Abstract:
A planar dielectric waveguide sensor is used to determine the concentrations of oxygenated and deoxygenated hemoglobin and other blood constituents such as pH, sugar. The planar waveguide core is in direct contact with the blood such that evanescent field of the light propagating within the core is selectively attenuated at specific wavelengths of interest. The planar waveguide has a construction that promotes strong interaction of the evanescent field with blood cells that contact it. In preferred embodiments, the waveguide is constructed of a low refractive index core to propagate an evanescent wave comparable in size to a red blood cell.

Description:
CROSS REFERENCE TO RELATED APPLICATIONS  
       [0001]     The present application claims priority to the U.S. Provisional Patent Application for a “Blood Oxygenation sensor, filed on Feb. 21, 2006, and now assigned application Ser. No. 60/775,531, which is incorporated herein by reference. 
     
    
     BACKGROUND OF INVENTION  
       [0002]     The present invention relates to methods and an apparatus for the in-vivo optical measurement of blood chemistry, and in particular the oxygenation of red blood cells.  
         [0003]     Prior art methods of measuring the oxygenation of blood are well known and utilize various means to measure light absorption at specific wavelengths in the Near Infrared (NIR) of the spectrum of electromagnetic radiation to distinguish between the concentration of the oxygenated hemoglobin and de-oxygenated or reduced hemoglobin in red blood cells (RBC&#39;s). Externally worn sensors, typically deployed on a finger of other thin external appendage, are widely deployed but give a gross average of the oxygenation of RBC&#39;s in the large number of veins and artery that are illuminated by an external light source. Of more interest to researchers and clinicians are local arterial measurements at a specific location.  
         [0004]     U.S. Pat. No. 5,280,786 to Wlodarczyk et al. issued on Jan. 25, 1994 for an Fiberoptic blood pressure and oxygenation sensor deployed on a catheter placed transcutaneously into a blood vessel. A sensing tip of the catheter includes a pressure-sensing element and an oxygen saturation-measuring element. One of the disclosed means for measuring the oxygen saturation of blood was a sensing tip comprising an optical fiber having at least a portion of the cladding removed. Other features of this invention included novel tip designs, measuring head features, and approaches for enhancing measurement though correlation of the saturation and pressure readings.  
         [0005]     While the device disclosed in the &#39;786 patent was an improvement of the prior art in providing a means to measure oxygenation in a single vein or artery it appears to still suffer several limitations of potential consequence to more extensive an long term clinical deployment in patents.  
         [0006]     One limitation recognized by the inventors in the current invention is that the geometry created by the removal of the fiber cladding and to enable the other features is that it creates an irregular protruding device in the blood stream. As such, a device is placed in smaller arteries and veins, where the flow velocity is higher; there is a greater likelihood that the protrusion will result in turbulent blood flow that can disrupt RBC&#39;s leading to hemolysis.  
         [0007]     For such reasons, and others it would be desirable to have a fiber optic sensor such as that disclosed in the &#39;786 patent that is at least one of more sensitive and smaller in size, and preferably both and does not suffer from the disadvantages of prior art.  
         [0008]     It is therefore a first object of the present invention to provide an optical means for precise local measurement of oxygenated hemoglobin and reduced hemoglobin that can potentially be smaller than the size provided by the prior art.  
         [0009]     It is a further objective of the present invention to provide an optical sensing means that is sufficiently compatible with blood that it can remain in a patient for a long period of time, and be deployed in smaller veins and/or arteries.  
         [0010]     It is yet another object of the present invention to provide such a sensor device that is easier to integrate with other biomedical devices and transducers  
         [0011]     It is also a further objective of the present invention to provide such a device that is capable of a far more accurate local and representative determination of concentrations of in-vivo blood components, including oxygenated hemoglobin and reduced hemoglobin.  
       SUMMARY OF INVENTION  
       [0012]     In the present invention, the first object of having a reduced size device that is non-protruding is achieved by providing in optical communication a laser or other light source, a beam splitter and a first detector to receive a portion of the energy directed to it by the beam splitter. The other portion of the energy directed by a beam splitter is transported to a planar waveguide in contact with the blood. Light from the laser or other source may be delivered to the planar waveguide by an optical fiber. The light attenuated by absorption of the evanescent wave in the planar waveguide is directed to a second detector. When the light is delivered by an optical fiber, the light direction means is preferably a mirror at the end of the optical fiber.  
         [0013]     A second object of the invention of providing higher sensitivity and more representative measurement of the local concentration of oxygenated and reduced hemoglobin is achieved by providing either a planar or an optical fiber waveguide sensing device in contact with the blood wherein the refractive index of the core of the waveguide is preferably less than about 1.47. The refractive index of the waveguide core exposed to the blood is more preferably between about 1.38-1.42, the refractive index of the human RBC&#39;s, to allow the evanescent light to penetrate deeply into more RBC&#39;s than prior art devices.  
         [0014]     The above and other objects, effects, features, and advantages of the present invention will become more apparent from the following description of the embodiments thereof taken in conjunction with the accompanying drawings.  
     
    
     BRIEF DESCRIPTION OF DRAWINGS  
       [0015]      FIG. 1A  a schematic illustration of the planar dielectric waveguide with the sandwiched between two cladding regions in which the guided waves propagate in the z direction.  
         [0016]      FIG. 1B  illustrates the electric field distribution  15  within single mode dielectric waveguide  10  of  FIG. 1A  and as it spreads into the cladding  11  where its strength is exponentially decaying in intensity.  
         [0017]      FIG. 2  is a graphic representation showing the dependence of the imaginary part of the effective refractive index n eff ″ on the imaginary part of the cladding refractive index n 2 ″.  
         [0018]      FIG. 3  is a graphic representation of showing the dependence of γ′, the exponential decay factor in the cladding from equation 9, on the imaginary part of the cladding refractive index n 2 ″.  
         [0019]      FIG. 4  shows the absorption spectrum of water in the 600-1100 nm range.  
         [0020]      FIG. 5  compares the absorption spectra of oxygenated and reduced hemoglobin in the ranges 450-1000 nm (top), and 650-1050 nm (bottom) as published by Cope 1991 (M. Cope. Ph.D. Thesis, Univ. College London, 1991)  
         [0021]      FIG. 6A  is a schematic illustration of the device deployed in the method for measuring concentrations of Hb and HbO 2 .  
         [0022]      FIG. 6B  is a cross sectional elevation through the schematic oximetry probe portion of the device in  FIG. 6A .  
         [0023]      FIG. 6C  is a cross section elevation of the oximetry probe portion of the device in  FIGS. 6A and 6B  showing the optical fiber, coupler and planar waveguide at section line C-C in  FIG. 6B   
         [0024]      FIG. 7  illustrate the dependence of β″ on saturation level SO 2    
         [0025]      FIG. 8  illustrates the power extinction (Eq. 16) as a function of the saturation level in a Log-Linear scale respectively to better show that the extinction increases logarithmically with the saturation level  
         [0026]      FIG. 9  illustrates in perspective the geometry of the dielectric waveguide  130  used for measuring oxygen saturation level of blood.  
         [0027]      FIG. 10  illustrates the single mode dielectric waveguide electric field distribution for the dielectric waveguide  130  of  FIG. 9   
         [0028]      FIG. 11  is a graph illustrating the refractive index of hemoglobin and water in the range 250-1050 nm.  
         [0029]      FIG. 12  compares the dispersion of the refractive index of the HOSP in the infrared, visible and ultraviolet ranges with that of SiO2. Note that the wavelength is indirectly shown as the energy in eV.  
         [0030]      FIG. 13A  is an illustration of the geometry of the planar dielectric waveguide consisting of the core  130  and cladding  135  as surrounded by water  5  in which a single red blood cell particle  200  placed on the waveguide surface.  
         [0031]      FIG. 13B  illustrates the results of modeling the electric field intensity distribution for the geometry shown in  FIG. 13A . The field propagates from the left to the right of the waveguide  130 .  
     
    
     DETAILED DESCRIPTION  
       [0032]     Referring to  FIGS. 1 through 13 , wherein like reference numerals refer to like components in the various views, there is illustrated therein a new and improved blood oxygenation sensor, generally denominated  100  herein.  
         [0033]     Before providing details on the construction and methods of using the various embodiments of the invention, theoretical consideration regarding the operative principles of such device will be described.  
         [0034]     A step index optical fiber/dielectric waveguide has, basically, two regions with different refractive indexes: core and cladding. The core has a refractive index larger than the cladding region surrounding it. Light waves are guided throughout the waveguide due to a phenomenon known as total internal reflection. Remarkably, that propagating field is not spatially limited only to the core, but also extends into the cladding and exponentially decays with the distance from the core. This exponentially decaying field is known as an evanescent field. The extension of the evanescent field depends on the core diameter and the refractive indexes of the core and the cladding.  
         [0035]     The absorption coefficient of silicate glass is very small in the visible and infrared ranges of the spectrum; because of that, the light absorption in the waveguide is usually neglected in the analysis of guiding modes. In contrast to the textbook analysis, in the treatment below, we consider a waveguide with a transparent core and an absorbing cladding media.  
         [0036]     The propagation of light within dielectric medium is governed by the Helmholtz equation:
 
Δ E+n   2   k   0   2   E= 0  (1)
 
 where E=E(x, y, z) is the electric field, n is the refractive index and k 0 =ω/c=2π/λ is a free-space light wavevector. 
 
         [0037]     The Helmholtz equation remains valid also in the case of an absorbing medium. In that case the refractive index is a complex quantity:
 
 n=n′+in″   (2)
 
 where n″ describes light absorption:
 
α=2k 0 n″  (3)
 
 where α is the absorption coefficient. 
 
         [0038]     For the sake of simplicity, we consider only a symmetric planar dielectric waveguide  10  as schematically illustrated in  FIG. 1A . The planar dielectric waveguide core  9  is sandwiched between two cladding regions  11 ′ and  11  in which the guided waves propagate in the z direction.  
         [0039]     We consider only the case of the transverse electric field. The case of the transverse magnetic field could be treated in the analogously. Seeking the solution of the Eq. (1) in the form E(x, y, z)=ŷE y (x)exp(iβz), we obtain the following equations for the propagating field  
                       ∂   2     ⁢     E   y         ∂     x   2         +       (         n   j   2     ⁢     k   0   2       -     β   2       )     ⁢     E   y         =   0     ,           (   4   )             
 
 where β is the propagation constant, n j , j={1, 2} is the refractive indexes of the core and the cladding respectively. Since the refractive index of cladding is a complex quantity the same is also true for β:
 
β=β′+ iβ″.   (5)
 
         [0040]     The imaginary part of β is responsible for decaying of the propagating wave. The intensity of the wave after propagation of a distance L is:
 
 I=I   0  exp(−2β″ L ),  (6)
 
 where I 0  is the input intensity. This intensity could be measured by a photodetector. Our goal is to relate β″ with the absorption coefficient of the cladding medium α=2k 0 n 2 ″. Alternatively, defining the effective refractive index n eff  as
 
β=k 0 n eff ,  (7)
 
 the above problem could be reformulated as finding a relation between n 2 ″ and n eff ″. The general solution of the Eqs. (2) has the form
 
 E   y1   =A  exp( ik   1   x )+ B  exp(− ik   1   x )
 
 E   y2   =C  exp(−γ x )  (8)
 
 where E y1 , E y2  are the fields in the core and the cladding respectively and k 1  and γ are the complex numbers given by
 
 k   1   =√{square root over (n 1   2 k 0   2 −β 2 )} 
 
γ=√{square root over (β 2   −n   2   2   k   0   2 )}  (9)
 
         [0041]     E y2  is the evanescent field. It extends into the cladding on the effective distance 1/Re[γ]. The constants k 1  and γ are calculated from continuity conditions of the field and its derivatives on the interface. Specifically, for the even modes: E y (x)=E y (−x), the following equation holds:
 
 k   t  tan( k   1   a )=γ  (10)
 
 wherein for the odd modes: E y (x)=−E y (−x) the following equation holds:
 
 k   1  cot( k   1   a )=−γ  (11)
 
 where a =d/2 is the half of the core thickness. From Eqs (9) follows:
 
 k   1   2 +γ 2   =k   0   2   NA   2 ,  (12)
 
 where NA is the numerical aperture of the waveguide NA=√{square root over (n 1   2 −n 2   2 )}. The transcendental equations Eqs. (10)-(11) together with Eq. 12 solve the above problem. The complex propagation constant β is calculated from Eqs (9):
 
β=√{square root over (n 1   2 k 0   2 −k 1   2 )}=√{square root over (n 2   2 k 0   2 +γ 2 )}  (13)
 
 If there is no absorption, then all quantities in Eqs. (9)-(13) are real numbers. In that case, the above equations coincide with the classical equations for modal dispersion in a planar symmetric dielectric waveguide. However, when the propagating medium is absorbing, the same equations solve the problem of finding dependence of the mode intensity decay rate 2β″ on the absorption coefficient of the cladding medium α. 
 
         [0042]      FIG. 1B  shows a single mode field intensity  15  in a non-absorbing dielectric waveguide  10  with core  9  and cladding layers  11 ′ and  11  above and below respectively. The field was calculated from Eqs. (7), (9) and (10) for a waveguide with a core thickness of d=2 μm, a light wavelength of λ=640 nm, a numerical aperture NA=0.11 and a core refractive index 1.45 (silica glass).  
         [0043]     The dependence of the decaying rate 2β″=2k 0 n eff ″ inside the dielectric waveguide in the example above on the absorption coefficient of cladding medium α=2k 0 n 2 ″ is shown on the  FIG. 3 .  
         [0044]      FIG. 2  shows, that for the values of n 2 ″ up to approximately 0.005 or, equivalently, up to absorption coefficient α=2k 0 n 2 ″=0.1 μm −1  the effective refractive index is approximately proportional to n 2 ″. The maximum of n eff ″ is at n 2 ″=0.02; after that value the decay rate decreases.  
         [0045]     The dependence of γ′ on n 2 ″ is shown on the  FIG. 3 . As seen from the plot γ′ is remains constant up to n 2 ″≈10 −3 , after that γ′ increases rapidly. That corresponds to the decrease of the penetration depth of the evanescent field (which is 1/γ′) in a strongly absorbing medium.  
         [0046]     The main constituents of blood which contribute towards absorption in visible and near infrared ranges are water and hemoglobin. While the former is constant, the concentrations of oxygenated hemoglobin (HbO 2 ) and reduced hemoglobin (Hb) change. Thus, the corresponding changes in absorption can provide clinically useful physiological information.  
         [0047]      FIG. 4  shows absorption spectrum of water in the 600-1100 nm range. The significant absorption occurs only in the infrared region. The spectra of Hb and HbO2, expressed in term of the specific extinction coefficient, can be seen in  FIG. 5 . While both absorb strongly in the blue and green regions of the visible spectrum; the absorption of Hb is slightly stronger beyond about 590 nm. Note the point at about 800 nm, where the two curves intersect. The specific extinction coefficient μ represents the level of absorption per mmol of compound per liter of solution per cm (usually quoted in unit&#39;s mmolar-1 cm-1). It is related to the absorption coefficient α as:  
             μ   =     k   ⁢           ⁢     α   c               (   14   )               
 where k=log 10 (e)=0.4343 is a constant and c is the concentration of the compound (in units of mmoles). 
 
