Disposable extracorporeal conduit for blood constituent monitoring

DESCRIPTION

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EXTRACORPORAL CONDUCT FOR BLOOD CONSTITUTE CONTROL

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1. Field of the Invention. ,

The present invention is directed to a device and system for obtaining a desired value of the concentration of biological constituents present in a pulsatingly flowing liquid, and more particularly directed to such obtaining in pulsatingly flowing blood in an environment. hemodialysis

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2. Basic Technique

The kidneys are located on either side of the spine. In a healthy patient, the kidneys work to stimulate red blood cell production and regulate blood content. The kidneys also produce hormones, which affect other organs and control growth. When functioning properly, the kidneys serve as a means of cleaning the blood by removing excess fluid and toxins. The filtration task is partially performed on each kidney by about one million nephrons. Nephrons are filtration units made up of tiny blood vessels. Each of such blood vessels is called a glomerulus. Every day will be processed by the kidney approximately 0.22m³ (200 quarters) of blood and liquids. The kidney removes about 0.0022m³ (two quarters) of water and toxic chemicals which are sent to the bladder in the form of urine for subsequent emptying through urination.

A patient whose kidneys are working below normal may be dialysed as a substitute for the function of blood cleaning normally performed by properly functioning kidneys. Dialysis is a process whereby the kidney's blood-cleansing function is replaced. The dialysis process was perfected for routine use in the 1960s and was invented around 50

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years. For purposes of discussion and illustration of hemodialysis, we now refer to Fig. 1. While Figure 1 incorporates a currently preferred embodiment of the present invention, it also incorporates a view of some common components, which are typical in an environment. generic hemodialysis. The generic environment and typical components will now be discussed.

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In hemodialysis, blood is taken from a patient 200 by means of absorption catheter means, an example of which is shown in Figure 1 in the form of an inlet catheter 122. Inlet catheter 122 is intravenously inserted into patient 200 at a point 180 and is used to define a bloodstream upstream of a blood filter used to filter out impurities contained in the blood. The blood filter is also called a dialyzer 130. Dirty blood flows from a patient's artery 20 to pump means, an example of which is pump 140. From pump 140, blood flows to the dialyzer 130. The dialyzer 130 has an inlet opening 230 and an outlet opening 240. Pump 140 performs the function of moving patient's dirty blood 200 into inlet port 230 through dialyzer 130 and out of outlet port 240.

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Specifically, the dirty blood contained in the catheter 122 is transported to the inlet port 230 of the dialyzer 130. Once it has passed through and been cleaned by the dialyzer 130, the blood can be further processed, such as a heparin drip. , in component 300 related to hemodialysis. The blood, now cleaned, is returned to patient 200, after the dialysis process, by means of outlet catheter means, an example of which is outlet catheter 124. Outlet catheter 124, which is also intravenously inserted into patient 200 at point 180, defines a blood passage, which is downstream of dialyzer 130, and which carries blood exited from dialyzer 130 back to patient 200.

As noted, the hemodialysis process uses a blood filter or dialyzer 130 to clean a patient's blood 200. As blood passes through the dialyzer 130, it travels in capillary tubes (not shown) within the dialyzer 130, which serve like membrane passages to dirty blood. Capillary tubes

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/ Remove poisons and excess liquids through a diffusion process. An example of excess fluid present in dirty blood is water and an example of poisons present in dirty blood are blood urea nitrogen (BUN) and potassium.

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Excess fluids and poisons are removed by a liquid dialysis fluid, which is a solution of chemicals and water. Clean dialysis fluid enters dialyzer 130 through an inlet tube 210 from a combination of controller and tank 170. Dialysis fluid wraps around the capillary tubes inside dialysis 130 as dialysis fluid flows through the dialysis fluid. dialyzer 130. The clean dialysis fluid diffuses excess fluids and poisons through the capillary tubes, and then sends excess fluids and poisons along with the dialysis fluid out of dialysis 130 through a outlet tube 220, thus cleaning the blood. The dialysis fluid exiting the outlet tube 220, after cleaning the blood, may be discarded.

In summary, dirty blood flows from a patient's artery 200 to pump 140 and then to dialysis 130. Dirty blood flows into dialysis 130 from an inlet catheter 122 and clean blood flows out of the dialyzer 130 through a patient return catheter 124, 200.

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Hemodialysis, which removes excess fluid from a patient's blood, has an acute impact on the body's fluid balance, in part due to the rapid change in blood volume. When the rate of fluid removal is faster than the body's plasma refill rate, intravascular blood volume decreases. This resulting imbalance has been linked to complications such as hypotension, unconsciousness, headaches, vomiting, dizziness and cramps experienced by the patient both during and after dialysis treatments. With hypotension and blunt shock occurring in up to 25% of hemodialysis treatments, hypovolemia caused by dialysis remains a major complication of hemodialysis.

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Many dialysis patients already have impaired circulatory responses due to the side effects of end-stage renal disease. A malfunction of blood pressure compensatory mechanisms due to intravascular volume depletion has been considered to be one of the major causal factors of dialysis hypotension.

In order to reduce the chances of dialysis hypotension, continuous measurement of circulating blood volume can optimize dialysis therapy regimens, control fluid balance and help achieve the patient's dry weight target on a quantitative basis. Volumetric controllers, while giving an accurate measure of the amount of liquid removed by ultrafiltration, provide no indication of how body plasma refilling mechanisms are responding to actual fluid removal. Factors such as food and water absorption and postural changes also significantly affect circulating blood volume during dialysis. Maneuvers such as eating, drinking, and posture illustrate how sensitive plasma refilling mechanisms are.

The hematocrit value gives an indication about the change in blood volume. Since the number of red blood cells in whole blood is not significantly altered by dialysis and the average corpuscular volume of red blood cells remains practically constant, it follows that changes in blood volume will be inversely proportional to changes in hematocrit. Therefore, the change in the patient's blood volume can be defined at any time during the course of dialysis treatment according to EQUATION 1.

EQUATION 1

Β V final Hctjjucial

Β Initial V Hotfinal

Where:

BVfinal

Bnicniciol

HCTfinai = Final Blood Volume = Initial Blood Volume = Final Hematocrit Value

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Clinically, however, it may be more useful to determine the percentage change in blood volume according to EQUATION 2.

