Abstract:
A method and device for the continuous real-time monitoring of relative blood volume change, based on registration of blood hemoglobin concentration, during long periods of time, such as dialysis session. A method and device for cardiac output measurement, based on optical dilution technique, during dialysis, surgeries, intensive care procedure. The effects of blood electrolyte composition change which result in a change of light beam geometry are eliminated by the relative orientation between a light source and a single light or photodetector. An illumination axis and a detection axis are oriented in an offset, non collinear configuration at a sufficient angle to substantially eliminate the scattering properties of the blood. A primary implementation of the device is extracorporeal paths, such as hemodialysis tubing systems, artery-vein extracorporeal artificial shunts, other extracorporeal systems. It also may be applied to vessels, tissues or to body parts being capable of transillumination.

Description:
Divisional of U.S. Ser. No. 08/950,244 filed Oct. 14, 1997 and issuing Mar. 21, 2000 as U.S. Pat. No. 6,041,246. 
    
    
     FIELD OF THE INVENTION 
     The present invention relates to the photometric analysis of blood properties to monitor changes in blood volume, blood proteins concentration, cardiac output and other hemodynamic parameters. More particularly, the invention includes a method and apparatus for employing a single light emitter and a single photodetector, the single photodetector is oriented with respect to the emitter to minimize a scattering effect of the light from the electrolyte composition of the blood, blood flow rate, blood hematocrit level and other factors. 
     BACKGROUND OF THE INVENTION 
     The optical density of blood corresponds to a number of factors and the measurement of the optical density of the blood has been used to determine certain blood parameters. 
     For example, during a hemodialysis session when fluid is being removed from the blood stream by a dialyzing machine, the concentration of hemoglobin, naturally occurring in the red blood cells, may decrease or, generally increase, with respect to the equilibrium processes of the fluid removal and mobilization from the body tissues. The change in the concentration of hemoglobin results in a corresponding change in the optical density of the blood and may be registered. 
     Monitoring changes in blood volume, as well as blood hemoglobin, can help to prevent such complications as hypotension due to continuous hypovolemia. Cardiac output of a patient, as well as other hemodynamic parameters of cardiovascular system, also provide useful information about the patient condition during hemodialysis, surgery, and in intensive care units. Using a dilution technique it is possible to calculate several characteristics of cardiovascular system, particularly cardiac output. 
     There are several research works and patents in the field of photometric blood monitoring. As shown in U.S. Pat. Nos. 3,830,569 to Meric; 4,243,883 to Schwartzmann; and 4,303,336 to Cullis, a device with suitable light source and photodetector may be used. Multiple detectors may be used, as shown by Meric. 
     As illustrated by U.S. Pat. Nos. 4,745,279 to Karkar et al.; 4,776,340 to Moran et al.; 5,048,524 to Bailey; and 5,066,859 to Karkar, additional detectors may be used to compensate various measurement artifacts, particularly for variations in intensity of light entering the blood stream. 
     However, the correlation between blood parameters and data obtained by photometric methods needs to be enhanced. 
     In U.S. Pat. No. 5,331,958 to Oppenheimer, a through photodetector and a remote photodetector are used for correction of light scattering due to sodium concentration fluctuations in the blood. Signals of the through and the remote photodetectors are amplified separately. The data from the remote detector is used to compensate or adjust the data of the through detector. These and other dual correction or compensation systems are subject to the inherent unreliability of multiple detectors, as well as calibration issues. 
     A disadvantage of other methods of photometric analysis of blood properties, which utilize conventional external tubing, includes a dependence on the physical characteristics of the tubing such as diameter, wall thickness, and light absorption factor. 
     Therefore, the need exists for a method and apparatus which require only a single light detector, wherein the light scattering effect is eliminated or minimized. In such an apparatus, the mechanical construction and data collection process is substantially streamlined as compared to multiple photodetector devices. There is a need for a method and apparatus having automatic calibration of the optical probe sensitivity for each particular tubing that the probe is employed with; and further having a dilution technique applicable for determining hemodynamic parameters of a patient. 
     SUMMARY OF THE INVENTION 
     In accordance with the present invention, there is a novel method and apparatus for the photometric analysis of blood properties. The present design allows for registration of blood optical density, provides signal adjustment for fluctuations in light scattering, which are related to blood electrolyte composition, as well as blood flow rate and other blood parameters. 