         [0048]     In accordance with a first embodiment of the invention, an oximeter device  100  for measuring concentrations of oxygenated or oxyhemoglobin (HbO 2 ) and reduced hemoglobin (Hb) is shown on  FIG. 6 , which shows that the various useful absorption bands of Hb and HbO 2  occur at wavelengths between about 550 nm and about 800 nm.  
         [0049]     Light from a laser source  105  is coupled into a 50/50 2×2 fiber optic beamsplitter  115 . Half of the light energy is coupled into an optical fiber  110  that ends with the oximetry probe  120  while the other half is measured on the detector D 2  ( 122 ) which measures a reference signal.  
         [0050]     In the optical fiber  110  the light propagates without losses until it enters the planar waveguide portion  130  exposed to the blood. Light enters planar waveguide portion  130  from fiber core  109  via a coupler  140 , as shown by the solid arrows. In this region, the intensity of the wave exponentially decreases with the propagating distance due to interaction of the wave with the absorbing medium (blood) via its evanescent field. When the waves reach the end of the planar waveguide  130  or the fiber  120  they are reflected from the mirror on the fiber&#39;s end, as shown by the dashed arrows. An optical fiber  110  is the preferred means for delivery light from the laser source  105 , as it can be readily adapted to fit into or form a catheter that is inserted into the body, and in particular the cardiovascular system. As an alternative to using coupler  140  to direct a portion of the light from optical fiber  110  to planar waveguide  130 , optical fiber  110  may simply terminate in a planar waveguide having a mirror on the end face, or other means to return light back in the direction of the optical fiber indicated by the dashed arrows. It is also preferable that the light transmitted through the internal fiber but not coupled into the planar waveguide should be absorbed, rather than reflected by the mirror  150 . Thus, it is more preferable that only the planar waveguide terminates in a mirror, with the end of the fiber optic terminating in an absorbing layer  180  or alternative structure or optical path that does not allow uncoupled light to reflect back to the detector.  
         [0051]     Although the preferred embodiment deploys a mirror at the end of the fiber optic  110  used to deliver light to the planar waveguide  130 , alternative embodiments include using a continuous optical fiber in the form of a loop that terminates at detector D 2  ( 122 ) wherein a second optical coupler would transmit light from the planar waveguide in the same direction as propagation such that it reaches detector D 2 .  
         [0052]     It should also be appreciated that rather than using a laser, a multi-wavelength light source might be deployed such as a broadband light source of multiple fiber optic lasers each tuned to a different wavelength.  
         [0053]     The backreflected wave is transmitted through the exposed region and finally, after splitting on the beamsplitter  115  reaches the detector D 1  ( 121 ). The light intensity decay in one pass of the exposed region is:
 
 I=I   0  exp(−2β″  L ),  (15)
 
 where β″ is the imaginary part of β in the Eq. (13) and L is the length of the exposed region. If the transmission losses of the optical fiber and losses on the mirror and beamsplitter are negligible in comparison with the losses in the exposed region, then the following equation holds:
 
exp(−4β ″ L )=2 I   1   /I   2   (16)
 
 where I 1 , I 2  are the signals on detectors D 1  and D 2  respectively. The factor 4 in the exponent in Eq.(16) appeared because the double pass of the exposed region. From Eq. (16) the characteristic distance L 0  of the exposed region is:  
               L   0     =       1     4   ⁢           ⁢     β   ″         .             (   17   )             
 
         [0054]     At this distance the intensity decreases e times in a double pass. Although a more precise model should account for reflections at the interfaces between the cladding of the optical fiber and absorbing medium (blood) as well the losses due to fiber mode transformation when the mode field enters into the exposed region, we expect comparable trends to those shown herein.  
         [0055]     The β″ can be expressed through the absorption coefficient of blood α by the solution of the modal dispersion equations.  
         [0056]     The absorption coefficient of the blood depends on the concentrations of oxygenated and deoxygenated hemoglobin in it. The total absorption coefficient in the blood is a sum of specific absorption coefficients:
 
