Abstract:
Systems and methods are provided for noninvasively measuring the levels of urea, blood osmolarity (or Na + ), plasma free hemoglobin and tissue water content in a patient&#39;s blood or tissue. Light of selected wavelengths is passed through blood or body tissue and the transmitted or reflected light is detected and the detected signals can be electronically compared and manipulated to provide the non-invasive, continuous and quantitative display of a patient&#39;s blood urea, blood osmolarity (or Na + ), plasma free hemoglobin and tissue water content.

Description:
RELATED U.S. APPLICATIONS 
     This is a Continuation-In-Part of Ser. No. 08/479,352, filed Jun. 7, 1995, now U.S. Pat. No. 5,803,908 which is a continuation of Ser. No. 08/317,726 filed Oct. 4, 1994, now U.S. Pat. No. 5,499,627, which is a divisional of Ser. No. 08/011,882 filed Feb. 1, 1993, now U.S. Pat. No. 5,372,136, which is a continuation of Ser. No. 08/598,169 filed Feb. 7, 1996, now U.S. Pat. No. 5,734,502, which is a Continuation of application Ser. No. 07/598,169 filed Oct. 16, 1990 now abandoned. 
    
    
     FIELD OF THE INVENTION 
     This invention relates to systems and methods for noninvasively and/or continuously and quantitatively measuring spectrophotometrically a patient&#39;s blood urea nitrogen, blood osmolarity, plasma free hemoglobin, and tissue water content. 
     BACKGROUND 
     Modern medical practice utilizes a number of procedures and indicators to assess a patient&#39;s condition. Blood urea nitrogen (“BUN”), plasma free hemoglobin (“PFH”), and tissue water content are important indications of a patient&#39;s condition. 
     BUN, which is the amount of urea or urea nitrogen per unit volume of blood expressed, typically, in milligram percent units (mg %) is typically a by-product of the catabolism of various body proteins principally found in muscle and liver tissues. It is present in extra (and intra) vascular spaces, but later processed and excreted through the kidneys into the urine. Specifically, in the case of end-stage renal disease patients, acute renal failure, or chronic renal failure, wherein the kidneys do not function properly to excrete this waste product, the BUN levels elevate. Subsequently, urea becomes a toxin to many other organ systems of the body including the brain, heart, skin, etc. 
     Medical professionals routinely desire to know the BUN, or dialysate urea or dialysate urea nitrogen (DUN) value, of the patient, because of the above-mentioned deleterious and serious side effects. To determine BUN using any of the techniques available today, it is necessary to draw a sample of blood by veni-puncture. Then, using widely accepted techniques, the sample of blood is subjected to biochemical and enzymatic reactions to determine the level of urea in the blood. 
     Conventional techniques require that a sample of blood be withdrawn from the patient for in-vitro analysis. Any invasion of the subject to obtain blood is accompanied by the problems of inconvenience, stress, and discomfort imposed upon the subject. The infectious risks are also present when the body is invaded, via needle-skin puncture. Additionally, withdrawing blood also creates certain contamination risks to paramedical professionals. Moreover, even in a setting where obtaining a blood sample does not impose any additional problems, for example during surgery, the available techniques require delay between the time that the sample is drawn and the BUN value is correctly processed. Still further, none of the previously available techniques allow for continuous monitoring of the subjects BUN as would be desirable during hemodialysis treatment procedures or even in intensive care treatment. 
     Specifically in hemodialysis, recent techniques have been developed to enzymatically determine the BUN level in the dialysate fluid as a marker of what is transpiring in the blood. However, these particular enzymatic techniques are likewise fraught with serious drawbacks, not the least of which is that the technique does not give a continuous measurement of the urea nitrogen even in the dialysate fluid. The enzymatic determination is accomplished by periodic, automated, sampling, wherein the enzymatic compounds and dialysate fluids are mixed and the urea nitrogen level is thereby determined. 
     Medical professionals routinely desire to know the tissue water (or hydration status) of the patient. For example, in hemodialysis (or in end-stage renal disease) patient tissue water increases dramatically due to the inadequate elimination of water from the interstitial and intravascular spaces, since the kidney no longer correctly functions. Hence, the patients become edematous and their tissue water content increases dramatically. In hemodialysis, the goal of therapy is to remove all of the toxins of the blood and body. Some of these toxins are the urea, potassium, and even water which can become a significant toxin to the patient. Therefore, removal of water from the tissue is crucial because this water overloaded state requires excessive energy expenditure by the heart to function. Hence, many dialysis patients are in a state of pulmonary edema, congestive heart failure, etc., due to the large “load” that the heart must push against. Therefore, to reach an appropriate “dry weight” (the patient&#39;s body weight when the kidneys were functioning normally) is an important dialysis therapy goal. 