     The present apparatus employs a single light emitter projecting a light beam along an illumination axis and a single receiving photodetector for measuring light along a detection axis, wherein the illumination axis is non colinear, offset or angled with respect to the detection axis. According to the method, the present invention includes projecting a light beam into a blood column along an illumination axis, and detecting the light emerging from the blood column by a single receiving photodetector receiving system along a detection axis that is non colinear to the illumination axis. 
     The effects of light scattering are taken into consideration and have been eliminated by the location and orientation of the single receiving photodetector detection axis with respect to the illumination axis. It is understood the present system may be employed with additional sensors or detectors, however these sensors are not required, or related to the light scattering compensation, but are employed for different purposes. 
     Preferably a light emitter of isobestic wavelength is used to avoid the effect of hemoglobin oxygen saturation. 
     The device may be used with an extracorporeal tubing system through which blood flows, as well as being applied to patient&#39; vessels and body tissues which are capable of being transilluminated. Further, the device may be used during a dialysis session as well as surgeries, intensive care procedures, measurements hemodynamic parameters of blood by dilution technique, or other blood property altering events. 
     The present method includes utilizing biologically natural indicators, as for example isotonic saline, glucose, or another solution, for measurements by dilution technique and calibration procedures. There is no need in special markers, such as dye-green or the like, which stay in a patient&#39;s body for a long time. 
     In the drawings and in the detailed description of the invention there are shown and described only principal embodiments of this invention and are of illustrative nature only, but not restrictive. Other embodiments and technical realizations are applicable, all without departing from the scope and spirit of the invention. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 is a plot showing the light intensity changes caused by equal diluting volumes of isotonic saline, water and hypertonic saline. 
     FIG. 2 is a plot showing the light intensity changes caused by the blood flow rate changes. 
     FIG. 3 is a diagram showing the light scattering dependence on hematocrit level, blood electrolytes and blood flow rate. 
     FIG. 4 a  is a diagram showing a particular angle α in an emitted light beam that is characterized by stable light portion. 
     FIG. 4 b  is a diagram showing a principal dependence of absolute relative error of blood optical density measurement on an angle between Illumination Vector (IV) and Detection Vector (DV), where S NORM  is signal change while electrolyte composition remains unchanged, S HYPER  is signal change when electrolyte content was increased. An angle a is the angle of zero error, and still some vicinity Da is a range of angles where accuracy of measurement remains acceptable. 
     FIG. 5 is a plot showing registered light intensity signals for three bolus injections of equal volume of normal saline, water and 5% saline with a sensor constructed and located to adjust for the scattering effect. 
     FIG. 6 is a plot showing the light intensity changes caused by the blood flow rate changes registered with a sensor constructed and located to adjust for the scattering effect (lower trace) versus changes registered with traditional straight transilluminating sensor (upper trace). 
     FIG. 7 is a schematic of the present single emitter-single detector system. 
     FIG. 8 is a schematic of an alternative embodiment of the present single emitter-single detector system. 
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
     Generally, the present method and apparatus include a probe  10  having a housing  20 , a light source  30  and a receiving photodetector  40  for cooperating with a length of light transmissive tube  50 . 
     The light transmissive tube  50  may be any of a variety of conduits including but not limited to a dialysis tubing system, an artery-vein extracorporeal tubing shunt, or blood vessel. 
     The housing  20  is sized to operably retain the light source  30  and the photodetector  40 . The housing  20  includes a light port  22  through light from the light source  30  may pass. The housing  20  includes a tube receiving passage  24  for receiving a length of the light transmissive tube  50 , wherein which blood may flow through the tube. The housing  20  is constructed to dispose the tube  50  intermediate the light port  22  and the photodetector  40 . 
     Referring to FIG. 7, the probe  10  embodying the present apparatus is shown in an illustrative form. The probe  10  includes the light source  30  for generating a light beam. The generated light beam travels along an illumination vector IV. The illumination vector IV of the light may be dictated by the light source  30  itself, or by the configuration of the housing  20 . Specifically, the light port  22  of the housing  20  may be configured as a slit or aperture through which light from the light source  30  passes, and the housing thus permits only light passing along a predetermined illumination vector IV to intersect the tubing  50 . As shown in FIG. 7, in a first embodiment the light source  30  and housing  20  produce an illumination vector that is perpendicular to a longitudinal axis of the tubing  50 . 
     To avoid the need in correction for oxygen saturation, a light source  30  with an isobestic wavelength of emission spectrum may be used. The light source  30  is driven by light source driver  32 . The light source driver  32  may include optical feedback for source of supplied current, or may be a dedicated laser diode driver. 