α=α 0 +α Hb +α HbO     2     (18)
 
 where α HbO     2   , α Hb  and is the absorption coefficients of oxygenated and deoxygenated hemoglobin and α 0  is the absorption coefficient of blood without hemoglobin. α 0  is assumed to be constant on a short time scale on which α HbO     2   , α Hb  changes. Using Eq. (14), Eq. (18) could be rewritten in the form:  
             α   =       α   0     +         μ   Hb     k     ⁢     c   Hb       +         μ     Hb   ⁢           ⁢     O   2         k     ⁢     C     Hb   ⁢           ⁢     O   2                     (   19   )             
 
 where c Hb , c HbO     2    are concentrations of the Hb and HbO 2  respectively. Note that the absorption coefficient and specific extinction coefficients are functions of λ. If a measurement is done at two or more wavelengths of light, then we have the system of linear equations  
               α   ⁡     (     λ   i     )       =         α   0     ⁡     (     λ   i     )       +           ⁢           μ   Hb     ⁡     (     λ   i     )       k     ⁢           ⁢     c   Hb       +           μ     Hb   ⁢           ⁢     O   2         ⁡     (     λ   i     )       k     ⁢           ⁢     C     Hb   ⁢           ⁢     O   2                     (   20   )             
 
 where λ i  {i=1, . . . , N} are wavelengths of the light. If the function α 0 (λ) is known, then the system of Eqs. (20) could be solved for concentrations of the Hb and HbO 2 . This system is over determined for N&gt;2; it has a single solution at N=2 if the determinate of the system is not zero. If N=1, then the system is under determined and could not be solved. However, if the total concentration of hemoglobin c in blood is known, then using the additional equation:
 
 c   Hb   +c   HbO     2     =c,   (21)
 
         [0057]     The concentrations c Hb , c HbO     2    could be determined as well for the case N=1.  
         [0058]     To better appreciate certain aspects and embodiments of the present invention we first illustrate the optimum and effective operative parameter for measuring blood oxygenation with a fiber optic sensor operating at the single wavelength λ=600 nm. That is, rather than utilizing the coupler and dielectric waveguide shown in  FIG. 6 , we consider the performance when the oximeter probe is simply the terminal portion of the optical fiber when the cladding is removed or reduced in thickness such that an evanescent field might extend into the blood. As in a conventional optical fiber, the refractive index of the fiber core is n 1 =1.45 and the core diameter is 2 μm. Blood is modeled as a medium with the refractive index n 2 =1.33+iα/2k 0 , where α is the blood absorption coefficient given by Eq. (19). The concentration of hemoglobin c is assumed to be 15 g/100 ml, or equivalently, 2.32×10 3  μmolar.  
         [0059]      FIG. 7  shows imaginary part of β calculated from Eq. (10), Eq. (12) and Eq. (13) as a function of saturation level SO 2  that is defined as the ratio between c HbO     2    and c Hb . From this figure, we see that β″ changes linearly as the saturation level of O 2  changes from 70% to 100%. The characteristic distance L 0  can also be modeled from Eq. (16) at the corresponding minimum of β″=0.06 (100% O 2)  saturation is 4.2 cm. Therefore, in this example an optimal length of the exposed region should be about 4 cm.  FIG. 8  shows power extinction (Eq. 16) as a function of the saturation level. The extinction increases logarithmically with the saturation level.  
         [0060]     We now consider in more detail the particular advantages of the embodiment of the invention utilizing a planar dielectric waveguide  130  shown in device  100  of  FIG. 6A -C. The geometry of this dielectric waveguide used for measuring oxygen saturation level of blood is shown in  FIG. 9 , waveguide  130  has a substrate  132  that acts as a lower cladding and a rectangular core  135 . In calculating the single mode dielectric waveguide electric field distribution by finite element methods for the dielectric waveguide  130  of  FIG. 9 , which is plotted in  FIG. 10  the refractive index of the waveguide core  135  was assumed to be n 1 =1.45 while the refractive index of the substrate is 1.40 for measuring the performance at the wavelength of λ=600 nm.  FIG. 10  has a shaded legend bar at the right side in which dashed lead lines connect the proper shaded portion of the bar to the corresponding shaded regions of the waveguide  130  to show the gradation in the evanescent field. The model of the device performance assumes the core region  135  of waveguide  130  is uncovered and is in the direct contact with blood. The saturation level of the blood is assumed to be 97%. Applying finite element methods to calculate mode field of the waveguide  130 , shown in  FIG. 11 , also resulted in the imaginary part of the effective refractive index being 1.88×10 −6 .  
         [0061]     Applying Eq. 16 and 17, this corresponds to the β″=0.2 cm −1  and to the effective waveguide length L=1.25 cm, which is about a third less than the optimum length (4 cm) of the exposed region for a circular optical fiber. Thus, deployment of the planar waveguide permits the portion exposed to the blood to be less than 3.5 cm.  
         [0062]     Another important advantage of the novel planar geometry sensor compared with the prior art disclosure of optical fiber sensors is the much lower potential for error in blood saturation measurement. Sources of potential errors include variations of the optical signal of the laser and inside the optical fiber and photodetector noises. This can be modeled by letting x=c HbO     2   /c be the saturation level. Then variation of x from Eq. (19) and Eq. (21) is  
               δ   ⁢           ⁢   x     =       0.43     c   ⁢           ⁢   Δ   ⁢           ⁢   μ       ⁢   δα             (   22   )             
 