EQUATION 2

100% x BVf-BVj =

Hctj-1

Hctf x 100%

Where:

= Final Blood Volume = Initial Blood Volume = Initial Hematocrit Value = Final Hematocrit Value

The use of hematocrit modification as a measure of the actual change in blood volume that occurs during dialysis is known. However, in order for the relationship between hematocrit modification and blood volume modification to be useful, the hematocrit must be precisely and continuously controlled in real time throughout the hemodialysis treatment session. While accuracy can be achieved by elaborate technical means to be clinically practical, real-time hematocrit and blood volume control should be easy to use, save nursing staff time, operate non-invasively and be justifiable. in terms of costs.

Several techniques employed to control intravascular blood volume modification include, due to ultrafiltration as a function of hematocrit value, microcentrifugation, electrical conductivity and photometry.

In microcentrifugation a microcentrifuge is used to measure the hematocrit. This process is unsuitable for controlling changes in blood volume in real time due to the amount of time that elapses between measurements, the large potential for sampling and reading error and the need to properly compensate for plasma trapped in red blood cell columns. Thus, due to the

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In the labor intensive nature of centrifuging the patient's blood samples on a time basis, this technique is totally inappropriate, impractical and too expensive for large-scale clinical application.

In an attempt to obtain hematocrit information in real time, electrical conductivity measurements have been used. Conductivity measurements are, however, adversely affected by abnormal electrolyte, anticoagulant and protein concentrations, which are all prevalent among dialysis patients. Given this, this particular technique is also full of significant technical errors.

Optical techniques, although generally unaffected by the above problems, have been susceptible to other instabilities. These include ambient light variations, intubation devices, changes in blood flow rate, inline pressures and oxygen saturation. In addition, light sources used in optical techniques require frequent calibration.

BRIEF SUMMARY OF THE INVENTION

According to the invention, as embodied and widely described herein, the invention is directed to an apparatus and a system incorporating the apparatus. The device is a disposable extracorporeal cuvette through which a pulsatile fluid flows.

The cuvette has an inlet opening, also called a first fluid conduit, and an outlet opening, also called a third fluid conduit. Between the inlet port and the outlet port there is a conduit, also called a second fluid conduit, which is in fluid communication with the inlet port and the outlet port.

The conduit has two opposite walls, which have a predetermined separation from each other. The two opposing walls constrain at least some of the fluid pulsatile flowing through the conduit. The conduit also has transducer means located on one of the opposite walls. The transducer means vary the predetermined separation between the two opposite walls in response to pressure pulsations in the pulsatile fluid.

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Alternatively, the conduit may be stated to have means, in contact with the pulsatile fluid flow contained therein, to make an elastic fluctuation with each pulse of the pulsatile fluid flow present within the conduit. Elastic fluctuation is in a direction which is essentially normal or perpendicular to the general direction of movement of the pulsatile fluid flowing within the conduit. The elastic nature of the buoyancy means ensures their return to their original position after performing the buoyancy.

The conduit, also called the second fluid conduit, is made of materials that allow at least two predetermined wavelengths of electromagnetic radiation, and preferably four wavelengths, to pass transmissively through them. The materials used are preferably inexpensive, so that the cuvette is economical both to be disposed of after use and to be manufactured.

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The term "fluid" is intended herein to alternatively mean liquid or gaseous substances. Also intended herein is that "pulsatile" means rhythmic or cyclic eruptions or pressure increases in a fluid flowing under pressure.

The cuvette incorporating system is intended to control the concentration of a particular or desired biological constituent. Preferably, the pulsatile fluid flowing is the blood of a patient and the desired biological constituent concentration to be controlled in the system is the concentration of red blood cells, also expressed as hematocrit value. The preferred system, in which constituent concentration control is performed, is a hemodialysis system, wherein the hematocrit value is controlled both before and after the blood cleaning process as a means of obtaining volume modification. during the hemodialysis process.

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Constituent concentration control calculations are performed using the technique described in PCT Patent Application Serial No. PCT / US93 / 03427 filed April 12, 1993 entitled "NON-INVASIVE CONTROL SYSTEM AND METHOD" SYSTEM AND METHOD FOR NONINVASIVEHEMATOCRITMONITORING), and is hereinafter referred to as The Embedded Technique.

The Incorporated Technique, known as noninvasive differential raciometric spectrophotometry, is described as follows. It is assumed that incident radiation passing into living tissue will pass through a combination of blood, tissue, and interstitial fluid compartments. Light attenuated by such living tissues can be expressed by means of the modified Beer-Lambert equation:

I = Io ~^(andb (2)

EQUATION (2) can also be written

In (I / Io) = - (and_B QC + Xv) + et Xt + (Xi) d + G (2a)

Where eb, and_t ee, represent the extinction coefficient, respectively in blood, tissue and interstitial fluid compartments; X_The and X_v represent the concentration of arterial and venous blood (Xb = X_The + X_v ), Xt represents the concentration of tissue absorbers, and Xi represents the relative concentration of water and dissolved components in the interstitial fluid compartment; d represents the intrasensor separation; and G is a constant of geometric configuration.

As the blood layer pulses, the terms of concentration change. The term d may be fixed by the geometrical configuration of the device. Taking the partial derivatives of equation (2) with respect to time and dividing them by equation (2) gives:

Sl / ôt

-______=₍and_B(ax_Q / at. + ax_v / at_} + e, sx, / 9, + e_; ax_f / ôt&gt; d + ôG / at (3)

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which can be simplified in each compartment and wavelength by letting X = X / 1, and G = G / 1, and

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λ = (e_B (X '_The + X '_v) + e_t X '_t +_HeyX'i) d + G '(4)

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Assuming that X and G do not vary significantly over the pulse time interval, then G = 0 and X _t = 0 and equation (4) can be simplified to

V \ = (e_B(X '_The+ X '_v) + e_iX '_i) d (5)

Examining the transport between X_The and X_v, a proportionality constant K can be formed, such as X_v = K_vXa, which represents the reactionary nature of the venous component and further reduces the above equation to

V \ = (e_B(lK_v) X '_The+ eiX'i) d (6)

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Given that X_The and Xj are not wavelength dependent (λ), Vx values ​​at different wavelengths can be differentially subtracted to produce an independent hematocrit term, which contains only ejXi information. Although the term V gos / V 1310 provides useful information regarding changes in hematocrit, it should be recognized that the simple V sos / V mo ratio is not sufficiently accurate to determine the hematocrit value unless the term GiX; be known or deleted. For example, the term GjX isos can be ignored since G805 is extremely small, while the term ejX imo is about 25% -50% of the value and_Bof the blood itself and can therefore not be ignored without affecting accuracy.

