Disposable extracorporeal conduit for blood constituent monitoring

IS 2 165 877 T3

DESCRIPTION

Disposable extracorporeal conduit to control blood constituents.

Field of the invention

The present invention relates to an apparatus and a system for obtaining a desired concentration value of the biological constituent present in a pulsating flowing fluid and is more particularly aimed at obtaining these results in pulsatingly flowing blood in a hemodialysis environment.

State of the art

The kidneys are located on both sides of the spine. In a healthy patient, the kidneys function to stimulate the production of red blood cells and regulate the content of the blood. The kidneys also produce hormones that 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 task of filtering each kidney is performed, in part, by approximately one million nephrons in the kidney. Nephrons are filtering units made up of tiny blood vessels. Each of these blood vessels is called a glomerulus. Each day the kidney treats approximately 0.22 m³(200 quarts) of blood and fluid. The kidney withdraws approximately 0.0022 m³ (two quarts) of water and toxic chemicals that are sent to the bladder in the form of urine for subsequent evacuation when urinating.

A patient whose kidneys are functioning below normal may be subjected to dialysis as a substitute for the normal blood cleansing function normally performed by the normally functioning kidneys. Dialysis is a procedure in which the function of the kidneys to clean the blood is performed in a substitute way. The diaolysis procedure for routine use was perfected in the 1960s, having been invented about fifty years ago. For the purposes of describing and illustrating hemodialysis, reference is now made to Figure 1. Although Figure 1 incorporates a view of a presently preferred embodiment of the present invention, it also incorporates a view of some common components that are normally found in a general hemodialysis environment. The general setting of hemodialysis and the usual components in bone are described below.

In hemodialysis, blood is taken from patient 200 via a collection catheter, an example of which is shown in FIG. 1 as inlet catheter 122. Inlet catheter 122 is inserted intravenously into patient 200 at zone 180 and is used to define an ascending blood path to a blood filter used to filter impurities from the blood. The blood filter is also called a dialyzer 130. Dirty blood flows from an artery of patient 200 to a pump means, an example of this is pump 140. From pump 140, blood flows to dialyzer 130. Dialyzer 130 has inlet port 230 and port outlet 240. Pump 140 performs the function of causing dirty blood from patient 200 to enter through inlet port 230, pass through dialyzer 130, and exit dialyzer 130 through outlet port 240.

Specifically, the dirty blood from the inlet catheter 122 is transported to the inlet port 230 of the dialyzer 130. After passing through the dialyzer 130 and having been cleaned by it, the blood may receive further treatment, such as a heparin drip , in the corresponding component 300 of hemodialysis. The now clean blood returns to the patient 200 after the dialysis treatment through an outlet catheter means, an example of which is the outlet catheter 124. The outlet catheter 124, which is likewise inserted intravenously into the patient 200 at Zone 180 defines a descending blood path from dialyzer 130, with dialyzer 130 returning the outgoing blood to patient 200.

As noted, the hemodialysis procedure uses a blood filter or dialyzer 130 to clean the blood from patient 200. As the blood passes through dialyzer 130, it travels through straw-like tubes (not shown) in the interior of dialyzer 130 serving as membrane passages for dirty blood. The straw-like tubes remove toxins and excess fluid through a diffusion procedure. An example of excess fluid in dirty blood is water and examples of toxins in dirty blood are blood urea nitrogen (BUN) and potassium.

Excess fluid and toxins are removed by means of a clean liquid dialysate fluid, which is a solution of chemicals and water. Clean dialysate enters dialyzer 130 through inlet tube 210 from a combined controller and reservoir 170. The dialysate surrounds the

ES 2 165 877 T3 straw-like tubes in dialyzer 130 as dialysate flows down through dialyzer 130. The clean dialysate collects excess fluid and toxins that pass through the straw-like tubes, by diffusion, and then returns excess fluid and toxins along with dialysate exiting dialyzer 130 through outlet tube 220, thereby cleaning the blood. The dialysate exiting the outlet tube 220 can be discarded after cleaning the blood.

In summary, dirty blood flows from an artery of patient 200 to pump 140 and then to dialyzer 130. The dirty blood flows into dialyzer 130 from inlet catheter 122 and clean blood flows out of dialyzer 130 through the exit catheter 124 returning to patient 200.

Hemodialysis, which removes excess fluid from a patient's blood, has a strong impact on the body's fluid balance due in part to the rapid change in the volume of circulating blood. When the rate of fluid withdrawal is faster than that of plasma replenishment from the body, intravascular blood volume decreases. This resulting fluid imbalance has been linked to complications such as hypotension, loss of consciousness, headaches, vomiting, vertigo, and cramps experienced by the patient, both during and after dialysis treatments. As hypotension and true shocks occur in up to 25% of hemodialysis treatments, hypovolemia induced by dialysis remains a major complication of hemodialysis.

Many dialysis patients already have compromised circulatory responses due to the side effects of end-stage kidney disease. A malfunction of the compensatory mechanisms of blood pressure due to the decrease in intravascular volume has been considered one of the main factors causing the hypotension induced by dialysis.

In order to reduce the risk of dialysis-induced hypotension, continuous measurement of circulating blood volume can optimize dialysis therapy regimens, control fluid balance, and help achieve the patient's dry weight goal. on a quantitative basis. Volumetric controllers, while providing an accurate measurement of the amount of fluid removed through ultrafiltration, do not provide any indication of how the body's plasma replenishment mechanisms respond to actual fluid removal. Factors such as food eaten, drinking water, and changes in posture also significantly affect the volume of blood circulating during dialysis. Acts such as eating, drinking, and posture illustrate how sensitive plasma replenishment mechanisms are.

The hematocrit value provides an indication of the change in blood volume. Since the number of red blood cells in whole blood is not significantly altered by dialysis, and the mean corpuscular volume of red blood cells remains essentially constant, it turns out that changes in blood volume will be inversely proportional to changes in blood volume. hematocrit. Therefore, the change in the patient's blood volume can be defined at any time in the course of dialysis treatment as in Equation 1.