     The receiving photodetector  40  is spaced from the light port  22  to locate a length of the light transmissive tube  50  therebetween. The receiving photodetector  40  is constructed to detect an intersection of light along a detection vector DV. As blood flows through the tube  50 , a blood column is exposed to the illumination vector IV and the photodetector  40 . Light initially entering the blood column along the illumination vector IV is absorbed and scattered in the blood flowing through the tubing  50 . The receiving photodetector  40  is oriented to receive only that part of the light passing through and emerging from the blood column that is scattered by the angle α with respect to the illumination vector IV. Thus, only light that is inclined from the illumination vector IV by the angle α reaches the receiving photodetector  40 . The detection vector DV of the receiving photodetector  40  is offset from the illumination vector IV by the angle α, thereby allowing only light scattered by angle α to reach the receiving photodetector. That is, the detection vector DV is inclined with respect to the illumination vector IV by the angle α. Although the receiving photodetector  40  is shown as coincident with the detector port, it is understood the detector port may define the interface with the blood column wherein the receiving photodetector is spaced from the blood column. That is, the application of secondary optics may allow the light source  30  to be spaced from the light port  22  and the receiving photodetector  40  to be spaced from the detector port. 
     Therefore, in the present design light is projected along a linear illumination vector IV and the detection vector DV of the receiving photodetector  40  is offset from the illumination vector by the angle α such that only light scattered by the angle α intersects the receiving photodetector. 
     The signal detected by the receiving photodetector  40  is amplified by an amplifier  42  to produce an amplified signal. The amplifier  42  may apply a logarithmic transformation function to be able to further operate in terms of optical density of the blood rather than merely detected light intensity. Conveniently, the amplified signal is applied to an analog-to-digital converter  44 , and passes into a microprocessor interface  46 . After processing the amplified and digitized signal to produce data, the microprocessor interface  46  causes the data to appear at a conventional monitor  48 , such as a computer. 
     An alternative embodiment of the optical probe  10  is shown in FIG.  8 . In this embodiment, the relative orientation between the light source  30  and the receiving detector  40  is modified. However, the illumination vector IV remains offset from the detection vector DV by the angle β. It is understood that the angle between the illumination vector IV and the detection vector DV, shown as α in FIG.  7  and β in FIG. 8 may not be identical. While the angles α and β could be equal, they may vary depending upon the specific configuration of the system. 
     As shown in FIG. 8, the probe  10  is constructed so that the housing  20  locates the light source  30  straight across from a receiving detector  40 . That is, the light port  22  and the photodetector  40  are located at the same longitudinal position along the tubing  50 , and are perpendicular to a longitudinal axis of the length of tubing as the tubing passes through the housing  20 . In this configuration, the illumination vector IV is non perpendicular to the longitudinal axis of the tubing  50 . That is, the illumination vector IV includes a component that extends along the longitudinal axis of the local length of the tubing  50  and hence blood column. In this configuration, the angle between the illumination vector IV and the detection vector DV is β. Locating the photodetector  40  and detection vector DV at the angle β from the illumination vector IV results in the photodetector registering the same portion of scattered light as the first configuration. 
     It is understood that some accommodation may be made for the defraction of the light as the light passes from the air through a wall of the tubing  50 , into the blood, from the blood into a second wall of the tubing and from the tubing into the air. However, it is believed such accommodation is relatively small and may be accounted for by a small shift in alignment of housing to correct diffraction by parallel layers of materials. 
     Theory 
     As stated, a change in hemoglobin concentration induces a change in the optical density of blood. An amount of light passing through a blood sample and emerging from the blood sample can be disposed and interpreted to determine blood properties. However, there are several factors that influence light scattering in blood. These factors include electrolyte composition, flow rate and hematocrit level of the blood. These influences introduce error in hemoglobin evaluation as well as injection bolus registration measurements made by a photodetector and light source having aligned illumination vectors and detection vector (that is, the detector is displaced straightway across the tubing from the light source at a common longitudinal position and illumination vector is collinear with the detection vector). Thus, in the systems of the prior art, the illumination vector and the detection vector are colinear. These influences are conveniently referred to as the “scatter effect” and are illustrated in FIGS. 1,  2  and  3 . 
     Referring to FIG. 4 a , as it has now been discovered, there is a particular angle α between the illumination vector IV of an emitted beam and a detection vector DV which is characterized by stable light portion, scattered in this direction. For fixed distance between a light source (distance from the light port  22  to the receiving photodetector  40 ), fixed blood column width and stable emitted light intensity, the signal received by the receiving photodetector at the angle α from the illumination vector IV is a function of hemoglobin concentration of blood and does not depend on light scattering. 