     Medical professionals in other specialties are desirous of knowing the tissue water content of non-renal, edematous patients for other reasons. Hormonal imbalances, menstrual cycle variations, congestive heart failure and other causes also result in pulmonary edema and peripheral edema. These states require the knowledge of the interstitial tissue water content. 
     Tissue water content is conventionally measured by bioelectrical impedance; however, bioimpedance can be costly and requires the injection of small electrical currents into the patient. Another technique involves measuring the amount of water in the tissue spaces by injecting radio-isotopes into a patient. This is done principally on a research basis, however, because of the attendant radiation risks. 
     PFH (Plasma Free Hemoglobin) is the amount of hemoglobin not contained inside a red blood cell, but rather free in plasma solution and is expressed, typically in milligram percent units (mg %). PFH is typically a result of red blood cell breakage or hemolysis, with spillage of the hemoglobin directly into the plasma. Specifically, in the case of end-stage renal disease patients, acute renal failure, or chronic renal failure, wherein the kidneys do not function properly and hemodialysis is required, the PFH levels in the blood may elevate due to tubing lines kinking and pump rollers crushing the red blood cells during the course of the hemodialysis treatment. This can occur in cardio-pulmonary surgeries as well. Subsequently, PFH itself becomes a toxin to many other organ systems of the body. 
     Medical professionals desire to know the PFH of the patient, because of the above mentioned deleterious and serious side effects associated with the presence of PFH. In conventional techniques, PFH is measured by drawing a sample of blood by veni-puncture. Then, using widely accepted techniques the sample of blood is subjected to biochemical reactions to determine the level of PFH in the plasma of the blood. 
     Blood osmolarity is the osmolar content of blood per unit volume of blood expressed, typically, in milliosmolar units. The osmolar content of blood (and/or the sodium content) should have a narrow range of values due to the body&#39;s compensatory abilities. However, in the case of end-stage renal disease patients, acute failure, or chronic renal failure, wherein the kidneys do not function properly and hemodialysis is required, the blood osmolarity varies greatly. 
     Medical professionals routinely desire to know the osmolarity or sodium value of the patient, because of the deleterious and serious side effects associated with levels outside the normal range. To determine the blood osmolarity (OSM) or blood sodium (Na + ) content using any of the techniques available today, it is necessary to draw a sample of blood by veni-puncture. Then, using widely accepted techniques the sample of blood is subjected to physical and biochemical reactions to determine the level of OSM or Na +  in the blood. 
     In view of the drawbacks in the available art dealing with invasive blood constituent determinations, it would be an advance in the art to noninvasively and quantitatively determine a subject&#39;s blood constituent including BUN, PFH, tissue water, osmolarity, and Na + . It would also be an advance to provide a system and method for noninvasive blood constituent monitoring which utilizes electromagnetic emissions as the information carrier for information relating to BUN, PFH, tissue water, osmolarity, and Na + . 
     OBJECTS OF THE INVENTION 
     It is an object of the present invention to provide a method and system for noninvasively measuring and monitoring tissue water. 
     It is another object of the present invention to provide a method and system for noninvasively measuring and monitoring urea in the blood or dialysate fluid. 
     It is a further object of the present invention to provide a method and system for noninvasively measuring and monitoring PFH in the blood. 
     It is yet an object of the present invention to provide a method and system for noninvasively measuring and monitoring osmolarity and Na +  of the blood. 
     SUMMARY OF THE INVENTION 
     In general, the present invention is directed to apparatus and methods for determining a biologic constituent value, such as BUN, PFH, tissue water and osmolarity, transcutaneously, continuously, and noninvasively. 
     One aspect of the present invention provides a method and system for the noninvasive measurement of tissue water content in which an emitter and a detector are positioned outside of a patient (either remote from the patient or, preferably, mounted on the skin of a patient), light from the emitter passes through a portion of a patient&#39;s body and is received by the detector. The detector measures light of a selected wavelength that is absorbed by water. The intensity of light at an initial time is measured by a detector; then at a later time, typically after a period of dialysis, the intensity of detected light is compared with the initial intensity and the percent change in tissue water is calculated. 
     In an alternate method, light of at least two selected wavelengths are detected and absorptions due to blood and/or skin tissues are accounted for. Thus a measurement of intensity transmitted (i.e. detected) versus intensity emitted at a wavelength selected for absorption by water allows calculation of the volume percent of water in a patient&#39;s tissue. 