Equation 1 ^BVfinal ^Hctinitial ^BV initial ^Hctfinal

Where:

BVf_inai _ Final blood volume

Initial BV _ Initial Blood Volume

HCTf_final _ Final hematocrit value

In the clinical setting, however, it may be more useful to determine the percentage change in blood volume represented in Equation 2.

Equation 2

100 % x

BV_F - BV_iBV r Hct_i

ÍHctf x 100%

Where:

BVf _ Final blood volume

IS 2 165 877 T3

BVj = Initial blood volume

HCTj = Initial hematocrit value

HCTf = Final hematocrit value

It is known that the change in hematocrit is used as a measure of the change in the actual volume of the blood during dialysis. However, in order for the relationship between the change in hematocrit and the change in blood volume to be useful, the hematocrit must be monitored precisely, continuously and in real time throughout the entire hemodialysis treatment session. Although precision can be achieved through elaborate technical means, to be clinically practical, real-time control of hematocrit and blood volume should be easy to use, save time for nurses, operate non-invasively, and be justified in terms of cost.

Various techniques used to control intravascular blood volume change due to ultrafiltration as a function of the hematocrit value include microcentrifugation, electrical conductivity, and photometry.

In microcentrifugation, a microcentrifuge machine is used to measure the hematocrit. This procedure is unsuitable for monitoring blood volume changes in real time, due to the time between measurements, the high probability of sampling and reading errors, and the need to adequately compensate for trapped plasma in the columns of Red blood cells. Hence, due to the intensely labor-intensive nature of centrifuging patient blood samples on a timely basis, this technique is totally inadequate, impractical, and excessively costly for large-scale clinical application.

In an attempt to obtain real-time hematocrit information, electrical conductivity measurements have been used. Conductometric measurements, however, are adversely affected by the abnormal concentrations of electrolytes, anticoagulants, and proteones, which are present, all of which are present in dialysis patients. Hence, this particular technique is equally plagued with significant technical errors.

Optical techniques, although generally unaffected by such problems, have been susceptible to other instabilities. These include variations in ambient light, intubation devices, changes in the rhythm of blood flow, lone pressures, and oxygen saturation. Furthermore, light sources used in optical techniques require frequent calibration.

Brief summary of the invention

According to the present invention, as explained in the embodiment widely described herein, the invention was intended for an apparatus and a system incorporating the apparatus. The device is a disposable extracorporeal cuvette, through which a fluid passes in a pulsating manner.

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

The conduit has two opposite walls that have a predetermined spacing from each other. The two opposite walls retain at least part of the pulsating fluid that flows through the conduit. The conduit also has a transducer means that was located on one of the opposite walls. The transducer means varied the predetermined spacing between the two opposing walls in response to pressure pulsations in the flowing pulsating fluid.

The conduit may alternatively be said to have a means, in contact with the pulsating fluid flowing within it, to create an elastic fluctuation with each pulsation of the pulsating fluid flowing in the conduit. Elastic fluctuation occurs in a direction that is essentially normal or perpendicular to the general direction of movement of the pulsating fluid that was flowing in the conduit. The elastic nature of the fluctuation medium ensures the return to the original position after the fluctuation is performed.

The conduit, also called the second fluid conduit, was made of materials that allow at least two predetermined wavelengths of electromagnetic radiation and, preferably, four wavelengths, to pass transmissively through west. The materials

ES 2 165 877 T3 used are preferably low cost so that the cuvette is economical both to dispose of after use and to manufacture.

The term "fluid" is intended herein to mean alternatively liquid or gaseous substances. "Pulsating" is also intended herein to mean rhythmic or cyclical impulses or increases in pressure in a fluid flowing under pressure.

The system that incorporates the cuvette is designed to control the concentration of a particular or desired biological constituent. Preferably, the pulsatingly flowing fluid is blood from a patient and the desired biological constituent concentration that is controlled in the system is the red blood cell concentration, also expressed as the hematocrit value. The preferred system in which the control of the concentration of constituents takes place is a hemodialysis system in which the hematocrit value is controlled both before and after the blood cleaning procedure, as a means of knowing the change in volume of blood. blood during the hemodiaolysis procedure.

Constituent concentration control calculations are performed using the technique described in the PCT patent application, serial No.^or PCT / US93 / 03427 filed April 12, 1993, entitled "System and Method for Noninvasive Hematocrit Control," and is hereinafter referred to as the Incorporated Art.

The Incorporated Technique, known as non-invasive differential ratiomometric spectrophotometry, is described below. Incident radiation passing over or through living tissue is assumed to pass through a combination of blood, tissue, and interstitial fluid compartments. The light attenuated by said living tissue can be expressed by the modified Beer-Lambert equation:

I i - (fibfXa + Xv) - £ tXt + eiXi) d + G (2)

Equation (2) can also be written:

ln (I / Io) = - © (Xa + Xv) + etXt + ε Xi) d + G (2a)

Where ε _b, ε _t, and ε j represent the extinction coefficient in the blood, tissue and interstitial fluid compartments, respectively; X_or and X_v represent the concentration of arterial and venous blood (X_b= X_or + X_v), X_t represents the concentration of tissue absorbers, and X¿ represents the relative concentration of water and dissolved components in the interstitial fluid compartment; d represents the intrasensor space; and G is a constant of the geometrical configuration.