 where Δμ=μ HbO     2   −μ Hb  and δα is the uncertainty in the measurement the level of blood absorption coefficient. For n 2 ″&lt;&lt;1α is proportional to β″ with a good accuracy. Writing β″=Γα, where Γ is a constant for a given fiber we could rewrite Eq. (22) in the form  
               δ   ⁢           ⁢   x     ⁢           =       0.43     c   ⁢           ⁢   ΓΔ   ⁢           ⁢   μ       ⁢           ⁢       δβ   ″     .               (   23   )             
 
         [0063]     The error in the determination of the saturation level from the intensities I 1 , I 2  is from Eq. (16) and Eq. (23)  
                 δ   ⁢           ⁢   x     ⁢           =     0.11   ⁢     ξ     Γ   ⁢           ⁢   Lc   ⁢           ⁢   Δ   ⁢           ⁢   μ           ⁢           ,           (   24   )             
 
 where ξ=δI 1 /I 1 −δI 2 /I 2  is the difference of intensity variations on the two photodetectors. 
 
         [0064]     From Eq. (24) we see that δx becomes smaller with the increase of Δμ, L and Γ. In the equation above Δμ is determined by the wavelength of the light and Γ depends on the configuration of the fiber/waveguide.  
         [0065]     Thus, now applying Eq. 24 to the conventional fiber optic sensor considered above, with the exposed length, L being fixed at 4 cm for maximum saturation: Γ=3.1×10 −3 , cΔμ=2.32 cm −1  and δx=2.7ξ. Thus, if for example ξ=0.01, then the error in the saturation level is about 3%.  
         [0066]     In the example of the inventive planar dielectric waveguide  130  of device  100  in  FIG. 6 , the exposed length, L at maximum saturation is 1.25 cm: Γ=0.019 and δx=1.4ξ. For the same level of intensity variations, the saturation error level is 1.4%, or less than half the error (3%) of the fiber optic sensor.  
         [0067]     Thus, another aspect of the invention involves the method of first providing a waveguide comprising a planar support as a cladding on a first surface with a second surface parallel to the plane of the first surface, and terminating with a reflective surface orthogonal to the direction of propagation, then placing the second surface in contact with blood and propagating light through the waveguide toward the mirror, after which the intensity of light reflected by the mirror is measured.  
         [0068]     Yet another important operative principle of an even more preferred embodiment of the current invention for measuring blood oxygen saturation level is to deploy a waveguide in the which the dimensions of the evanescent field is comparable to, and most preferably, much more than the dimensions of the red blood cell. It should be appreciated that if the evanescent field that interacts with the RBC is much smaller than a RBC the signal will be strongly influenced by position of a particular red blood cell relative to the waveguide. The blood component hemoglobin is concentrated within erythrocytes or red blood cells that have a torus-like shape with the diameter of each corpuscular being is about 8 μm and having a thickness of about 2 μm. Further, the evanescent field should be of a nature that allows it to also penetrate deeply into the red blood cell, that is at least a micron, or preferably at least about 2 microns, but more preferably about 4 microns. These two conditions are fulfilled for a low-dielectric-constant (low-k) planar dielectric waveguide described below.  
         [0069]     The spatial extension of evanescent field is proportional to λ/Δn, where Δn is the difference of refractive index between core and the cladding. The refractive index of the human red blood cells is in the range 1.38-1.42 depending on concentration of the hemoglobin in it, as shown in  FIG. 11 . Therefore, the refractive index of the core of the waveguide should be as close to these values as possible to allow the evanescent light penetrate deeply into the red blood cell. The refractive index of SiO 2  waveguide is about 1.45. Thus, the typical penetration depth at that difference in refractive indexes between silica waveguide and red blood cell is 200-300 nm, or about between a tenth and a sixth of the thickness (2 μm) of the red blood cell. Accordingly, it is preferable that the waveguide  130  has a refractive index (n) is less than 1.45 at the absorption bands of Hb and HbO 2 . It is more preferable that the refractive index of the waveguide be in the range of 1.38-1.45.  