     Another aspect of the present invention also provides a method and system for the measurement of urea concentration in the blood of a patient by measuring the quantities of light at two wavelengths where the function of the extinction coefficient versus urea concentration at each given wavelength holds BUN information that is different in at least one of curvature, offset, linearity or sign from the other wavelength. 
     A further aspect of the present invention provides a method and system for measuring osmolarity (or sodium content) of a patient&#39;s blood. 
     Yet another aspect of the present invention provides a method and system for measuring plasma free hemoglobin through an optical technique. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 shows a typical hemodialysis tubing circuit and connections. 
     FIG. 2 shows a perspective view of apparatus used to noninvasively measure tissue water content. 
     FIG. 2A shows an enlarged cross-sectional view of a patient&#39;s finger in the apparatus of FIG. 2, configured in a transmission mode. 
     FIG. 3 shows absorption coefficient values for oxyhemoglobin, reduced hemoglobin and water. 
     FIG. 4 shows a plot of Log (Io/I) at 1300 nm wavelength versus hematocrit of a blood sample as it is diluted with saline to show that intensity of light has slight dependence on Hematocrit. 
     FIG. 5 shows an absorption spectrum of 14% urea at room temperature. 
     FIG. 6 represents a hypothetical plot of log light intensity vs. urea concentration (milligram percentage, mg %). 
     FIG. 7 shows plots of the % change in absorbance of a beam of light at 810 nm and 810 nm/1300 nm v. sodium. 
     FIG. 8 shows plots of hematocrit (HCT) as measured by three techniques vs. mean cell volume (MCV). 
     FIG. 9 shows plots of hematocrit (HCT) as measured by three techniques vs. mean cell volume (MCV). 
     FIG. 10 shows plots of the % change in absorbance of light at 810 nm and 810 nm/1300 nm vs. PFH. 
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     In general, the present invention is directed to apparatus and methods for determining a biologic constituent value transcutaneously, continuously, and noninvasively. This is achieved by passing at least one wavelength of light onto or through body tissues such as the finger, earlobe, or scalp, etc., see FIG. 2 and 2A, and then compensating for the effects of the other non-water, body tissues using a modified Beer Lambert Law as a theoretical basis. An example of such a measurement technique is found in U.S. Pat. No. 5,372,136 and is incorporated herein. 
     Although the present invention will describe in great detail the transillumination of blood in an extracorporeal conduit, it will be appreciated that reflectance spectrophotometry may alternatively be employed when transillumination is difficult to accomplish. As used herein, the term “body part” is intended to include skin, earlobe, fingertip, lip, etc., but also “extracorporeal conduit”, see FIG. 1, may refer to a disposable blood chamber or in-vitro blood containers such as tubes and cuvettes. 
     In preferred embodiments, measurements are conducted using the apparatus (or modified versions thereof) described in U.S. Pat. Nos. 5,456,253 (column 1, line 18 through column 14, line 67; FIGS. 1-13) and 5,372,136 (column 1, line 12 through column 14, line 39; FIGS. 1-16), and U.S. patent application Ser. No. 08/479,352 which are incorporated herein as if reproduced in full below. 
     By way of background, kidneys are located on either side of the spine. In a healthy patient, kidneys function to stimulate red blood cell production and regulate the content of the blood. Kidneys also produce hormones that affect other organs and control growth. When functioning properly, kidneys serve as a means for cleaning the blood by removing excess fluids and toxins. The filtering task in each kidney is performed in part by the some one million nephrons in the kidney. The nephrons are filtering units made up of tiny blood vessels. Each such blood vessel is called a glomerulus. Every day, roughly 200 quarts of blood and fluids will be processed by the kidney. The kidney removes about two quarts of water and toxic chemicals which are sent to the bladder as urine for subsequent voiding thereof by urination. 
     A patient whose kidneys are performing substandardly may be dialyzed as a substitute for the blood cleansing function normally performed by property functioning kidneys. Dialysis is a process by which the function of the kidney of cleaning blood is substitutionarily performed. The process of dialysis was perfected for routine use in the 1960&#39;s, having been invented some 50 years ago. For the purposes of discussion and illustration of hemodialysis, FIG. 1 is now referred to. While FIG. 1 incorporates a view of a presently preferred embodiment of the present invention, it also incorporates a view of some common components which are typical in a general hemodialysis environment. The general environment of hemodialysis and typical components therein will now be discussed. 
     In hemodialysis, blood is taken out of a patient  200  by an intake catheter means, one example of which is shown in FIG. 1 as an input catheter  122 . Input catheter  122  is intravenously inserted into patient  200  at a site  180  and is used for defining a blood passageway upstream of a blood filter used to filter the impurities out of the blood. The blood filter is also called a dialyzer  130 . The unclean blood flows from an artery in patient  200  to a pump means, an example of which is pump  140 . From pump  140 , the blood flows to dialyzer  130 . Dialyzer  130  has an input port  230  and an output port  240 . The pump  140  performs the function of moving the unclean blood from patient  200  into input port  230  through dialyzer  130 , and out of dialyzer  130  at output port  240 . 