When the layer of blood pulses, the terms of concentration change. The term d can be set by the geometric configuration of the device. Taking the partial derivatives of equation (2) with respect to time and dividing by equation (2) gives:

^{dI / dt} = © (5xa / 5t + dxv / dt) + £_tdX_t/ dt + εί 0¾ / dt) d + dG / dt (3) that can be simplified in each compartment and wavelength by assigning X = X / t, and G = G / t, and so that it gives:

V / = - [φ] ​​λ

V / = © (X'a + XV) + £ t Xt + εί X <) d + G '(4)

Assuming X_t and G do not vary significantly during the pulse time interval, G = 0 and X can be simplified _t = 0, and equation (4) is then:

V'X = © (Xa + XV) + εί Xí) d (5)

Examining the transport between X_or and X_v, we can form a proportional constant K such as X_v = K_v V_or, which represents the reactionary nature of the venous component and additionally reduces the equation that precedes

V / = © (1 - Kv) Xa + είX ') d (6)

IS 2 165 877 T3

Since X., and X. are not wavelength dependent (λ) the V \ values ​​at different wavelengths can be differentially subtracted to produce an independent hematocrit term containing only £ jXj information. Although the term V₈₀₅/ V 1310 provides useful data on relative changes in hematocrit, it should be recognized that the simple ratio V _s05/ V ₁₃₁₀ it is not accurate enough to determine the hematocrit value unless the term -; X ... · is known or eliminated. For example, the term -jX may not be taken into account_i80S since ε.₈₀₅ is actually very small, while the term ejX_i1310 is approximately 25-50% of the ε value _b1310 of one's own blood and, therefore, cannot be disregarded without affecting precision.

Figures 15 and 19 suggest that a linear combination of V \ with λ = 805 nm and λ = 970 nm would have a nearly constant value for a range of Hct values. As the extinction coefficients are well known -_is0sand -j9₇₀, or they can be determined empirically, it can be found that a constant of proportionality R₁ produce:

-¡970X1 = V970 - R1V805 (7)

This correction term can now be applied with a second constant of proportionality R₂(where R₂ is approximately equal to -_i1310 /-_i970) of the term V ₁₃₁₀ to precisely eliminate your sensitivity -_i1310Xj, hence:

-b1310 (1 - Kv) X '_to = Vj310 - R2 (V970 - R ^ s) (8)

This corrected term can now be used ratiometrically with V _80s to remove the (1-K_v) X, and leave the proportion of the pure extinction coefficient represented below by equation (9) and shown graphically in figure 16.

^ε =_________^V© ( ^-b1310 ^Ví310 ^{- R}2^(V970 ^{- R}1^V805⁾

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

A. Although the wavelengths λ = 805 nm and λ = 1310 nm are nearly isobestic, the actual function of ε compared to Hematocrit at each given wavelength must have hematocrit data that differ in curvature, or deviation , or in linearity, or in sign with each other. See figure 15. If the functions ε_λ compared to the hematocrit are not different enough, then the ratio ε _by1/ ε ^_λ2 it will have no hematocrit data. See Figures 20A and 20B and Figures 21A and 21B. Although the above discussion refers to the isobestic wavelengths of λ = 805 nm and λ = 1310 nm, it was seen that other isobestic wavelengths such as λ = 570 nm, λ = 589 nm and λ = 1550 nm.

B. Also, the wavelengths should be selected close enough to each other so that the optical path lengths, d, are approximately the same. Longer wavelengths are preferred as they have less sensitivity to scattering, s:

^S «= Λ ⁽¹⁰⁾

C. The geometric or spatial relationship of the emitters and sensors is important. For example, if vertically aligned emitters are used in an earlobe measuring device, it is possible that the uppermost emitter will illuminate a different amount of blood-filled tissue than the lowermost emitter. If only one sensor is used, then there will be a disparity between X_b at each wavelength. Likewise, the sensor-emitter spatial separation distance is very important since the pressure applied to the tissue between the sensor and the emitters affects the resistance of the capillary and arteriolar vessels. This changes the X when the pressure (or distance) changes. This change in X then regulates the function V \. Therefore, the sensor-emitter separation distance must be such that the pressure applied to the earlobe, the fingertip or other body member, does not affect the function of V \. This sensor separation distance is empirically determined and should generate not less than 40 mm Hg of applied transmural pressure.

A horizontal alignment of the emitters with respect to the single sensor can be arranged so that the emitters and sensors illuminate and detect suitable areas of X_y1 and from X, \₂. It is important

ES 2 165 877 T3 note that the term d, the sensor-emitter separation, will be different between λ_χ and λ₂ by the cosine of the angle between the sensor and the emitter. Therefore, if there is any misalignment with respect to normal, the term d will not be canceled to obtain equation (9).

The preferred arrangement is one in which all the emitters (660, 805, 950 and 1310 nm) are located on the same substrate. It is preferred because the emitters will then illuminate essentially the same area X_b.

D. In the case of reflectance spectrophotometry, an aperture is required for the sensor and for each emitter. A sensor-emitter separation is also required so that the reflectance of the first layer of tissue, R_t, (a bloodless layer of the epithelium) does not further increase a multiple scattering effect, that is, the total reflectance, R, measured would also contain spurious information on the reflectance of the epithelial layers, where:

R - Rt +

T_t².Rb (1 - Rb.Rt) (11) where R is the total reflectance, R_t is the reflectance due to the first epithelial layer of the tissue, R_b is the reflectance due to the blood layer and T_t, is the transmission through the first layer of tissue.

The reflectance equations that describe R_t or R_b must now add up all the backscattered light that the sensor detects, that is:

Rb -

<img file="ES2165877T3_D0001.tif" />

(source function) (spread function) (12)

Although equation (9) describes the theory of the non-invasive hematocrit device, the four estimates (AD) are important for repeatability and accurate operation of the hematocrit device.

Assuming elements A to D are properly treated, then (9) becomes:

^gbA1 _—^(s1 + ^ki⁾(13) £ bA2 (s2 + k₂) where s is a scattering constant and k is an absorption constant, and where in whole blood:

s - o-_sHct (1 - Hct) (14) k - &lt;r_toHct (at isobestic wavelengths) (15) where σ_s is the dispersion cross section and σ_α is the absorption cross section.