         [0070]     Thus, another embodiment of the invention is use of a dielectric material for a waveguide with refractive index lower than of silica to increase the penetration depth of the evanescent field, and thus obtain both a greater and more representative measurement of the blood oxygenation. One such preferred low-dielectric-constant material is spin-on hybrid siloxane-organic polymer, such as that known as HOSP and available from Honeywell Advanced Microelectronic Materials (Tempe, Ariz.). Thin films of HOSP could be prepared by a spin-on coating technique. The dispersion of refractive index of HOSP in wavelength range of 180 nm to 2.35 μm is shown in the  FIG. 12 .  
         [0071]      FIG. 13B  illustrates the results of modeling the electric field intensity distribution for the geometry shown in  FIG. 13A . The field propagates from the left to the right of the waveguide  130 . The planar waveguide modeled as having been fabricated by deposition of a film of the HOSP material as the core  130  on a silicon substrate that serves as one side of the cladding. The other side of the cladding  135  of the waveguide may be formed by ion implantation into the HOSP layer. The thickness of the waveguide core is 0.5 μm. The refractive index of the core  130  of the waveguide was 1.375 and the refractive index of the cladding was 1.36. The propagation of light in the planar waveguide was then modeled as if surrounded by blood that was a solution of erythrocytes in water with a refractive index, n, of 1.33. The complex refractive index of red blood cell was modeled as 1.38+iα/2k 0 , where α is the absorption coefficient. The model geometry is shown in the  FIG. 13A . This figure shows only a small part of the waveguide (L=40 μm) and a single particle on it. Preferably, the waveguide  130  has a length of at least about 10 microns. Such waveguides as modeled herein will result in an accurate determination of concentrations of blood components by sampling more blood cells and sampling each red blood cell to a greater depth. Further, as the planar waveguide can have a low measurement error with a small area, such a precise measurement of oxygenated and reduced hemoglobin can be made at a particular local as the oximeter probe is inserted in a catheter or other implanted medical devices.  
         [0072]     It should be appreciated from the foregoing that although a preferred form of a waveguide is a planar waveguide flush with the surface of the catheter or probe, the performance of an optical fiber (with a circular cross section, or any other non-planar waveguide form) is improved when the core refractive index is less about 1.45 at the absorption bands of Hb and HbO 2    
         [0073]     In other embodiments of the invention either the planar waveguide or a non-planar waveguide using a low refractive index core may be integrated with a catheter in signal communication with cardiac monitoring equipment, or a pacemaker or electro-cardiac defibrillator to change the pacing rate or provide a defibrillating pulse when local low blood concentrations are detected so as to prevent cardiac or other tissue ischemia.  
         [0074]     While the invention has been described in connection with a preferred embodiment, it is not intended to limit the scope of the invention to the particular form set forth, but on the contrary, it is intended to cover such alternatives, modifications, and equivalents as may be within the spirit and scope of the invention as defined by the appended claims.