     Specifically, unclean blood in input catheter  122  is transported to input port  230  of dialyzer  130 . After passing through and being cleansed by dialyzer  130 , the blood may receive further processing, such a heparin drip, in hemodialysis related component  300 . The now clean blood is returned to patient  200  after the dialyzing process by means of an output catheter means, an example of which is output catheter  124 . Output catheter  124 , which is also intravenously inserted into patient  200  at site  180 , defines a blood passageway which is downstream from dialyzer  130 , taking the blood output by dialyzer  130  back to patient  200 . 
     As mentioned, the hemodialysis process uses a blood filter or dialyzer  130  to clean the blood of patient  200 . As blood passes through dialyzer  130 , it travels in straw-like tubes (not shown) within dialyzer  130  which serve as membrane passageways for the unclean blood. The straw-like tubes remove poisons and excess fluids through a process of diffusion. An example of excess fluid in unclean blood is water and an example of poisons in unclean blood are blood urea nitrogen (BUN) and potassium. 
     The excess fluids and poisons are removed by a clean dialysate liquid fluid, which is a solution of chemicals and water. Clean dialysate enters dialyzer  130  at an input tube  210  from a combined controller and tank  170 . The dialysate surrounds the straw-like tubes in dialyzer  130  as the dialysate flows down through dialyzer  130 . The clean dialysate picks up the excess fluids and poisons passing through the straw-like tubes, by diffusion, and then returns the excess fluids and poisons with the dialysate out of dialyzer  130  via an output tube  220 , thus cleansing the blood. Dialysate exiting at output tube  220  after cleansing the blood may be discarded. 
     The general hemodialysis process and environment is seen in FIG.  1  and has been described above. A summary of this process is that patient  200 , whose kidneys are performing substandardly, is dialyzed. The unclean blood flows from an artery in patient  200  to the pump  140  and then to dialyzer  130 . Unclean blood flows into dialyzer  130  from input catheter  122 , and then clean blood flows out of dialyzer  130  via output catheter  124  back to patient  200 . 
     It is preferable that the pump  140  causes the blood flowing into, through, and out of dialyzer  130  to flow in a pulsatile fashion. 
     Installed at either end of dialyzer  130  is a spectrophotometry means for defining a blood flow path, for emitting radiation into the blood in the flow path, and for detecting radiation passing through both the blood and the flow path. The spectrophotometry means includes a cuvette means  10  for defining the blood flow path, and an emitter/detector means  100  for directing and detecting radiation. Within the emitter/detector means is both an emission means for directing radiation and a detector means for detecting radiation. Once such spectrophotometry means is discussed in detail in U.S. Pat. No. 5,456,253 and is incorporated by reference herein. 
     Emitter/detector apparatus  100  enables the detection by a photodetector (not shown) of the portion of radiation which is directed by a photoemitter (not shown) to cuvette  10  and passes through both the blood therein and the cuvette  10 . As shown in FIG. 1, the cuvette  10  is installed at either end of dialyzer  130 . Each cuvette  10  has a photoemitter and a photodetector thereon. 
     The emitter/detector means is electrically connected to a calculation means. In a preferred embodiment of the system, an example of the calculator means is depicted in FIG. 1 as computer  150  which is electrically connected to the photoemitter and the photodetector on emitter/detector apparatus  100  by means of cable  120 . 
     Intake catheter  122  takes blood to cuvette  10  situated before input port  230  of dialyzer  130 . Emitter/detector apparatus  100  at input port  230  of dialyzer  130  subjects the blood therein to at least two radiation wavelengths of electromagnetic radiation for the purposes of analysis, via spectrophotometry, so that the concentration of a desired biological constituent can be derived. Each photodetector, at both input port  230  and output port  240  of the dialyzer  130 , communicates the detected radiation at least a first and a second wavelength via cable  120  to computer  150 . 
     Computer  150  calculates both before dialysis and after dialysis concentrations of the sought-after or desired biological constituent. Computer  150  then displays, respectively, at a first display  152  and a second display  154 , the derived concentration of the biological constituent in their analogue or digital representations. 
     It should be understood that the improvements and modifications of the present invention can be applied to a wide variety of blood monitoring apparati and, thus, are not limited to certain preferred embodiments such as those described in the above-cited U.S. patents. 
     The theoretical basis for the spectrophotometric technique mentioned above is the Beer Lambert Law as shown below. 
     