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

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

εχχ _— σ₈ι (1 - Hct) + σαΐ () ελ2 σ_δtwo (1 - Hct) +

Therefore, the proportion ε_λ1/ ε_λ2 it is a function of the hematocrit. From Figure 16, a look-up table or polynomial curve adaptation equation can be obtained and used in the final hematocrit results displayed. Knowing the real hematocrit value, it is easy to see (Figure 14) that a wavelength of 660 nanometers can be selected to obtain a ratio ε from which the oxygen saturation value independent of the hematocrit is obtained. For example, equation (16) would become:

ES 2 165 877 T3 £ b660 _ ^σ~ 66θ (1 - Hct) +&gt;, 660 + ^Sa ° 2⁽&gt; i0660 ^— &lt;^ ar660⁾ £ b805 ^σδ805^{(1 - Hct)} + ^σα805 + ^Sto^OR2⁽^ ao805 ^— &lt;^ as805⁾ (17)

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

Figure 18 graphically demonstrates the need for a hematocrit independent blood saturation device. When the hematocrit value or the percent oxygen saturation decreases, the percent saturation error becomes unacceptable for clonal use . For example, it is not uncommon to see patients who have a low hematocrit (about 20%) who also have breathing difficulties (low oxygen saturation). Hence clinicians simply need more accurate oxygen saturation values.

Knowing the oxygen saturation values ​​and the hematocrit, the computation of the oxygen content is simple and can be directly visualized (a value that the doctor did not previously have as a non-invasive result, in real and continuous time):

[Oxygen content] = Hct.S_toOR₂.K (18) where K is an emporically determined constant.

With reference to equations (16) and (9) the computer must make a decision on the convenience of using the Taylor expansion approximation on the logarithm. This algorithm is kept in the software as a decisive decision for the reading and sampling algorithms. The Taylor approximation is only valid for small I / t values.

The aforementioned techniques describe conditions and equations in which the wavelengths are chosen so that the hematocrit value obtained does not have any interference from oxygen saturation, and therefore an independently determined hematocrit value.

However, you can also choose λ₂ (the reference wavelength) from equation (13) of 1550 nm. In the radiation zone from 900 to 2000 nm, the absorption coefficients of blood depend on the hematocrit and water, while at 805 nm the absorption coefficient of blood depends only on the hematocrit. Therefore, using a combination, the wavelengths 660, 805, and 1550, will also give a technique to determine the hematocrit (ε₈₀₅/ ει₅₅₀) and oxygen saturation (ε₆₆₀/ ε₈₀₅).

The present invention can 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 spectrophotometric techniques can be used.

Constituent concentration calculations, described in the Incorporated Technique, are performed with the data obtained by passing multiple wavelengths of light through the extracorporeal blood conduit or a part of the patient's body. Means with Incorporated Technique for emitting and detecting said multiple wavelengths of light and for analyzing the various detected light intensities are also described and provided. The spatial arrangement, both for detecting and emitting light, is detailed in such a way as to provide the optimum repeatability of the signals and data obtained from them. Finally, memory and calculation means are included that are capable of storing, manipulating, displaying, and printing the detected signals in a variety of ways. The Embedded Technique allows an end user, nurse, physician or patient to obtain a desired biological constituent concentration value, such as the hematocrit value, the oxygen saturation value or the oxygen content value, by representing them as values. digital real-time.

Although the apparatus of the present invention is preferably applicable in the field of kidney dialysis, it can also be used in cardiovascular surgery or in other medical fields where blood is present in extracorporeal tubes. In these fields, the present invention obtains the hematocrit value of the blood, the oxygen saturation of the blood, the oxygen content of the blood and the volume change of the blood. These results are all obtained through the present invention, without the need to obtain a stagnant blood sample with invasive methods. Advantageously, obtaining the hematocrit value in these cases provides repeatable and reliable determinations, in a non-invasive and continuous manner, of the hematocrit value of a patient regardless of the perfusion status or the cardiac output volume of the patient.

IS 2 165 877 T3

Brief description of the drawings

In order to better understand how to achieve the aforementioned and other advantages, a more particular description of the invention will be set forth with reference to the specific embodiments thereof which are illustrated in the attached drawings. It should be understood that said drawings represent only usual embodiments of the invention and that, therefore, they should not be considered as limiting the scope of the invention, presenting the invention in its embodiment that is currently considered the best to put into practice and use said invention being described with additional specificity and in detail through the attached drawings in which:

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

Figure 2 is an enlarged section, perspective view, of a cuvette shown in Figure 1.

Figure 3 is a perspective view of Figure 2 with the various components thereof disconnected.

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

Figure 5 is a cross-sectional view taken along the line at section 5-5 of Figure 3.

Figure 6 is a cross-sectional view, taken along the line of section 6-6 of Figure 3, showing the fluid flow paths within the cuvette.

Figure 7 is a cross-sectional view, taken along section line 7-7 of Figure 3.

Figure 8 is a cross-sectional view taken along section line 8-8 of Figure 2, depicting the radiation of four (4) wavelengths directed 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 a cross-sectional view of the alternative embodiment of the cuvette, taken along section line 8-8 of Figure 2, depicting radiation at four (4) wavelengths directed toward a Flexible silicone membrane inside the cuvette.

Figure 11 is the graph of the proportion e_bi / e _b2 compared to the hematocrit, where e_bi is the extinction coefficient of whole blood to radiation from a first wavelength of radiation and where e_b2 the extinction coefficient of whole blood at radiation of a second wavelength of radiation, in which the first and second wavelengths are isobastic wavelengths such as that Figure 11 represents an important function of the hematocrit.

Figure 12 represents the relationships of phasic variables between the dimensions of the component of the cuvette represented in figure 8, which includes the thickness of the blood layer and its changes (d and Ad, respectively), the thickness of the membrane disposable (T) and the membrane area (A).