Figures 15 and 19 suggest that the linear combination of Vx at λ = 805 nm and λ = 970 nm will never have a near constant value for a range of Hct values. Given that

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extinction coefficients e, 805 and 61970 are well known, or can be empirically determined, a precise proportionality constant Rj can be found to produce

6.970 X'i-V'970 "Ri V * 805 (7)

This correction term can now be applied with a second proportionality constant R₂ (where R2 is approximately equal to Gibius / 6.970) for the term V i3! 0, in order to remove exactly its sensitivity 6ϋ₃ιοΧ i, so what

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6_Bi3io (l -K_v) X '_The = V '₁₃io-R2 (V'970-Ri V '_8O5) (8)

This corrected term can now be used raciometrically with V 805 to remove (lK_v) Xa and leave the pure ratio of the extinction coefficient represented by Equation (9) below and graphically shown in Figure 16.

6b805 V'8O5 (9)

Gbl310 V'1₃10 ~ R2 (V'970 — Ri V'8O5)

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It should be noted that the following assumptions and requirements are essential in hematocrit determinations (but in the case of pulse oximetry these requirements may not be of the same degree of significance).

A. Although the wavelengths λ = 805 nm and λ = 1310 are of nearly equal specifications, the true function of e with respect to the hematocrit at each given wavelength must contain the hematocrit information that is different in curvature. , or the spacing, or the linearity, or the sign, of the other. See Figure 15. If the 6χ functions are not sufficiently different from the hematocrit, then the ratio 6bxi / ^ bu will not contain hematocrit information. See Figures 20A and 20B and Figures 21A and 21B. Although the previous discussion refers to the equal specification wavelengths of λ = 805 nm and λ = 1310 nm, it should

I ΐ

¹¹ consider that other wavelengths of equal specifications, such as λ = 570 nm, λ = 589 nm and λ = 1550 nm may also be used.

B. In addition, the wavelengths should be selected sufficiently close to each other such that the optical passage lengths d are approximately the same. Longer wavelengths are preferred as they have lower dispersion sensitivity, s:

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(10)

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C. The geometric or spatial relationship of emitters and sensors is important. For example, if vertically aligned emitters are used in an ear lobe measuring device, then the highest emitter may illuminate a different amount of blood-filled tissue than that illuminated by the lower emitter. If only one sensor is used, then there will be a disparity between Xb at each wavelength. In addition, the sensor - emitter spatial separation distance is very

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important because the pressure applied to the tissue between the sensor and the emitters affects the yielding of the arteriolar and capillary vessels. This modifies ο X as the pressure (or distance) changes. This modification in X therefore modulates the Vx function. Therefore, the distance of the sensor - emitter separation must be such that the pressure applied to the earlobe, fingertip or other body member does not affect the Vx function. This sensor separation distance is empirically determined and should generate less than 40 mm Hg of applied transparietal pressure.

A horizontal alignment of the emitters relative to a single sensor can be arranged so that the emitters and sensors illuminate and detect identical regions of Χχι and Χχ2. It is important to note that the term d, the sensor - emitter separation, will differ between λι and λ₂ cosine of the angle between the sensor and the emitter. Therefore, if any misalignment with respect to normal occurs, the term d will not be canceled to obtain equation (9).

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The preferred arrangement is where all emitters (660, 805, 950 and 1310 nm) are placed on the same substrate. This is preferred because the emitters will then illuminate virtually the same X region._B.

D. For reflectance spectrophotometry, an aperture is required for the sensor and each emitter. A sensor - emitter separation is also required so that the reflectance of the first tissue layer, Rt (a non - blood epithelial layer) does not further exaggerate a multiple scattering effect, i.e. the total reflectance, R, measured. would also contain spurious information of the layer reflectance

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epithelial, where:

TS-Rb

R = R_t + ------ (1-Rb'Rt) (11) where R is the total reflectance, R_t is the reflectance due to the first layer of epithelial tissue, Rt, is the reflectance due to the blood layer, and T_t It is transmission through the first layer of fabric.

The reflectance equations that describe R_t or R_B should now be the sum of the total light the sensor detects, that is:

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R_B - Jff (source function) (scatter function) (12)

Although equation (9) describes the noninvasive hematocrit device theory, the four assumptions (AD) are important for the repeatability and accuracy of hematocrit device functioning.

Assuming that items A through D have been adequately addressed, (9) we take:

and_Bxi _ (si + ki)

SbX2 (s₂ + k₂) (13) where s is a scattering constant and k is an absorption constant, and where in whole blood:

s = σ₅ Hct (1-Hct)

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k = σ₃ Hct (at wavelengths of equal specifications) (15) where σ₅ is the perpendicular section of the dispersion and the_The It is the perpendicular section of the absorption.

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From the above it is taken that g, the extinction coefficient, is not a simple function of the absorption coefficient, k, normally determined in pure solutions. Rather, it contains a diffusion or dispersion term, which should be considered in a non-pure solution medium, such as whole blood and tissue.

Finally, substituting for (14) and (15) in (13):

£ λΐ _ σ₅ι (l-Hct) + a_Thei ei2 a_s2 (1-Hct) + aa2 (16)

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Therefore, the ratio gj.i / εχ2 is a function of the hematocrit. From Figure 16, a reference table or polynomial curve fitting equation can be obtained and used for the final presentation of hematocrit results. Knowing the true value of hematocrit, it is easy to see (Figure 14) that a wavelength at 660 nanometers can be chosen to obtain a g ratio from which the oxygen saturation value is derived, independent of the hematocrit. For example, equation (16) would be:

Sb660 σ₅660 (1-Hct) + Oa660 + S_The O₂ (σ_αο660 - ^) ^andb805 &lt;TS8O5 (1'Hct) + Oa805⁺ Are₂ (The_to805 -Caseoi) (17)

Equation (17) simultaneously shows the mutual dependence of hematocrit and oxygen saturation.