       
         I=I o e −E(x)d    (1)  
       
     
     Wherein I o  is the intensity of the incident source radiation, I is the transmitted intensity of the source radiation through the sample, E is the extinction coefficient of the sought for constituent, x is the concentration of the sample constituent in the tissue (or blood conduit), and d is the optical path length (distance). 
     Utilizing the Beer-Lambert Law, quantitative measurements have not been possible in the body or whole blood since the scattering of the incident photons passing into and through the measuring container (or skin) regions is extensive and highly variable. This scattering spoils the Beer-Lambert Law by adding a variable loss of radiation to the measurement and also extends the path length of the incident radiation by an unknown amount as well. Therefore modifications to the Beer-Lambert Law are required and will be hereafter shown. 
     Since it is important to know when the patient&#39;s interstitial water has been completely removed from the tissue or that the patient&#39;s “dry weight” has been achieved, the following equations describe the methodology by which a value or graphic representation, digital or analog, can be determined. A modified Beer Law equation for tissue can be approximated as: 
     
       
         I≈I o e −[(3K(K+S))     ½     ]d    (2)  
       
     
     When K and S are the bulk absorbance and scattering coefficients of tissue. For human tissue K is &lt;&lt;S and hence for small changes in tissue absorption a more accurate expression is:                I   =       I   o     ·     α     S   ·     d   2         ·            -   α     ·   d           ,       where                 α     =       (     3      KS     )     ½               (   3   )                                
     Since S, the scattering term, is relatively constant for a given tissue, the absorbance term, K, takes the form: 
     
       
           K=K   b   X   b   +K   w   X   w   +K   s   X   s    (4)  
       
     
     Where: 
     K b =absorbance due to blood 
     K w =absorbance due to water 
     K s =absorbance due to skin tissues 
     X b =volume of blood per volume of tissue 
     X w =volume of water per volume of tissue and 
     X s =volume of skin per volume of tissue. 
     MEASURING TISSUE WATER 
     In hemodialysis usage, the tissue water content is most important, since the clinician is attempting to measure the tissue hydration status after which no further water can be removed. Hence measuring the optical power at about 1300 nm from the start (initial, (I 1 ) 13 ) of hemodialysis to some time (t) is given from equation 3 measured at separate d&#39;s, d 1  and d 2 ; and where I 01 =I 02 :                  log                     (         I   1          d   1   2           I   2          d   2   2         )     i       =         -     [       (     3        K   i          S   i       )     ½     ]       ·   d                   Δ                     and           (   5   )                 log                     (         I   1          d   1   2           I   2          d   2   2         )     t       =         -     [       (     3        K   t          S   t       )     ½     ]       ·   Δ                   d             (   6   )                                
     To determine the % change in X w , % ΔX w , the following obtains:                {     log                       (         I   1          d   1   2           I   2          d   2   2         )     i     /   log                       (         I   1          d   1   2           I   2          d   2   2         )     t       }     =     γ   =       (       K   i            S   i     /     K   t            S   t       )     ½               (   7   )                                
     Again, since S i ≈S t  and substituting equation 4 into equation 7, the following obtains: 
     
       
         γ=[( K   b   X   b   +K   w   X   w   +K   s   X   s ) I /( K   b   X   b   +K   w   X   w   +K   s   X   s ) t ] ½   (8)  
       
     
     But at a wavelength of about 1300 nm, K b  and K s  are about equal to K w , hence:                γ   2     =         K   w     ·     (     1   +       X   b     /     X   wi       +       X   s     /     X   wi         )     ·     X   wi           K   w     ·     (     1   +       X   b     /     X   wt       +       X   s     /     X   wt         )     ·     X   wt                 (   9   )                                
     and hence: 
     
       
         γ 2   ≈X   wi   /X   wt    (10)  
       
     
     Therefore the percentage change in tissue water (X w ) becomes: 
     
       
         [γ 2 −1]×100=[( X   wi   /X   wt )−1]×100   (11)  
       