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

Figure 14 is a graph representing the coefficients of oaptic absorption of oxyhemoglobin (HbO₂), reduced hemoglobin (Hb), and water (H₂O) compared to wavelength.

Figure 15 is a graph representing the relationship between the extinction coefficient of light at three different wavelengths compared to the hematocrit of whole blood.

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

IS 2 165 877 T3

Figures 17A-17E provide a flow chart showing the phases performed during a currently preferred method of the Incorporated Technique, using the pulsating component of the subject's blood flow to provide accurate blood oxygen saturation and hematocrit values. .

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

Figure 19 is a graph of e_bso5 / e_b97o compared to hematocrit.

Figures 20A-20B are graphs of ε versus hematocrit at two non-preferred wavelengths and ε₄/ ε₂ compared to hematocrit at such non-preferred wavelengths.

Figures 21A-21B are graphs of ε compared to hematocrit at two non-preferred wavelengths, and ε₄/ ε ₂ compared to hematocrit at such 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 that is present in the blood of a patient undergoing hemodialysis treatment.

1. The system

The general hemodialysis procedure and environment are represented in Figure 1 and have been described above. A summary of such a procedure is that patient 200, whose lungs are functioning below normal, is subjected to dialysis. Dirty blood flows from an artery of patient 200 to 140 and then to 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.

It is preferable that pump 140 flows blood into, through, and out of dialyzer 130 so that it flows in a pulsating fashion.

Installed at either end of the dialyzer 130 is a spectrophotometric means to define a blood flow path, to emit radiation into the blood in the flow path, and to detect radiation passing through the blood and of the flow path. The spectrophotometric means includes a cuvette means for defining the path of the blood flow, and an emitting / detecting means for directing and detecting radiation. Inside the emitting / detecting means there are an emission means to direct the radiation and a detecting means to detect the radiation.

In the preferred embodiment, as shown in Figures 3 and 8, an example of the emitter / detector means is represented by the emitter / detector apparatus 100. An example of the emission means is indicated by a photoemitter 102. The emitter / detector apparatus 100 also has detection means, an example of which is represented as a photodetector 104. An example of the cuvette means is represented in Figures 3 and 8 with the cuvette 10.

The emitter / detector apparatus 100 allows the detection by the photodetector 104 of the portion of radiation that is directed by the photoemitter 102 to the cuvette 10 and which passes through the blood found in it and in the cuvette 10.

As shown in Figures 2 and 8, a cuvette 10 is installed at either end of the dialyzer 130. Each cuvette 10 has a light emitter 102 and a photodetector 104 thereon. In the preferred embodiment of the system, a photoemitter 102 and a photodetector 104 are represented as being held together by a spring-loaded C-shaped clip on the photoemitter / detector apparatus 100.

The emitting / detecting means has an electrical connection with a computing means. In a preferred embodiment of the system, an example of the calculation means is represented in Figure 1 as the computer 150, which is electrically connected to a photoemitter 102 and to the photodetector 104 in the emitter / detector apparatus 100 by means of the cable. 120.

The inlet catheter 122 draws blood to the cuvette 10 located before the inlet port 230 of the dialyzer 130. The emitter / detector apparatus 100, at the inlet port 230 of the dialyzer 130, subjects

ES 2 165 877 T3 blood at least two wavelengths of electromagnetic radiation for the purposes of analysis, by means of spectrophotometry, so that the concentration of a desired biological constituent can be obtained. Each photodetector 104, at the inlet 230 and outlet 240 ports of dialyzer 130, communicates detected radiation at at least one first and one second wavelengths, via cable 120, to computer 150.

The computer 150 calculates, before and after the diaolysis, the concentrations of biological constituent that are intended or desired to be found. Next, the computer 150 respectively displays, in a first representation 152 and in a second representation 154, the obtained concentration of the biological constituent, in analogue or digital representations. The calculation means, represented here as for example the computer 150, preferably has the multiple capacity of simultaneous computation in real time and of representation of the hematocrit and oxygen saturation values ​​as well as the change of the volume percentage. of blood from a patient undergoing hemodialysis.

The selection and predetermination of the radiation wavelengths are based on the desired biological constituent whose concentration value is sought. Photoemitter 102 preferably emits and photodetector 104 preferably detects four (4) predetermined wavelengths for spectrophotometric techniques described in the Incorporated Art. Accordingly, cuvette 10 would preferably comprise materials that allow the four (4) predetermined wavelengths of directed radiation to pass through them.

two. The apparatus

a) A preferred embodiment

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

As Figure 4 shows, there is an upper frame 12 that was assembled inside a lower frame 8 to form a trough 10. The upper frame 12 may be installed on the lower frame 8 by means of an adhesive. Other equivalent means such as rub welding or ultrasound welding can also be used. The purpose of properly adhering the upper frame 12 to the lower frame 8 is to create between them a hermetic and fluid-impermeable joint, so that the fluids that are passed through the cuvette 10 do not leak or ooze at the points of contact. joined between the upper frame 12 and the lower frame 8. The lower frame 8 has handles or fins 14 with which the bowl 10 can be operated manually.

The conduit incorporates a transducer medium. As indicated, the transducer means varied the predetermined spacing between the two opposing walls with each pressure pulsation in the fluid. In the presently preferred embodiment, an example of a transducer means is depicted in Figures 4, 5, 7 and 8 as wall 30 having an opposite wall 32 thereto.

The fluid flows pulsatingly in the conduit, within the area limited between a vertical wall 46 and the opposite walls 30 and 32. The fluid, within the conduit, follows the flow paths indicated by arrows 40 and 42 in the figure. 6. The flow path seen in Figure 8, at 44, which describes the flow of the fluid as it enters the inlet 16 in the conduit, adopted the flow path 40 inside the area 36 or, alternatively, the path of flow 42 into area 34. Fluid flowing between opposing walls 30 and 32 is described by area marked 36. Figure 12 represents area 36 as "A". Fluid flowing out of opposing 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 the area 34 is preferably greater than that of the area. 36.