Figure 18 graphically demonstrates the need for a hematocrit-independent blood saturation device. As the hematocrit value or the

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oxygen saturation percentage decreases, the saturation percentage error becomes unacceptable for clinical use. It is not uncommon, for example, to see patients with a low hematocrit (about 20%) who also have breathing problems (low oxygen saturation). Therefore, the clinician simply needs more accurate oxygen saturation values.

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Knowing the hematocrit and oxygen saturation values, the calculation of Oxygen Content is trivial, and can be presented directly (a value hitherto unavailable to the clinician as a non-invasive, continuous, real-time result):

[Oxygen Content] = Hct · S_TheO2 · K (18) where K is an empirically determined constant.

Referring to equations (16) and (9) a decision must be made by the computer as to whether to use the Taylor expansion approximation for the logarithm. This algorithm is retained in programs as a qualifying decision for averaging and reading algorithms. Taylor's approximation is valid only for small I / t values.

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The above techniques describe conditions and equations in which wavelengths of equal specifications are chosen such that the hematocrit value obtained has no interference with oxygen saturation, therefore an independently determined hematocrit value.

One can, however, also choose Δ2 (the reference wavelength) in equation (13) at 1550 nm. In the radiation region of 900 to 2000 nm, the blood absorption coefficients depend on hematocrit and water, while at 805 nm the blood absorption coefficient only depends on hematocrit. Therefore, using in combination, wavelengths of 660, 805, and 1550 will also provide a technique for determining hematocrit (egos / ei).₅₅o) and oxygen saturation (β6βο / ^ 805) 15

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The present invention may be applied to the determination of other components (including but not limited to glucose or cholesterol) in any range of the electromagnetic spectrum in which the techniques may be used.

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spectrophotometric.

Constituent concentration calculations, described in the Incorporated Art, are performed on derived data by passing multiple lengths of,_t light wave through the conduct of extracorporeal blood or a body part of a patient. Means for providing and detecting such multiple light wavelengths and for analyzing the various light intensities detected are also described and provided by the Incorporated Art. The spatial arrangement for both detecting and emitting light is detailed to provide optimal repeatability of the signals and data derived therefrom. Finally, memory and calculation means are included which are capable of storing, manipulating, displaying and printing the detected signals in a variety of ways. The Embedded Technique allows an end user, nurse, clinician, or patient to guarantee a desired value of the concentration of biological constituents, such as the hematocrit value, oxygen saturation value, or oxygen content value, by means of their presentation in the form of digital values ​​in real time.

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Although the apparatus of the present invention is preferably applicable in the area of ​​renal dialysis, it may also be employed in cardiovascular surgery or in other fields of medicine where blood is present in extracorporeal tubing. In these environments, the present invention derives hematocrit value, blood oxygen saturation, blood oxygen content, and change in blood volume. These leads are all made by the present invention without the need for an invasively obtained stagnant blood sample. Advantageously, deriving the hematocrit value in such cases provides non-invasively and continuously repeatable and reliable determinations of a patient's hematocrit value, regardless of the patient's perfusion status or cardiac output.

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BRIEF DESCRIPTION OF DRAWINGS

In order to more fully understand the manner in which the above and other advantages of the invention are obtained, a more particular description of the invention will be provided with reference to specific embodiments thereof which are illustrated in the accompanying drawings. In the understanding that these drawings only represent typical embodiments of the invention and are therefore not considered to limit its scope, the invention, in its present form best understood to make and use thereof, will be described herein. additional specificity and detail through the use of the accompanying drawings, in which:

Figure 1 is a view of an environment of a patient undergoing hemodialysis treatment and shows a system incorporating principles of a presently preferred embodiment of the invention including a pair of cuvettes having spectrophotometry components thereon.

Figure 2 is an enlarged perspective and partial cross-sectional view of a cuvette shown in Figure 1.

Figure 3 is a perspective view of Figure 2, with the various respective components turned off.

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Figure 4 is an exploded view of the cuvette shown in Figure 3.

Figure 5 is a cross-sectional view taken along Section Line 5-5 of Figure 3.

Figure 6 is a cross-sectional view taken along Section Line 6-6 of Figure 3 showing the fluid flow passages within the cuvette.

Figure 7 is a cross-sectional view taken along Section Line 7-7 of Figure 3.

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Figure 8 is an open perpendicular cross-sectional view taken along Section Line 8-8 of Figure 2, showing radiation at four (4) wavelengths toward a flexible wall within the cuvette.

Figure 9 is an exploded view of an alternative embodiment of the cuvette shown in Figure 3.

Figure 10 is an open perpendicular cross-sectional view of an alternate embodiment of the cuvette taken along Section Line 8-8 of Figure 2, representing radiation for four (4) wavelengths directed toward a Flexible silicone membrane inside the cuvette.

Figure 11 is a graph of the ratio ebi / eb2 to Hematocrit, where ebi is the extinction coefficient of whole blood in radiation with a first radiation wavelength and eb2 is the extinction coefficient of whole blood at a radiation of one. second radiation wavelength, wherein the first and second wavelengths are wavelengths with equal specifications such as that of Figure 11 which represents a strong hematocrit function.

Figure 12 represents the variable physical relationships between the component dimensions of the cuvette shown in Figure 8, including the thickness of the blood layer and its modifications (from Ad, respectively), the thickness of the disposable membrane (T) and the area of ​​the membrane (A).

Figure 13 represents the error in approximation ΔΙ / I as a function of the changes Ád in the thickness of the blood layer (d) within the cuvette.

Figure 14 is a graph showing the optical absorption coefficients of oxyhemoglobin (HbO₂), reduced hemoglobin (Hb), and water (H₂O) in relation to wavelength.

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Figure 15 is a graph showing the relationship between the light extinction coefficient at three different wavelengths relative to the whole blood hematocrit,

Figure 16 is a graph showing the relationship between the ratio of extinction coefficients of two rays having different wavelengths relative to the hematocrit.

Figures 17A-17E provide a flowchart showing steps taken during a presently preferred method of the Incorporated Technique using a pulsatile component of the subject's blood flow to provide accurate hematocrit and oxygen saturation values. Blood

Figure 18 is a graph showing the variation of oxygen saturation as a function of hematocrit.

Figure 19 is a graph of the_B805 / e_B970 in relation to hematocrit.

Figures 20A-20B are graphs of and relative to Hematocrit, at two non-preferred wavelengths and ei / e₂ relative to Hematocrit at these non-preferred wavelengths.

