Abstract:
Both the number and the size of microair bubbles in a bloodstream are accurately determined optically, independently of oxygen saturation, by monitoring the intensity of light transmission in the 800-850 nm range through the bloodstream and indicating the count and amplitude range of peaks in the monitored intensity.

Description:
This application is a continuation, of application Ser. No. 08/841,015, filed Apr. 29, 1997, now abandoned. 
    
    
     FIELD OF THE INVENTION 
     This invention relates to the detection and quantification of microair in blood, and more particularly to an infrared system for counting microair bubbles in a bloodstream and determining their size. 
     BACKGROUND OF THE INVENTION 
     During open-heart surgery, microscopic air bubbles having a diameter on the order of 60-300 μm are frequently entrained into the blood circuit of the heart-lung machine in spite of careful defoaming of the blood passing through the machine. These microair bubbles have been suspected of causing strokes, memory loss and other undesirable effects in the patient. It is therefore important in a heart-lung machine to detect the presence and size of these bubbles remaining in the bloodstream after filtration so that their origin can be traced and appropriate remedial measures can be taken when the presence of microair is detected in the blood circuit. 
     Microair detection in the prior art has conventionally been done by transmitting an ultrasound beam through a bloodstream flowing past the detector. The problem with this approach is that ultrasound is expensive and is neither able to accurately evaluate the size of individual bubbles, nor detect them as individual bubbles when they are close together. Also, ultrasound measures discontinuities in the bloodstream and therefore cannot distinguish between microair and tiny blood clots. Also, the accuracy of ultrasound measurements becomes poor for very small diameter bubbles. Furthermore, in pulsing ultrasound applications, the propagation velocity of sound required ultrasound pulses to be at least 10-20 μsec apart for a fast-moving 1.25 cm diameter bloodstream, so that tracking is not continuous. Consequently, a more accurate and discriminating method of detecting microair was needed. 
     Prior to the present invention, optical detection of microair was of limited use because it only operated in a binary (bubble present or absent) mode. Quantitative detection was considered impractical because light transmission through a bloodstream is strongly affected by hematocrit and oxygen saturation, which vary unpredictably during surgery. 
     SUMMARY OF THE INVENTION 
     The present invention allows highly accurate detection and measurement of microair over a wide range of bubble sizes by detecting the translucence of a bloodstream to infrared radiation having a wavelength of about 800-850 nm in at least two directions at a substantial angle to each other. The detection is made by detecting spikes in the amplitude of infrared signals received by an array of infrared sensors disposed around a column of blood when microair bubbles pass the field of view of the sensors. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 is a perspective view of a microair detector according to the invention; 
     FIG. 2 is a section along line  2 — 2  of FIG. 1; 
     FIG. 3 is a front view of an infrared emitter/receiver useful in the invention; 
     FIG. 4 is a wavelength-absorption diagram showing the light absorption coefficients of various media as a function of wavelength; 
     FIG. 5 illustrates an alternative embodiment of the invention; and 
     FIG. 6 is a time-amplitude diagram showing the amplitude pattern produced by the passage of microair bubbles through the detector. 
    
    
     DESCRIPTION OF THE PREFERRED EMBODIMENT 
     FIGS. 1-3 show an optical microair detector system  10  according to the invention. Blood  11  containing microair bubbles  13  flows through a rigid, transparent tube  12  preferably formed of polycarbonate. Rigidity of the tube  12  is necessary because the pulsations of the blood flow in a heart-lung machine expand and contract the walls of a flexible tube enough to create a rhythmic noise in the detector and degrade the signal quality. 
     Disposed around the tube  12  in a light-tight enclosure  14  clamped around the tube  12  by fasteners  15  are two or more opposed sets  16  of combination light sources  18  and photodetectors  20 . The light source  18  and photodetector  20  of each set  16  may be nested within each other as shown in FIG. 2, or disposed side by side as shown in FIG.  3 . The light sources  18  illuminate the opposing detectors  20  through the blood in the tube  12 . The number of sets  16  will depend upon the diameter of the tube  12  and the dimensions of the sets  16 , but they are preferably so disposed (FIG. 3) that the entire cross section of the tube  12  is either directly in the light path of a set  16 , or at least is substantially illuminated by the side scatter of a light beam from a set  16 . 
     In the above-described apparatus, the detection of intensity changes in light signals passed through the tube  12  is the key to detecting the presence of a bubble  13 . These changes in the light intensity can be caused either by highly obstructing particles (clots, bone chips, etc.) or by air emboli or bubbles  13  that reflect light at their surfaces and allow the remaining light to pass through without the absorptive effects of blood. It is possible to determine what kind of obstruction has passed through the detector, as well as the size of that obstruction, by tracking the resulting variations in the light signal impinging upon the photodetectors  20 . 
     The variations in the impinging light intensity caused by bubbles  13  are due to the difference between the optical properties of the blood  11  and the bubbles  13 . Light travelling through the blood  11  is both scattered and absorbed by the different component particles of blood, such as red blood cells, water molecules, and platelets. Light is absorbed both by the hemoglobin found in red blood cells and the water molecules. Scattering, where the light is deflected by some angle, generally results when the light interacts either with red blood cell bodies or phospholipids. 
     In very simple (non-blood) media where scattering is negligible and absorption is the primary effect, Beer&#39;s law can be used to model the light intensity as it passes through the medium: 
     