Inlet 16 and outlet 18 are linearly aligned to either side of the conduit and share a common longitudinal axis passing between them. The cylindrical conduit between inlet 16 and outlet 18 has a longitudinal axis passing through opposing walls 30, 32 which is the same as the common longitudinal axis of inlet 16 and outlet 18. Figure 12 depicts the wall. 30 as the membrane thickness "T". As shown in Figure 8, the opposite wall 30 is preferably thinner than the opposite wall 32.

When the fluid in flow path 44 causes a pressure pulsation to arise, the opposite wall

ES 2 165 877 T3 flexes by arching while the opposite wall 32 remains relatively immobile. The deformation of the wall 30 is represented by the flexian line 62 in Figure 8. After the pressure pulsation of the fluid in the flow path 44, the wall 30 returns to the position represented in a broken line, so that the The separation between wall 30 and 32 is indicated by distance 60. The relationship between the distance 60 and the flexion line 62 is represented graphically in FIG. 12, respectively by d and Ad.

As mentioned, wall 30 is an example of the transducer means. Other transducer means operating in an equivalent manner could be incorporated into the conduit, such that the distance represented by line 60 in FIG. 8 can vary between opposite walls 30 and 32. By way of example, and without limitation, the means The transducer could be made essentially of silicone. Alternatively, it could be made so that a small section of the opposing wall 30 varies the distance between the opposing walls 30, 32. Said portion may be provided with a spring or have elastic means to return the small wall section 30 to its original position of lower pressure, prior to the pulsation. In such an embodiment, the opposite wall 30 need not be thinner than the opposite wall 32.

Figure 8 represents the oáptic paths 64, 66, 68 and 70 of electromagnetic radiation directed at four (4) different wavelengths. Each wavelength is selected to have spectrophotometric compatibility with a specific biological constituent in the pulsating fluid. As described, the portion of radiation detected and directed is used to obtain the concentration value using the Incorporated Technique. The wavelengths can be set in the emitter / detector apparatus 100 or they can be set by adjusting the computer 150, when the photoemitter 102 and the photodetector 104 are dynamically adjustable in terms of wavelength in the computer 150.

The opposite wall in the conduit incorporating the transducer means also has a means for receiving an emission medium extending from the wall. The other opposite wall has a means for receiving a sensing means. In the presently preferred embodiment, the means for receiving a photoemitter is indicated in Figures 3, 4, 5 and 7, as the combination of a first ring-shaped surface 22 with a second ring-shaped surface 24 projecting both on the opposite wall 30. The two ring-shaped surfaces 22 and 24 are designed to accept a cylindrical light emitter 102, shown in Figures 2, 3 and 8.

One medium extends from the other opposite wall to receive a sensing medium. In the presently preferred embodiment, the ring-shaped surface 26 is concaentric to the ring-shaped surfaces 22 and 24 and extends from the opposite wall 32 so as to accept the cylindrical photodetector 104.

As regards the bowl 10, it is preferable that the ring-shaped surfaces 22 and 24 are concontric with each other and concontric with the ring-shaped surface 26 and that the ring-shaped surface 22 has a smaller internal diameter. to the ring-shaped surface 24.

An alternative embodiment of the cuvette includes means that respond to pulsations of pressure in the fluid flowing through the conduit and that dampen variations in the predetermined spacing between two opposite walls. An example of this alternative embodiment is represented in Figures 9 and 10 in which the flexing of a membrane or diaphragm 30a is dampened by an air pocket 12b.

The diaphragm 30a may comprise silicone having a preferred thickness of 0.020 inches, or PVC or PETG having a preferred thickness of 0.005 inches. Other silicone-like diaphragms of suitable thickness are equally capable of equivalent performance and are therefore considered equivalent. The diaphragm 30a has a circular periphery 30b that is inserted in a frame 12a and a lower frame 8a, preferably made of plastic for medical purposes, which sand the silicone diaphragm 30a between them.

The upper frame 12a has a convex surface 12c facing wall 32a and a silicone membrane 30a. The convex surface 12c is a side of a third wall, in which the silicone membrane 30a and the wall 32a form the first and second walls.

The blood flows in the corpuscle in area 34a and semilinear in area 36a. When the blood pulses, the silicone membrane 30a bends from distance 60a with an increment of 62a, while the wall 32a and surface 12c remain relatively immobile. Silicone membrane 30a is

ES 2 165 877 T3 is an example of a transducer that varies the predetermined separation thickness of the blood flow path in the conduit with each pressure pulse. The presence of the air bag 12b will cushion the flexing of the silicone membrane 30a. The damping effect reduces sudden movements of the membrane 30a when it is moved by pressure pulsations in the fluid.

Air bag 12b is preferably hermetically sealed by forming an airtight connection between lower shell 8a, upper shell 12a, and the circular periphery of silicone membrane 30b.

The air bag 12b and the diaphragm 30a in the upper frame 12a, in an alternative embodiment not shown in the figures, may be complemented by a second air bag (not shown) positioned opposite and symmetrical and a second flexible diaphragm ( not shown) in the lower frame 8a so that the duct has two flexible diaphragms and two airbags. The second air bag should be formed in a recessed portion of the wall 32a on which the second flexible diaphragm rests in a manner similar to that shown in Figure 10 of the arrangement of the flexible diaphragm 30a and of the air bag 12b in the wall 12c. In said alternative embodiment of the cuvette of the invention, the blood flows in contact with the two flexible membranes and between them. The pulsating pressures in the flowing blood cause both flexible diaphragms to bend. The flexion in each of the two flexible diaphragms is cushioned by their respective air pockets.

