Abstract:
A retinal blood flow velocimeter projects an illumination beam through a steering system onto a retinal vessel, and forms a separate tracking image back through the steering system. A fast tracking loop detects motion of the tracking image and moves the steering system to null image motion and keep the illumination beam centered on the vessel. The beam is reflected from the vessel, picked up by detectors at two fixed angles, and processed by spectral analysis. In one preferred embodiment the illumination beam and the steering system follow entirely separate paths through the steering system. