be removed with the filter described in the preferred embodiment . curves 560 , 550 and 540 correspond to blood flows of 40 , 50 and 60 ml / min accordingly . animal blood with hematocrit of 27 % at the temperature of 37 ° c . was used in the experiment . tmp was between 100 and 250 mm hg and shear rate of blood inside the filter was maintained in the 1000 to 2000 sec - 1 range . the desired amount of 5 - 10 ml / min of ultrafiltrate was filtered out of blood before it was returned into a vessel using a 18 gage needle . filter had surface are of 0 . 1 m2 with membrane permeability of 33 ml / hr / m2 / mmhg supplied by minntech . filter bundle was made out of 700 fibers 22 cm long as described previously in this invention . fig7 shows yet another potential embodiment of the invention . this apparatus is of a batch type and uses one needle 60 . blood is withdrawn from the vein and stored in the bag 51 . during the withdrawal phase roller pump 52 rotates counter clockwise . approximately 50 ml of blood can be stored in the bag 51 . valve 59 is open and valve 58 is closed so that the filter 54 is bypassed . blood is then ejected from the bag 51 through the filter 54 back into the patient &# 39 ; s vein through the needle 60 . valve 58 is opened and valve 59 is closed and the pump 52 rotates clock wise . during this phase filtration occurs and the ultrafiltrate is collected in the bag 55 . pressure sensors 53 and 56 are used to control the process so that vein does not collapse or over - extend and tmp is not too high . air detector 57 is used to protect the patient from embolism . an anti - coagulant may be needed in view of the temporary storage of blood in the bag 51 . the invention has been described in connection with the best mode now known to the applicant inventors . the invention is not to be limited to the disclosed embodiment . rather , the invention covers all of various modifications and equivalent arrangements included within the spirit and scope of the appended claims .
a fluid removal apparatus comprising a blood removal catheter for insertion into a peripheral vein or artery and having a size 16 standard gage needle or less ; a filter having a blood inlet port coupled to the blood removal catheter , a blood outlet port , an excess fluid removal port , and a blood flow passage with porous membrane which passes fluids to the fluid removal port and retains solutes of 60 , 000 daltons or greater , and a blood return catheter for inserting into a peripheral vein or artery and having a size of 16 standard gage needle or less .