As another alternative embodiment of the cuvette of the invention, the duct has a single flexible diaphragm 30a and a single air pocket 12b as shown in Figure 10, and also presents a modification of the rigid wall 32a to have a plurality of small holes or micro-striae on its surface inside the duct so that the holes are in contact with the blood flowing through the duct. The orifices, being relatively small in diameter, are practically impervious to the blood flowing through the duct, but they serve to dampen the force and absorb some of the pressure from the pulsation in the pulsatingly flowing blood. Air damping in the microstrose decreases the flex distance 62a and reduces both the acceleration and the speed of the flexible diaphragm 30a when it is moving between its extreme positions at distance 60a and distance 62a.

In the two alternative embodiments above, which are not shown, the additional damping with air in the cuvette of the invention serves to further increase the accuracy of the spectrophotometry readings by reducing the jerky movement and the rippling movement of the diaphragm. flexible during pulsating blood cycles.

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

The air bag 12b is preferably hermetically sealed by the formation of a hermetic joint between the lower frame 8a, the upper frame 12a and the circular periphery of the silicone membrane 30b.

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

Other means for receiving an emitting means and a sensing means are contemplated and need not be ring-shaped surfaces. Said means could have a structure that supports the emitting means and the detecting means in close proximity to the cuvette means so as to be spatially appropriate for the Incorporated Art.

At the inlet 16 of the bowl 10 there is a luer lock connector 80 and at the outlet 18 of the bowl 10 there is a luer lock connector 82. The luer lock connectors 80 and 82 are connected respectively to the catheters 110 and 112. The catheters 110 and 112 are integral with the installation of the cuvette 10 in the inlet hole 230 and the outlet hole 240 of the dialyzer 130.

Other equivalent embodiments of the cuvette are contemplated. However, due to the mathematics described in the Incorporated Technique, it is clear that the other ways of carrying out the cuvette in a

ES 2 165 877 T3 extracorporeal systems require a conduit in which blood flows unimpeded and steadily. The duct must have balanced dimensions with respect to the thickness of the blood layer in the duct (such as distance 60 in figure 8), the thickness of the flexible membrane in the duct (such as wall 30 in figure 8), and relative to the area of ​​the flexible membrane (such as area 36 in Figure 6). Each of these parameters is empirically adjusted, as shown in figure 12, so that the electronics, as mentioned in the Incorporated Art, provide optimal AC pulsation signals in order to use the ΔΙ / I in Equations B and C, which follow below.

b) Structural variables of the cuvette

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

(i) The transducer medium

The transducer medium will be modulated with each cycle of the pump medium pump, such as pump 140 in Figure 1, to produce a small Δd (such as a flexible line 62 in Figure 8) so that Equations A, B and C, which follow, are excited in the annotation ε _b1/and _b2. (See figure 11, as an example). A brief explanation of the mathematics, more extensively explained in the Embedded Technique, is the following:

The Beer-Lambert equation is the theoretical basis of the Incorporated Technique as taught by Equation A.

Equation A

I = Ioe^-£-^x'^{d + G}

Where:

G = Elongation factor of variable optical path

I = measured intensity

I_or = Incident intensity ε = Extinction coefficient of the mean

X = Concentration of the mean d = Thickness of the mean

In a practical hemodynamic application of Equation A, when the layer of blood d (such as distance 60 in Figure 8) is pulsed due to the pumping medium (such as pump 140 in Figure 1), the concentration x remains constant, but the blood layer d (such as distance 60 in Figure 8) will change by Δd (such as flexian line 62 in Figure 8) due to the transducer medium (such as wall 30 in Figure 8). Hence, taking the partial derivatives of Equation A with respect to time, and then dividing Equation A by itself, will give Equation B.

Equation B dI / dt dd

--Γ- = ^X '

Where:

»Δ.Ι« Δ I / I

I and -_b = extinction coefficient of whole blood.

In order to obtain the result in EQUATION B, it is assumed that G in Equation A is negligible due to the fixed dimensions of the duct.

As a practical application of Equation B, it is assumed that two wavelengths are now chosen, such as 805nm and 1300 nm, both wavelengths being isobastic, as was done in Technique

IS 2 165 877 T3

Incorporated. The Built-In Technique calls for a ratio of Equation B to be taken at the first wavelength to Equation B at the second wavelength. Thus, the ratio of ΔΙ / Ι at wavelength 1 to the ratio ΔΙ / Ι at wavelength 2 results in the cancellation of Δd and x, leaving only e_bi / e_b2 in Equation C.

Equation C (ΔΙ / Ι) ι = £ bi (ΔΙ / Ι) 2 'b'2

Where:

c_bn = Extinction coefficient in whole blood at the nth wavelength; yn = Wavelength

The advantage of selecting these wavelengths is such that, at these wavelengths, the ratio of the extinction coefficients, e_bi / e _b2, will be an important function of the hematocrit in the blood pulsating through the cuvette. This important function is graphically represented in figure 11.

The Built-in Technique is advantageous because various terms in the mathematics of the preceding equations eliminate themselves, as represented by Equations B and C, thereby greatly simplifying the determination of the concentration of a biological constituent in the flowing fluid. pulsating. As indicated, a principle component that is removed is the concentration x, which is divided into the proportion. Likewise, the term Ι is canceled_ο in Equation B, where Ι_ο is the term for the incident intensity.

(ii) Fins or grip zones

The fins or gripping areas, such as the fins 14 of Figures 3 to 6, are preferably thin to allow the installation of the cuvette in areas where a hermetic fit is necessary to facilitate the connection of the cuvette to the blood tubing. . The fins or gripping areas also make fingerprints on the flexible membrane less frequent, such as wall 30 in Figure 8, (iii) Flow volume in the cuvette.

The cuvette will preferably maintain adequate blood flow through the flexible membrane while allowing the larger volume to flow through the outer annular area, an example of which is wall 30 with flow path 40 as compared to wall 46 and flow path 42. A construction of this arrangement prevents turbulence in the conduit, slows down immediately below the flexible membrane sensor, and reduces sensitivity to flow in the conduit.