       
         I(x)=I o e −μ   a   x   
       
     
     where I(x) is the intensity of the transmitted light at a distance x travelled through the medium, I o  is the incident light intensity, and μ a  is the absorption coefficient. 
     However, the situation in the blood  11  is not so simple. In blood, the scattering effect cannot be neglected, and in fact, is a much larger factor than absorption. The absorption coefficient (μ a ) for light in blood at a wavelength of 800 nm is approximately 1 cm −1 . This value is the inverse of the mean free path (the mean distance travelled by the individual photons through the blood medium before absorption). This number is highly dependent on several factors, including the oxygen saturation of the blood, the hematocrit (% by volume of red blood cells), and the wavelength of light used. FIG. 4 shows the three major blood components and their absorption coefficients&#39; dependence upon wavelength. In FIG. 4, curve  30  denotes de-oxyhemoglobin, curve  32  denotes oxyhemoglobin, and curve  34  denotes water. The values for oxygen saturation (fraction of oxyhemoglobin to total hemoglobin) and hematocrit may change over the course of the surgery, which may in turn cause variation of the absorption of the blood as the relative concentrations of the three absorbing components in FIG. 4 change. Because slight concentration changes cause large changes in absorption in certain parts of the spectrum (e.g. 900-1000 nm), the invention uses a wavelength in the 800-850 nm range, where the de-oxyhemoglobin is level (i.e. the dependency on oxygen saturation is eliminated) and the water absorption is minimized. 
     The scattering coefficient (μ s ) for light in blood is approximately 300 cm −1  assuming μ s =Σ (fraction of component) (μ s  of each component) for components such as water, platelets, red blood cells with oxy-hemoglobin and de-oxyhemoglobin, etc. This coefficient describes all scattering occurrences, including both forward and backscattering. Another factor, g, is introduced to describe the mean cosine of the scattering angle during scattering events. The effective scattering μ&#39; s =μ s  (1g) is a measure of the degree to which large angle scattering events occur in the medium. 
     In, blood g≈0.974 so μ&#39; s ≈8 cm −1 , still much higher than the μ a ≈1 cm −1 . 
     The apparatus of this invention can distinguish between microair bubbles  13  and small particulates such as blood clots and bone chips. Because of their increased absorption over blood, the light intensity drops significantly behind these particles, forming an effective shadow, instead of the increase in intensity due to a bubble. This makes particulates and bubbles easy to distinguish. 
     In order to increase the accuracy of the bubble size determination, it is advantageous to use a plurality of axially spaced detector sets  16  (FIG.  5 ), or to use CCD arrays which can track individual bubbles as they move through the detector&#39;s field of vision. Because bubbles rise within the bloodstream at a rate generally proportional to their size, a time correlation can be obtained by observing the intensity signal at spaced points along the tube  12 . This time correlation in turn can be used to check and increase the accuracy of the optical bubble measurement. 
     FIG. 6 shows the effect  22  of the passage of bubbles  13  of different sizes past the detector  10  of FIG. 1 (the more negative the output voltage in FIG. 6, the more light has passed through the blood  11 ). By measuring the amplitude (and, in the embodiment of FIG. 5, the timing) of the peaks  22 , and the relation of artifacts  23  to the peaks  22 , a conventional comparator  24  can provide an indication, to a numerical display  26 , of the count of bubbles  13  in various size ranges. 
     It is understood that the exemplary optical detection and quantification of microair in blood described herein and shown in the drawings represents only a presently preferred embodiment of the invention. Indeed, various modifications and additions may be made to such embodiment without departing from the spirit and scope of the invention. Thus, other modifications and additions may be obvious to those skilled in the art and may be implemented to adapt the present invention for use in a variety of different applications.