     Based on this understanding, the present optical probe  10  was designed. For this probe  10 , the scatter effect is essentially eliminated due to selection of adequate angle as illustrated in FIG. 4 b . Registered detected light intensity signals for three bolus injections of equal volume of normal saline, water, and 5% saline are shown in FIG.  5 . Blood flow related signals, shown in FIG. 6, also represents substantial improvement. 
     Signal Processing 
     Signal processing for the evaluation of the blood optical properties, changes in liquid balance of the blood and cardiac output calculation are based on following: 
     Assuming the number of erythrocytes remains constant during blood dialyzing, changes in liquid content induce changes in blood hemoglobin concentration. Relative volume change can be expressed as:                  Δ                 V       V   0       =     (     1   -       C   0     C       )             (     Equation                 1     )                                
     where C 0  is the initial, or basic hemoglobin concentration; C is the current hemoglobin concentration; V 0  is the initial volume and ΔV is the change in volume. Concentration can be evaluated from knowing the amount of light emerged from the blood by applying Beer-Lambert Law: 
     
       
         I=I s e −xCh    (Equation 2)  
       
     
     where I is intensity of light being absorbed emerging from the blood column of a thickness h; I s  is light intensity of the light source; k is the absorption coefficient; and C is concentration of the medium. Relative volume change can be expressed through the light intensity as:                  Δ                 V       V   0       =     (     1   -       ln          I   s       I   0           ln          I   s     I           )             (     Equation                 3     )                                
     where I s  is light intensity of the light beam; I 0  is the intensity of emerged light for initial concentration of blood hemoglobin and I is the intensity of emerged light for current meaning of concentration. 
     Relative hemoglobin concentration change can be expressed through the light intensity as:                  Δ                 C       C   0       =     (     1   -       ln        I     I   0           ln          I   s       I   0             )             (     Equation                 4     )                                
     An important parameter of the system is the sensitivity of the probe  10 . Different brands of tubing  50 , even different sections of tubing of the same brand may have different diameters and wall thickness. An absorption factor of the tubing material can also vary. Thus, sensitivity of the probe  10  should be known or determined for every particular tubing sample. It is also important to know any possible change in sensitivity of the probe  10  over relatively long periods of monitoring the blood properties. 
     A registered change in the optical signal received by the photodetector  40  is proportional to a relative concentration change,            Δ                 C       C   0       ,                          
     with a proportion coefficient (a factor for sensitivity of the optical probe  10  to relative blood hemoglobin concentration change), K:                  Δ                 C       C   0       =       K   ·   Δ                   U             (     Equation                 5     )                                
     where ΔU is the change in the optical signal received by the photodetector  40 , measured in volts, for example. To determine K, a known change is made in the relative hemoglobin concentrations and a corresponding signal change ΔU is registered. 
     Calibration 
     To provide correct data registering during long term monitoring of blood properties, measurements of hemodynamic parameters by dilution technique, and determining different absolute blood parameters, an optical sensor should be calibrated for particular operating conditions, such as current tubing system, initial blood parameters, as well as environmental changes occurring over the duration of the measurements. 
     When conventional hemodialysis is conducted, there are two major ways to calibrate the system. The calibration includes producing a known change in a blood optical property and registering the corresponding change in the measured (detected) signal. 
     In a first, more preferable method, the probe  10  is placed on the venous side of the dialysis tubing system, downstream from the dialyzer. In the first method of calibration, the ultrafiltration rate is changed by a known amount, or completely terminated. When the ultrafiltration rate is altered, the venous flow rate becomes greater by ultrafiltration rate change Q UF . The change in the ultrafiltration rate produces a decreasing hemoglobin concentration. 
     Relative concentration changes in a venous line can be expressed through blood flow rates as follows:                        Δ                 C       C   0                 v     =     -       Q   UF       Q   B                 (     Equation                 6     )                                
     where Q B  is the known blood flow rate in an arterial side and Q UF  is the known ultrafiltration rate change. Thus, desired coefficient K is:              K   =       -       Q   UF       Q   B         *   Δ                   U   UF               (     Equation                 7     )                                
     where ΔU UF  is change in the detected registered signal as caused by the change in the ultrafiltration rate. 