     
     As mentioned above, for the determination of the tissue water content in patients with congestive heart failure, etc., the absolute value of the tissue water may be desired. Hence, the following indicate the mathematical operations required to determine the absolute value of tissue water, see FIG.  6 . The following operations indicate the need for additional wavelengths in order to eliminate or compensate for the scattering effects of other competing biologic constituents in tissue. From equations 3, 4, 5 and 6 knowing Δd, it is necessary to measure S exactly (or eliminate S) to calculate X w  directly. S has been measured previously and found to be 0.75/mm, and relatively constant for human tissue. There may be a need to collect data from additional wavelengths in order to eliminate or compensate for the scattering effects of other competing biologic constituents in tissue. 
     In summary, one aspect of the present invention is directed to apparatus and methods for determining the biologic constituent value, the tissue water value, transcutaneously and noninvasively. This is achieved by passing at least one wavelength of light onto or through body tissues such as the finger, earlobe, scalp, etc. and then compensating for the effects of other body tissues not related to water. The light can also be passed directly through blood in a conduit. In one embodiment within the scope of the present invention, the wavelength of light is selected to be near 1300 nanometers (nm). At that particular wavelength, blood is almost independent of the hematocrit value but the water absorption coefficient at 1300 nm is very large compared to that of blood. Hence, the measurement at 1300 nm is independent of the hemoglobin content of the tissue per se. Another significant advantage of the present invention is the capability of monitoring multiple wavelengths simultaneously other than 1300 nm, where water absorption is even greater than that at 1300 run. However, at those wavelengths (1480 nm, 1550 nm, 1800 nm and 1900 nm) the simultaneous compensation for the hemoglobin value is required. 
     DETERMINING BLOOD UREA NITROGEN 
     A modified Beer-Lambert equation can also utilized for the determination of urea in the blood as follows: 
     
       
           I=I   o   e   [−E(x)d+B(d, E(x))]   (12)  
       
     
     where B (d,Ex)) is an optical pathlengthening function, Io is the intensity of the incident source radiation, I is the transmitted intensity of the source radiation through the sample, E is the extinction coefficient of the sought for constituent, x is the concentration of the sample constituent in the tissue (or disposable blood conduit) and d is the optical separation distance. 
     To determine BUN according to the present invention, a measuring wavelength (M) and a reference wavelength (R) must be selected. These wavelengths may be selected close enough to one another such that the pathlengthening factors are approximately the same for each wavelength (longer wavelengths are preferred since they exhibit less sensitivity to scattering). For example, the selection of a measuring wavelength at 2190 nm and a reference wavelength at 1900 nm may be appropriate since the scattering functions (pathlengthening factors) are approximately the same at these wavelengths, and the difference between the peak BUN absorption at 2190 nm and the minimal urea absorption at 1900 nm holds significant BUN information, as seen in FIG.  5 . 
     The actual function of extinction coefficient E (either E M  or E R ) versus the urea concentration at each given wavelength must hold BUN information that is different in at least one of curvature, offset, linearity, or sign from the other wavelength, see the hypothetical curves in FIG.  6 . If the functions of E versus urea concentration are not sufficiently different for each wavelength, then the ratio E M /E R  will not hold BUN information. Even though wavelengths of 2190 nm and 1300 nm are the preferred wavelengths (see FIG.  5 ), it will be appreciated that other wavelengths such as 2200 nm (2980 nm or 6160 nm) and 1480 nm (or 1900 nm) may also satisfy the condition of having adequate urea detected with respect to water. 
     For monitoring the BUN in living tissue of a patient, the pulsatile characteristics of the blood require the utilization of the form and mathematical operations presented in U.S. Pat. No. 5,372,136, and using the described ΔI/I technique in order to eliminate certain intrinsic tissue and extrinsic light source effects. 
     In non-pulsatile applications such as hemodialysis or where blood or dialysate is flowing through a chamber or cuvette, logarithmic operations will give the appropriate computed BUN values, as shown in the following formulation for the two example wavelengths: 
     
       
         [ BUN]   raw   =log  ( I/I   o ) 2190   /log  ( I/I   o ) 1300   =E   2190   /E   1300    (13)  
       
     
     But since the whole blood medium will have some hematocrit dependence at 2190 nm, use of the following formula compensates for the hematocrit effects: 
     
       
         [ BUN]   corrected   =[E   2190   /E   1300   ]•[F [log  ( I/I   o ) 8   /log  ( I/I   o ) 13 ]]  (14)  
       