Figures 21A-21B are graphs of and relative to Hematocrit at two non-preferred wavelengths and ei / e₂ relative to Hematocrit at these non-preferred wavelengths.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present invention contemplates a system and apparatus therein for determining the concentration of a biological constituent present in the blood of a patient undergoing hemodialysis treatment.

1. The system

The process and generic environment of hemodialysis are seen in Figure 1 and described above. A summary of this process is that patient 200, whose kidneys are functioning below normal, is dialysed. Dirty blood flows from a patient's artery 200 to 140 and then to the dialyzer 130. Dirty blood flows into dialyzer 130 from inlet catheter 122, and then clean blood flows out of dialyzer 130 through outlet catheter 124 back to patient 200.

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It is preferable that the pump 140 causes blood flowing in, through and out of the dialyzer 130 to flow pulsatably.

Installed at both ends of the dialyzer 130 are spectrophotometry means for defining a blood flow passage to emit radiation into the blood contained in the flow passage and to detect radiation passing through both the blood and the bloodstream. of the flow passage. Spectrophotometry means include cuvette means for defining blood flow passage, and emitter / detector means for directing and detecting radiation. Within the emitting / detecting means are simultaneously emitting means for directing radiation and detecting means for detecting radiation.

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In the preferred embodiment, as shown in Figures 3 and 8, an example of the sender / receiver means is represented by the sender / receiver apparatus 100. An example of the transmission means is indicated by a photo sender 102. The sender / detector apparatus 100 has also detecting means, an example of which is depicted as a photodetector 104. An example of the cuvette means is shown in Figures 3 and 8 as a cuvette 10.

The emitter / detector apparatus 100 allows photodetector 104 to detect the portion of radiation that is directed by the photo-emitter 102 to the cuvette 10 and passes through both the blood contained therein and the cuvette 10.

As shown in Figures 2 and 8, a cuvette 10 is installed at both ends of the dialyzer 130. Each cuvette 10 has a photo emitter 102 and a photodetector 104 installed therein. In the preferred embodiment of the system, photo emitter 102 and photodetector 104 are shown to be held together by spring-loaded type C forceps in a photo emitter / detector apparatus 100.

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The transmitter / detector means are electrically connected to the calculation means. In a preferred embodiment of the system, an example of calculating means is shown in Figure 1 as computer 150, which is electrically connected to photo-emitter 102 and photodetector 104 in transmitter / detector 100 via cable 120.

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The inlet catheter 122 carries the blood to the cuvette 10, located before the inlet opening 230 of the dialyzer 130. The emitter / detector apparatus 100 in the inlet opening 230 of the dialyzer 130, subjects the blood contained therein to radiation with at least two wavelengths of electromagnetic radiation for spectrophotometric analysis so that the concentration of a desired biological constituent can be obtained. Each of the photodetectors 104, at both the inlet opening 230 and the outlet opening 240 of the dialyzer 130, communicates the detected radiation of at least one first and a second wavelength through cable 120 to the computer 150.

Computer 150 calculates, both before and after dialysis, the concentrations of the desired or desired biological constituent. The computer 150 then displays, respectively on a first display 152 and a second display 154, the concentration of the biological constituent obtained in analog or digital representations. The calculation means shown here, for example, as the computer 150, preferably has the multiple capability of simultaneous real-time computing and display of hematocrit and oxygen saturation values, as well as percent change in blood volume in a patient undergoing hemodialysis.

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The choice and predetermination of radiation wavelengths is based on the biological constituent from which the concentration value is to be determined. Photo-emitter 102 preferably emits and photodetector 104 preferably detects four (4) predetermined wavelengths for spectrophotometry techniques taught in the Incorporated Art. Accordingly, the cuvette 10 will preferably be made of materials which allow the four (4) predetermined direct radiation wavelengths to pass therethrough.

<img file="PT700268E_D0048.tif" />

2. The device

a) A Preferred Embodiment

In the presently preferred embodiment, an example of the cuvette means is the disposable fluid cuvette 10 of Figures 1 to 8. The inlet and outlet for the cuvette are respectively indicated at 16 and 18, including a cylindrically shaped portion of the cuvette 10, herein called the conduit, or alternatively the second fluid conduit.

<img file="PT700268E_D0049.tif" />

As shown in Figure 4, there is an upper compartment assembly 12 which is mounted on the lower compartment assembly 8 to form the cuvette 10. The upper compartment assembly 12 may be installed on the lower compartment assembly 8 by middle of an adhesive element. Other and equivalent means such as friction welding or ultrasonic welding may also be employed. The purpose of properly sealing the upper housing 12 against the lower housing 8 is to create an impermeable, fluid-sealed connection between them so that fluids conducted through the cuvette 10 do not flow, infiltrate, or subsume at the connection points between the upper compartment 12 and lower compartment 8. Lower compartment 8 has handles or tabs 14 through which cuvette 10 can be manually moved.

The conduit incorporates transducer means. As stated, the transducer means vary the predetermined separation between the two opposite walls with each pulse pressure in the fluid. In the presently preferred embodiment, an example of \ * Α.

transducer means is shown in Figures 4, 5, 7 and 8 as wall 30 having a wall 32 opposite it.

<img file="PT700268E_D0050.tif" />

The pulsatile flow of fluid flows in the conduit within the limited zone between a vertical wall 46 and opposing walls 30 and 32. The fluid contained in the conduit assumes the flow passages indicated by arrows 40 and 42 in Figure 6. A The flow passage seen in Figure 8 at 44, which describes the flow of fluid as it enters the inlet port 16 into the conduit, will assume the flow passage 40 within. area 36 or alternatively will assume the passage of flow 42 within area 34. Fluid flowing between opposing walls 30 and 32 is described by the area indicated by 36. Figure 12 represents area 36 by Ά '. Fluid flowing out of opposite walls 30 and 32 flows in the direction indicated by arrow 42 in the area marked 34 in Figures 6 and 7. The volume of fluid in area 34 is preferably larger than that of area 36.

<img file="PT700268E_D0051.tif" />

Inlet port 16 and outlet port 18 are linearly aligned on both sides of the duct and share a common longitudinal axis passing between them. The cylindrical conduit between inlet port 16 and outlet port 18 has a longitudinal axis passing through opposing walls 30, 32, which is normal to the common axis of inlet ports 16 and outlet 18. Figure 12 represents the wall 30 as the membrane thickness 'Τ'. As shown in Figure 8, the opposite wall 30 is preferably thinner than the opposite wall 32.