     The relative change in hemoglobin concentration of the arterial blood is calculated through a measured venous side concentration change according to:                               Δ                 C       C   0                 AR     =       (     1   -       Q   UF       Q   B         )        Δ                   C     C   0                V           (     Equation                 9     )                                
     where                   Δ                 C       C   0                 AR                          
     is the relative change in arterial hemoglobin concentration,                   Δ                 C       C   0                 V                          
     is the relative change in venous hemoglobin concentration, Q UF  is ultrafiltration rate, and Q B  is blood flow rate in the arterial side. 
     Dialysis machines of some brands turn ultrafiltration off during dialysis session several times to conduct internal self-calibration. Advantageously, these self-calibrating procedures can be used for calibration of the sensitivity optical probe to the relative concentration change. 
     Another method of calibrating the sensitivity of the probe is to employ an indicator which dilutes the photometric density of the blood passing through the tubing  50 . For example, an injection of an indicator, such as normal saline or another conventional solution, is made into an injection port of the dialysis system. The indicator injection should be made upstream from the probe  10 , so that all the indicator passes through the probe between the light port  22  and the photodetector  40 . The indicator injection causes the optical property of the blood to change, thus changing the light intersecting the photodetector  40  along the detection vector. The change in the amount of detected scattered light is registered by the probe  10 , and specifically the photodetector  40 . Coefficient K can then be calculated through the volumes:              K   =       -       V   Inj       V   +     V   Inj           *   Δ                   U   Inj               (     Equation                 8     )                                
     where V inj  is injection&#39;s volume; V=(Q B −Q UF )*ΔT inj , ΔT inj  is the transit time of the injection bolus; and ΔU inj  is integrated over ΔT inj  registered signal change. 
     The calibration can also be made with the routine injection of a dilution indicator into an injection port on an arterial side of the dialysis system. The probe  10  is then placed on arterial side, downstream from the place of injection. Again, all the injected indicator must pass through the probe  10 . No recalculation is required in this case. 
     A calculation of cardiac output (CO) of a patient can be made by dilution techniques with the following method. 
     By injecting a known amount of an indicator into a blood flow, the diluting effect of the indicator over a period of time can be accurately determined by a sensor responsive to changes of photometric properties of the blood. The sensor is positioned so that the indicator passes the sensor, with the measured diluting effect being used to determine various blood parameters. The measurement may be made in an extracorporeal blood system in which optical measuring probes are secured, for example, to tubing leading to exterior blood treatment equipment such as hemodialysis machine, extracorporeal artery-vein passive shunt, or the like. Thereafter, a bolus, or known volume, of an indicator material is injected into the bloodstream, and measurements are made of changes in the photometric properties of the blood to determine the passage of the bolus past the optical probe  10 . The changes in such photometric properties of the blood can then be plotted and used to determine various hemodynamic parameters, particularly cardiac output of a patient. 
     With the aim of calibration for optical properties of different tubing, the above mentioned calibration injection of known amount of an indicator may be employed. The calibration injection must be made directly prior to the location of the probe, while simultaneously measuring the blood flow through the blood system. A blood flow measuring device can be, for example, ultrasonic flow meter such as those manufactured by Transonic Systems, Inc., Ithaca, N.Y. The cardiac output will be calculated as follows:              CO   =           V   ind     ·     S     cal   -   ind             V     cal   -   ind       ·     S     b   -   ind           ·     Q     A   -   cal                 (     Equation                 13     )                                
     where V cal-ind  is the volume of the calibrating injection, S cal-ind  is the dilution area under a curve generated by the measurement of dilution produced by a calibrating injection. 
     Generally, cardiac output of a patient&#39;s heart may be calculated as:              CO   =         K   s     ·     V   ind         S     b   -   ind                 (     Equation                 14     )                                
     where V ind  is total volume of injected indicator; S b-ind  is the area under a dilution curve representing the total optical density changes in the blood column over a time period; and K s  is the calibration coefficient for a particular set of tubing and probe, as determined earlier. 
     As an advantage of using the probe for cardiac output measurement, it should be noted the insensitivity of the probe to temperature variations in blood, indicator, and the tubing carrying the blood. Further, since the probe can be secured in extracorporeal tubing system such as passive arterial-venous shunt or dialysis tubing, the present method of cardiac output measurement substantially reduces the invasive penetration into a patient&#39;s vital organs, such as the heart, as compared to Swan-Ganz catheter. 
     While a preferred embodiment of the invention has been shown and described with particularity, it will be appreciated that various changes and modifications may suggest themselves to one having ordinary skill in the art upon being apprised of the present invention. It is intended to encompass all such changes and modifications as fall within the scope and spirit of the appended claims.