(iv) The flexible membrane

The stiffness of the flexible membrane depends on the thickness of the membrane and the area of ​​the membrane. Such stiffness reduces the absolute movement of d, which is the thickness of the blood flowing through the cuvette, such as distance 60 in Figure 8 or 60a in Figure 10, and hence its variation with aesthetic pressure. The membrane will preferably be flexible enough to allow Δd, such as the bending line 62 of Figure 8, with each ΔΡ, which is the change in pressure due to a pulsation in the flowing fluid. This ΔΡ facilitates the production of the ratio ΔΙ / Ι in Equation C and thus, the elimination by mathematical cancellation of the variable Ιο in Equation B.

Therefore, in order for the disposable extracorporeal cuvette of the invention to function effectively, it is important that there be a change, Δd, of d, such as flex line 62 and distance line 60, respectively, in Figure 8, or such as bending line 62a and distance line 60a, respectively, in FIG. 10. Preferably, Δd will be a function of the absolute thickness d of the blood sample, the thickness of the membrane, the aereal or the diameter of the membrane, the pressure change due to the pulsations in the fluid and the peak pressure change. at peak of the same, or the absolute pressure, and the elastic volumetric module of the plasty in which the flexible membrane and the cuvette are preferably constructed. These parameters allow Δd of the flexible membrane in the duct so that ΔΙ / Ι in Equation C would be operable, an example of which is shown in Figure 12. Figure 13 represents the error of the true value if Δd is improperly selected.

IS 2 165 877 T3

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, light emitters, and photodetectors. Sensor electronics and digital and analog schematics will also be used herein in connection with the invention being described.

Four. Determination of blood volume

In order to use the equations presented here and in the Incorporated Technique, determining the change in intravascular blood volume during hemodialysis using the hematocrit value requires the assumption that follows to be made. In renal dialysis, when the blood passes through the dialysis membrane, it is assumed that no red particles pass through the dialysis membrane. Only plasma or aqueous fluids, electrolytes and small molecules will pass through the dialysis membrane. Therefore, it can be said that any change in the hematocrit value due to dialysis is inversely proportional to the volume of blood according to Equations D to F that follow:

Equation D

PCV =

RBC

BV

Equation E

PCV₁

PCV2

RBC₁/ BV₁

RBC₂/ BV₂

Equation F

Hct₁

Hct₂

BV₁

Bv2

Where:

PCV_n = Volume of packed cell at time n.

RBC_n = Volume of red blood cells at time n,

BV_n = Volume of whole blood at time n, and

Hct_n = Hematocrit value at time n.

Since the volume of red blood cells (RBC) is a constant during dialysis, assuming that no red blood cells pass through the dialysis membrane, Equation F.

By monitoring the change in the hematocrit value from time n to time n + 1, the change in blood volume can be observed during that same time period. Specifically, when the rate of change of the blood volume of a patient is controlled during the duration of dialysis, the rate of plasma repositioning in said patient can also be determined. This system allows precise control of the hematocrit value, in order to accurately determine and control the percentage change in blood volume, which is directly related to the plasma replenishment rate when a person undergoes dialysis. Knowing this parameter, the doctor can adjust, in real time, the ultrafiltration index in order not to subject the patient to a reduced dialysis or to an exceeded dialysis.

As the foregoing discussion relates to non-invasive analysis of blood hematocrit data, it will be seen that the above-described emitter / sensor circuitry is equally suitable for in vitro analysis of blood hematocrit value.

The principles within the scope of the present invention require: (1) the optimization of fixed, spatial and geometric parameters in the cuvette and (2) that allow, at the same time, the flexible movement (Ad) of a membrane, and (3 ) that allow the ΔΙ / Ι calculation of Equation C for a quick and accurate determination of hematocrit and blood volume change. These principles are found in the preferred design.

IS 2 165 877 T3

The cuvette 10 conduit in the blood tubing system described herein is an optical path where blood flows through the conduit, having the preferred design depicted in Figures 1-8. The conduit allows light to flow through the conduit. passes from photoemitter 102 to photodetector 104 through the blood sample that passes through cuvette 10. The cuvette 10 is designed with a wall 30, which is a flexible membrane, which is moved with pulsation by the pressure changes in the blood conduit pipe due to the pump 140, which pulses the blood to through the extracorporeal conduction tubing.

The important elements in the cuvette are the thickness of the blood, such as the distance 60 in Figure 8, and the change in thickness of the blood, as shown by the flexion line 62 in Figure 8. This change is due to the pulsations created by the pump 140, the first and second opposing walls 30 and 32 and the areas of the same, and finally the volumetric plastic module of the first wall 30, all of this combine so that the First wall 30 fluctuates with each pulsation of the fluid.

Through cuvette 10, an emitter / detector apparatus 100 is placed. Photoemitter 102 then emits light through the blood sample in the conduit and is measured at photodetector 104, as shown in Figure 8. A computer 150 He then performs spectrophotometric techniques for the subsequent visualization in real time of the concentration values ​​of the constituent before and after the dialysis. From these values, if the concentration value sought is that of red blood cells, an instantaneous determination in real time of the change in blood volume can be obtained. In this way, hematocrit values ​​are obtained in a non-invasive way using electromagnetic radiation as an information vehicle. The present invention can be used in a disposable extracorporeal conduit system to determine the critical parameter of intravascular blood volume.

It will also be appreciated that the present invention will also provide a system and apparatus that can provide immediate and continuous monitoring data of the patient's blood volume. It additionally provides continuous, non-invasive data regarding the oxygen saturation status of the patient's blood, regardless of the patient's hematocrit value. Even under conditions of low blood perfusion to the patient, the extracorporeal control described here provides greater precision and ease of use.

The modes of realization described should be considered in all respects as merely illustrative and not restrictive. The scope of the invention, therefore, is indicated by the appended claims rather than in the foregoing description. All changes that are within the meaning and equivalence of the claims must be understood to be within the scope of the claims.