As the fluid contained in the flow passage 44 causes a pulsating pressure eruption, the opposite wall 30 flexes to an arcuate shape while the opposite wall 32 remains relatively still. The deformation of the wall 30 is represented by the bending line 62 seen in Figure 8. After the pulse pressure of the fluid contained in flow passage 44, wall 30 returns to its position shown in phantom, so that the separation between wall 30 and 32 is indicated by distance 60. The relationship between distance 60 and bending line 62 is graphically represented in Figure 12, respectively by Ad.

As noted, wall 30 is an example of transducer means. Other equivalent functioning transducer means may be incorporated into the conduit so that the distance represented by line 60 in Figure 8 between opposing walls 30 and 32 may be varied. By way of example, and not by way of limitation, the transducer means may be constructed essentially of silicone.

<img file="PT700268E_D0052.tif" />

Alternatively, a small section of the opposite wall 30 may be constructed to vary the distance between the opposite walls 30, 32. Such a portion may be spring-biased or have other flexible means for turning the small portion of the wall 30 back. to its original position of least pressure, prior to the thrust. In such an embodiment, the opposite wall 30 need not be thinner than the opposite wall 32.

<img file="PT700268E_D0053.tif" />

Figure 8 depicts optical passages 64, 66, 68, and 70 of directed electromagnetic radiation at four (4) different wavelengths. Each wavelength is selected for spectrophotometric compatibility with a specific biological constituent in the pulsatile fluid flow. As described, the detected portion of the directed radiation is used to obtain the concentration value using the Incorporated Technique. The wavelengths may be set in the emitter / detector apparatus 100 or may be set by adjusting computer 150, where photo-emitter 102 and photodetector 104 are dynamically adjustable to wavelength by computer 150.

The opposite wall in the conduit incorporating the transducer means also has means for receiving emission means projecting from the wall. The other of the opposite walls has means for receiving detecting means. In the presently preferred embodiment, the means for receiving a photo-emitter are shown in Figures 3, 4, 5 and 7 as a combination of a first ring-shaped surface 22 with a second ring-shaped surface 24, which are shown in FIG. both elongate from opposite wall 30. The two ring-shaped surfaces 22 and 24 are tapped to accept the cylindrical photo-emitter 102 shown in Figures 2, 3 and 8.

From one to the other of the opposing walls extend means for receiving receiving means. In the presently preferred embodiment, the ring-shaped surface 26 is concentric with the ring-shaped surfaces 22 and 24 and protrudes from the opposite wall 32 to accept the cylindrical photodetector 104.

<img file="PT700268E_D0054.tif" />

With respect to cuvette 10, it is preferable that the ring-shaped surfaces 22 and 24 are concentric with each other and are concentric with the ring-shaped surface 26, and that the ring-shaped surface 22 has a smaller internal diameter. than the ring-shaped surface 24.

An alternative embodiment of the cuvette includes means responsive to the pressure impulses of fluid flowing through the conduit and attenuating variations of predetermined separation between two opposing walls. An example of this alternative embodiment is shown in Figures 9 and 10, wherein the flexing of a membrane or diaphragm 30a is attenuated by the air pocket 12b.

<img file="PT700268E_D0055.tif" />

Diaphragm 30a may be silicone having a preferred thickness of about 0.5 mm (0.020 inch), or PVC or PETG having a preferred thickness of about 0.127 mm (0.005 inch). Other silicone diaphragms of the appropriate thickness are also capable of similar performance and as such are considered equivalent. Diaphragm 30a has a circular periphery 30b which is fitted in a housing 12a and in the lower housing 8a, which are preferably made of medical grade plastics and sandwich the silicone diaphragm 30a therebetween.

Upper housing 12a has a convex surface 12c facing simultaneously to wall 32a and silicone membrane 30a. Convex surface 12c is a side of a third wall, wherein the silicone membrane 30a and wall 32a form the first and second walls.

Blood flows circularly in area 34a and semi - linearly in area 36a. As blood pulses, silicone membrane 30a bends from distance 60a in a

<img file="PT700268E_D0056.tif" />

62a, while wall 32a and surface 12c are relatively still. The silicone membrane 30a is an example of a transducer, which varies the thickness of the predetermined separation of the blood flow passage in the conduit with each pressure pulse. The presence of air pocket 12b will attenuate the flexion of the silicone membrane 30a as it is moved by the pressure of the fluid pulses.

<img file="PT700268E_D0057.tif" />

The airbag 12b is preferably hermetically sealed by forming an air impermeable connection between the lower compartment 8a, the upper compartment 12a and the circular periphery of the silicone membrane 30b.

<img file="PT700268E_D0058.tif" />

The airbag 12b and the diaphragm 30a in the upper compartment 12a, in an alternative embodiment, not shown in the Figures, may be supplemented by a second, symmetrically opposite second airbag (not shown) and a second flexible diaphragm (not shown). shown in the lower compartment 8a, so that the duct has two flexible diaphragms and two air pockets. The second air pocket would be formed in a wall depression 32a over which the second flexible diaphragm is arranged in a manner similar to that in Figure 10 the arrangement of the flexible diaphragm 30a and the air pocket 12b in the wall 12c is shown. . In such an alternative embodiment of the cuvette of the invention, blood flows in contact with and between the two flexible membranes. The pulsatile pressures of the flowing blood causes both flexible diaphragms to flex. The flexion in each of the flexible diaphragms is attenuated by the respective air pocket.

As another alternative embodiment of the cuvette of the invention, the conduit has only the flexible diaphragm 30a and a single air pocket 12b as shown in Figure 10 and also features a modification of the rigid wall 32a to have a plurality of small holes or microstrices, on its inner surface with respect to the conduit, so that the holes are in contact with the blood flowing through the conduit. The holes, being of relatively small diameter, are virtually impermeable to the blood flowing through the conduit, but nonetheless serve to attenuate the force and absorb some of the pulse pressure from the pulsatile blood flow.

<img file="PT700268E_D0059.tif" />

Pneumatic damping on microstrings decreases bending distance 62a and reduces both acceleration and velocity of flexible diaphragm 30a as it moves between its extreme positions at distance 60a and distance 62a.