     
     where F[(log 8 /log 13 )] is a function of the hematocrit. It is likely that other competing substances will be detected at 2190 nm, those can also be compensated with similar functional operators. 
     In a preferred embodiment, the electronic structure and memory components for a BUN measuring system are similar to that described in U.S. Pat. No. 5,372,136. In some preferred embodiments, the wavelengths 1300 nm, 1800 nm, 1900 nm, and 2190 nm are selected. Telcom Device Corp. of Camarillo, Calif. manufactures the corresponding LEDs with product numbers: 1300 nm LED, 1.8 LED, 1.9 LED and 2.2 LED. A preferred source for the detector may be photodiode, PD24-04, manufactured by IBSG, St. Petersburg, Russia. Although the foregoing discussion relates to noninvasive analysis of BUN (or dialysate urea nitrogen, DUN) information in the hemodialysis setting, it will be appreciated that emitters, sensors, and circuitry can be adapted for invasive, in-vitro analysis of BUN, or the transcutaneous, in-vivo analysis. 
     To summarize, one embodiment of the present invention that measures BUN, one wavelength of light is selected to be at or near the peak absorption level of urea and another wavelength (the reference) selected at an absorption minimum of urea (or urea nitrogen) with respect to water. One such peak wavelength for urea (or urea nitrogen) is at 2190 nanometers (nm) and one such reference wavelength with respect to water may be 1300 nm wavelength of light. Other wavelengths of significant absorption due to water (the reference) and minimal absorption due to urea or urea nitrogen are also present at 1480 nm, 1550 nm, 1800 nm, 1900 nm, etc. 
     In the presence of blood however, at 2190 nm, hemoglobin (or hematocrit) absorption (and scattering) also exists, therefore one must compensate the 2190 nm absorption value with the hematocrit (or hemoglobin) value. However, at the 1300 nm wavelength, the absorbency due to hematocrit or hemoglobin is minimal. 
     MEASURING OSMOLARITY AND SODIUM CONTENT 
     The functional relationships between Na +  and osmolarity (OSM) are well known to those skilled in the art, wherein Na +  variations cause the greatest variations in OSM. The modified Beer-Lambert equation (2) can be utilized to determine osmolarity and Na + . Further, it is also well known by those skilled in the art that variations in OSM or Na +  cause a direct change in the mean cell volume (MCV) of red blood cells (RBCs). When [Na + ] increases by 12 millequivalents/liter (meq/L), the microcentrifuge derived hematocrit decreases by one Hematocrit unit (or the MCV decreases by ˜2-3%). The proposed method takes advantage of the fact that the log (8)/log (13) ratio is insensitive to [Na + ] changes, whereas the log (8) alone, see FIGS. 7,  8  and  9 , is very sensitive to Na +  or MCV changes. Log (8) is equal to log (I/I o ) at the 810 nm wavelength and log (13) is equal to log (I/I o ) at the 1300 nm wavelength. 
     The actual function of E versus the OSM or Na +  concentration at each given wavelength must hold OSM (or Na + ) information that is different in at least one of curvature, offset, linearity, or sign from the other wavelength, see FIG.  7 . If the functions of E versus OSM are not sufficiently different, then the ratio E 1 /E 2  for the two wavelengths will not hold OSM information. FIG.  7 . shows the direct affect of Na +  on the optical absorbance (% change in absorbance), for a single wavelength (log (8)) and a dual wavelength device          (       log                   (   8   )         log                   (   13   )         )     .                          
     For a dual wavelength device, the affect of Na +  on the ratio of log (8)/log (13) is minimal. That is, the ratiometric use of two wavelengths cancels competing factors such as Na + . 
     FIGS. 8 and 9 show that either Na +  or osmolar changes in blood affect the mean cell volume of a red blood cell. In the present invention by measuring one or two appropriate wavelengths a difference will be measured as a function of mean cell value and thus of Na +  or osmolarity. Please note that the lines plotted in FIGS. 8 and 9 should be linear. They are not because the graphs represent actual experimental data measured with the present invention. 
     Even though wavelengths of 810 nm and 1300 nm are preferred wavelengths, the wavelengths may be selected further apart from one another such that the pathlengthening factors are exaggerated for each wavelength. Therefore, a shorter and longer wavelength are preferred since they exhibit even more sensitivity to scattering. The selection of the measuring wavelength at 585 nm and the reference wavelength at 1550 nm may be more appropriate since the scattering functions (pathlengthening factors) are exaggerated at these two selected wavelengths. 
     In non-pulsatile applications such as hemodialysis or wherever blood is flowing through a blood chamber or cuvette, then logarithmic operations will give the appropriate computed OSM or Na +  values, see FIGS. 7,  8  and  9 , and as shown in the following: 
     
       
         [ Na   +   ]=[log  (8)]• F[log  (8)/ log  (13)]  (15)  
       
     
     where F( ) is a function of hematocrit, and 
     
       
           OSM=g [Na   +   ]+b    (16)  
       