In the two alternative and not shown above embodiments, the additional pneumatic damping in the cuvette of the invention serves to further increase the accuracy of spectrophotometry readings by reducing bounce motion and flexible diaphragm fluctuation during pulsatile blood cycles. . ,. . &gt; - ·

The ring-shaped surfaces 24a and 26a extend respectively from the third wall on surface 12c, and from second wall 32a. Similar to that shown in Figure 8, wavelengths 64a, 66a, 68a and 70a pass transmissively through membrane 30a from photo-emitter 102 to photodetector 104.

The air bag 12b is preferably hermetically sealed by forming an air tight bond between the lower compartment 8a, the upper compartment 12a and the circular periphery of the silicone membrane 30b.

The ring-shaped surfaces 24a and 26a extend respectively from the third wall on surface 12c and the second wall 32a. Similar to that shown in Figure 8, wavelengths 64a, 66a, 68a and 70a pass transmissively through the membranes 30a from photo-emitter 102 to photodetector 104.

Other means for receiving emission means and sensing means are contemplated and need not be ring-shaped surfaces. Such means may be a structure supporting the emitting means and the detecting means in close proximity to the cuvette means, so as to be especially suitable for the Incorporated Art.

In the inlet opening 16 of the cuvette 10 there is a threaded connection 80 and in the outlet opening 18 of the cuvette 10 there is a threaded connection 82. The threaded connections 80 and 82 are respectively connected to catheters 110 and 112. Catheters 110 and 112 are integral with the installation of the cuvette 10 in the inlet opening 230 and the outlet opening 240 of the dialyzer 130.

<img file="PT700268E_D0060.tif" />

Other equivalent embodiments of the cuvette are contemplated. However, from the mathematics described in the Incorporated Art, it is clear that other embodiments of the cuvette in an extracorporeal system require a conduit in which blood flows unimpeded and constantly. The duct should have its dimensions balanced on the thickness of the blood layer in the duct (such as the distance 60 in Figure 8), the thickness of the flexible membrane in the duct (such as the wall 30 in Figure 8), and the area of ​​the duct. flexible membrane (such as area 36 in Figure 6). Each of these parameters is empirically adjusted, as shown in Figure 12, so that electronics, as stated in the Embedded Technique, starve AC pulse signals in order to use the ΔΙ / I in EQUATIONS B and C, below.

<img file="PT700268E_D0061.tif" />

b) Bucket Structural Variables

Several physical characteristics of alternative embodiments of the conduct of the invention will be discussed below.

i) Transducers

The transducer means will modulate with each pump cycle of the pump means, such as pump 140 of Figure 1, to produce a small Ád (such as bending line 62 in Figure 8) like that of EQUATIONS A, B and C below are excited for the ebi1 ^ -vi notation (see Figure 1 by way of example). A cursory explanation of mathematics, more fully explained in Embedded Technique, is as follows:

The Beer-Lambert Equation is the theoretical basis of Embedded Technique, as shown by EQUATION A.

<img file="PT700268E_D0062.tif" />

<img file="PT700268E_D0063.tif" />

EQUATION A

I = I_Oand^{e, x, d + G}

Where:

G = Variable optical passage elongation factor

I = Measured Intensity

Io = Incident Intensity ε = Medium Extinction Coefficient x = Medium Concentration d = Medium Thickness

In the practical hemodynamic application of EQUATION A, as blood layer d (such as distance 60 in Figure 8) pulses due to pump means (such as pump 140 in Figure 1), the concentration x remains constant, but blood layer d (such as distance 60 in Figure 8) will change by Ad (such as bending line 62 in Figure 8) due to transducer means (such as wall 30 in Figure 8 ). Thus, taking the partial derivatives of Equation A over time and then dividing Equation A by itself will give Equation B.

EQUATION B

- dl / dt = eb · x · dd

It

Where:

ôd «Ad dT ar dT» ΔΙ / Ι T

Gb = whole blood extinction coefficient.

<img file="PT700268E_D0064.tif" />

In order to obtain the result in EQUATION B, G is assumed in EQUATION Ã to be negligible due to the fixed dimensions of the conduit.

<img file="PT700268E_D0065.tif" />

As a practical application of EQUATION B, suppose that two wavelengths, such as 805nm and 1300 nm, which are both wavelengths of equal specification, are now chosen, as was done in the Embedded Technique. The Embedded Technique asks for a ratio to be taken from EQUATION B at the first wavelength against EQUATION B at the second wavelength. Thus, the ratio of ΔΙ / I at wavelength 1 to the ratio ΔΙ / I at wavelength 2 results in the cancellation of both Ad ex, leaving only et, i in EQUATION C.

EQUATION C (ΔΙ / Ι) ι (ΔΙ / Ι)₂ ebi

Eb2

Where:

ebn⁼ Extinction coefficient in whole blood at the nth wavelength; en = Wavelength.

<img file="PT700268E_D0066.tif" />

The advantage of selecting these wavelengths as such is that at these wavelengths, the ratio of extinction coefficients, ebi1, vi, will be a strong function of the hematocrit in the pulsating blood flow through the cuvette. This strong function is graphically visible in Figure 11.

The Incorporated Technique is advantageous in that several mathematical terms of the above equations eliminate themselves as shown in EQUATIONS B and C, thereby dramatically simplifying the determination of the concentration of a biological constituent in the pulsatile fluid that flows. As noted, one principle component that is eliminated is concentration x, which divides itself in proportion. Additionally, the term I_O is canceled in EQUATION B where I_O is the second term for incident intensity.

(ii) Wings or Handles

<img file="PT700268E_D0067.tif" />

The wings or handles, such as wings 14 in Figures 3 to 6, are preferably thin to allow installation of the cuvette in areas where there is a tight fit to facilitate its connection to the blood tubing. The wings or handles also decrease the potential for fingerprinting on the flexible membrane, such as on wall 30 of Figure 8.

(iii) Bucket Flow Volume

Preferably the cuvette will maintain adequate blood flow through the flexible membrane while allowing for greater volume of flow through the outer annular area, an example of which is wall 30 with flow passage 40 relative to the wall. 46 and flow passage 42. A construction with this arrangement avoids turbulence within the conduit, slows down just below the flexible membrane sensor and reduces the flow rate sensitivity within the conduit.

<img file="PT700268E_D0068.tif" />

(iv) The Flexible Membrane

The rigidity of the flexible membrane depends on the thickness of the membrane and its area. Such rigidity minimizes the absolute motion of d, which is the thickness of blood flowing through the cuvette, such as the distance 60 in Figure 8 or 60a in Figure 10, and thus its variation with static pressure. Preferably the membrane will be sufficiently flexible to allow Ad, such as bending line 62 in Figure 8, with each AP, which is the change in pressure due to a pulsation in fluid flow. This ΔΡ facilitates the occurrence of the ΔΙ / Ι ratio of EQUATION C, and thus the elimination by mathematical cancellation of variable I_On EQUATION B.

<img file="PT700268E_D0069.tif" />

<img file="PT700268E_D0070.tif" />

For this reason, in order for the extracorporeal disposable cuvette of the invention to function effectively, it is important that there be a modification, Ad, of d, such as bending line 62 and distance line 60, respectively, in Figure 8, or such as bending line 62a and distance line 60, respectively, in Figure 10. Preferably, Ad will be a function of the absolute thickness d of the blood sample, the thickness of the membrane, the area or diameter of the membrane, the change in pressure due to pulses in the fluid and its peak-to-peak pressure change, or absolute pressure, and the total volume of the elastic modulus of the plastic from which the flexible membrane and cuvette are preferably constructed. These parameters allow the Ad of the flexible membrane within the conduit such that the ΔΙ / I in EQUATION C is operable, of which an example is shown in Figure 12. Figure 13 shows the error from the true value if Ad is improperly selected.

3 ELECTRONIC ASPECTS OF THE SYSTEM

The electronic components described in the Incorporated Art are essentially the same as would be used herein with respect to integrated circuits, photo-emitter and photodetectors. Digital schemes and similar electronic sensors would also be used herein in conjunction with the presently disclosed invention.

<img file="PT700268E_D0071.tif" />

4 DETERMINATION OF BLOOD VOLUME

In order to use the equations presented herein and in the Incorporated Art, determining the change in intravascular blood volume during hemodialysis using the hematocrit value requires that the following assumption be made. In the case of renal dialysis, as blood passes through the dialyzing membrane, it is assumed that no red blood cells pass through the dialyzing membrane. Only plasma and aqueous fluids, electrolytes and small molecules will pass through the dialyzer membrane. Accordingly, any change in hematocrit value due to dialysis can be shown to be inversely proportional to blood volume by the following EQUATIONS D to F.

<img file="PT700268E_D0072.tif" />

EQUATION D

PCV = RBC BV

EQUATION AND

PCVi

PCV₂

RBC1 / BV1

RBC₂ / bv₂

EQUATION F

Hcti

Hct₂

<img file="PT700268E_D0073.tif" />

BV₂

BV]

Where:

CV_no = Joint Cell Volume at time n,

RBC_no = Red Blood Cell Volume at time n,

BV_no = Volume of complete blood at time n,

Hct_no = Hematocrit value at time n.

Since the volume of red blood cells (RBC) is a constant during dialysis, since it is assumed that no red blood cells pass through the dialyzing membrane, EQUATION F obtains.

<img file="PT700268E_D0074.tif" />

By controlling the change in hematocrit value from moment n to moment n + 1, the change in blood volume over the same time period is seen. Specifically, by controlling the rate of change in a patient's blood volume over time while on dialysis, the rate of plasma refilling in a given patient can also be determined. This system allows precise control of the hematocrit value to determine and precisely control the percentage change in blood volume, which is directly related to the rate of plasma refill as a person undergoes dialysis. Knowing this parameter, the clinician is able to adjust, in real time, the rate of ultrafiltration, in order to neither sub - dialysis nor over - dialysis a patient.

<img file="PT700268E_D0075.tif" />

As the previous discussion related to noninvasive analysis of information (blood hematocrit), it will be appreciated that the aforementioned emitter / detector circuits are also adapted for in vitro analysis of this blood hematocrit value.

The principles within the scope of the present invention require: (1) optimization of the fixed, spatial and geometric parameters in the cuvette and (2) at the same time allowing flexible movement (Ád) of a membrane, and (3) allowing cálculo / Τ calculation of EQUATION C for rapid and accurate determination of hematocrit and blood volume modification. These principles are found within the scope of preferred design.

The cuvette duct 10 in the blood tubing system described herein is an optical passageway through which blood flows through the duct, which is of the preferred design seen in Figures 1 to 8. The duct allows light to pass from the photo emitter 102 photodetector 104 through the blood sample passing through the cuvette 10. The cuvette 10 is designed with a wall 30 which is a flexible membrane that moves with impulses due to pressure changes in the blood line due to pump 140 which pulsatibly pumps blood through the line. of extracorporeal tubing.

Important elements of the cuvette are blood thickness, such as distance 60 in Figure 8, and the change in blood thickness, as shown by bending line 62 in Figure 8. This modification is due to the pulsations created by pump 140. , first and second opposing walls 30 and 32 and their areas, and finally the plastic volume modulus of the first wall 30, which all combine to allow the first wall 30 to float with each pulse of fluid.

Through the cuvette 10 an emitter / detector apparatus 100 is placed. The light is then irradiated by the photo-emitter 102 through the blood sample present in the conduit and is measured on the photodetector 104, as shown in Figure 8. Spectrophotometric techniques are then performed. by computer 150 for later real-time presentation, or before and after dialysis, of constituent concentration values. From such values, if the concentration value sought is that of red blood cells, the determination of instantaneous change in blood volume can be obtained in real time. Thus, hematocrit values ​​are obtained noninvasively through the use of electromagnetic radiation as an information carrier. The present invention may be used in a disposable extracorporeal duct system to determine the critical intravascular blood volume parameter.

<img file="PT700268E_D0076.tif" />

It will also be appreciated that the present invention will also provide a system and apparatus which can provide the subject with immediate and continuous blood volume control information. It also provides continuous, noninvasive information regarding the oxygen saturation status of the patient's blood, regardless of the patient's hematocrit value. Even under conditions of low blood perfusion in a patient, extracorporeal control described herein provides increased accuracy and ease of use.

<img file="PT700268E_D0077.tif" />

The embodiments described should be considered in all respects as illustrative only and not restrictive. The scope of the invention is therefore indicated by the appended claims rather than the foregoing description. All modifications falling within the meaning and equivalence range of the claims should be encompassed within their scope.