     
     where OSM is a function of Na +  and g [Na +]  is a function of Na + . Further, g is slope and b is offset. Both g and b are empirically determined using known methods that employ a look-up table. 
     MEASURING PLASMA FREE HEMOGLOBIN 
     The modified Beer-Lambert equation (2) can also be utilized to determine PFH. In the present invention PFH is determined by using an optical technique that does not distinguish between hemoglobin in red blood cells and hemoglobin in plasma. Rather, when light at 800 nm is shined through blood, each of the elements (red blood cells, plasma and hemoglobin) extinguish a certain amount of light energy as shown by the formula:                E   bulk     =       E   RBC     +     E   plasma     +     E     Hgb                 in                 plasma       +     E       H   2        O       +     E     other                 …                 (   17   )                              =       (       S   RBC     +     K   RBC       )     +     K   p     +     K     Hgb                 in                 plasma       +     K       H   2        O       +     K   other               (   18   )                                
     where S is a scattering coefficient and K is an absorption coefficient              =         (       S   RBC     +     K   RBC     +     K   Hgb       )                                          membrane   itself             inside     of                 RBC               +     K   p     +     K   Hgb     +     K       H   2        O       +     K     other                 …                 (   19   )                 Thus:                                        E   bulk     =                    (       S   RBC     +     K   RBC     +     K   Hgb       )                                          membrane   itself             inside     of                 RBC               +                                (       K     p                 of                 plasma       +     K     Hgb                 in                 plasma         )     +     K       H   2        O       +     K     other                 …                         (   20   )                                
     From this it can be seen that if there is no hemoglobin in plasma, then K Hgb =0. Otherwise, K Hgb  adds to the total hemoglobin in plasma and red blood cells. 
     As an example, at 800 nm wavelength:          S   RBC     =   2.5               K     RBC                  membrane     =   .05                 K   Hgb                    inside   membrane       =   .5             K   plasma     =   .01             K       H   2        O       =   .01             K       Hgb                  in                 plasma         =   .5                   E   bulk     =     2.5   +   .05   +   .5   +   .01   +   .5   +   .01                 =     3.57                 with                 Hgb                 in                 plasma                 =     3.07                 with                 no                 Hgb                 in                 plasma                                  
     FIG. 10 shows plots of the % change in absorbance of light at 810 nm and 810 nm/1300 nm versus PFH. In this way, FIG. 10 illustrates that using only a single wavelength (800 nm) produces a large % change due to PFH. On the other hand, the ratio of two wavelengths (log (8)/log (13)) nulls out the effects of PFH. 
     It should be noticed that the following assumptions and requirements are essential in PFH determination. 
     A. Even though wavelengths of 810 nm and 1300 nm are the preferred wavelengths, the actual function of E versus the PFH concentration at each given wavelength must hold PFH and hematocrit information that is different in curvature, or offset, or linearity, or sign from the other wavelength, see FIG.  10 . If the functions of E versus PFH are not sufficiently different, then the ratio E1/E2 for the two wavelengths will not hold PFH information. It will be appreciated that other wavelengths such as 585 nm and 1550 nm would also satisfy the condition of having adequate PFH detected with respect to water. 
     B. Further, the wavelengths may be selected further apart from one another such that the path-lengthening factors are exaggerated for each wavelength. Therefore, a shorter and longer wavelength are preferred since they exhibit more sensitivity to absorption and scattering. The selection of the measuring wavelength at 585 nm and the reference wavelength at 1900 nm may be more appropriate since the scattering functions (path-lengthening factors) are exaggerated at those two selected wavelengths. 
     C. FIG. 10 shows that as PFH varies the log (8)/log (13) ratio is unaffected. Whereas, as PFH varies the single wavelength alone, log (8), varies greatly. In other words, the absorption effects due to hemoglobin, whether inside the red blood cell or in the plasma itself, are seen by the detector as a bulk absorbance. When two distinct wavelengths are used, each wavelength will carry, individually, bulk absorbance values. When these two wavelengths are mathematically operated upon ratiometrically the PFH is minimized, see FIG.  10 . 
     D. If monitoring the PFH in living tissue of a patient, the pulsatile characteristics of the blood would require the utilization of the form and mathematical operations as presented in U.S. Pat. No. 5,372,136, using the ΔI/I technique in order to eliminate certain intrinsic tissue and extrinsic light source effects. 
     E. In non-pulsatile applications such as hemodialysis or wherever blood is flowing through a disposable conduit, disposable blood chamber, or cuvette; then, logarithmic operations will give the appropriate computed PFH values, see FIG.  10  and as shown in the following: 
     
       
         [ PFH]=A[log  (8)• F ( log  (8)/ log  (13))]+ B    (21)  
       
     
     where A is slope and B is offset. Both A and B are empirically determined using known methods that employ a look-up table. 
     In this way, it can be appreciated that the present invention is directed toward apparatus and methods for determining the biologic constituent value of the PFH noninvasively. This is achieved by passing at least two wavelengths of light onto or through body tissue such as the finger, earlobe, or scalp or through a disposable extracorporeal conduit for the noninvasive monitoring. Compensating for the effects of competing body tissues and/or the variations in hematocrit is achieved with additional mathematical operations. 
     The present invention may be embodied in other specific forms without departing from its spirit or essential characteristics. The described embodiments are to be considered in all respects only as illustrative and non-restrictive. The scope of the invention is, therefore, indicated by the appended claims rather than by the foregoing description. All changes which come within the meaning and range of equivalency of the claims are to be embraced within their scope.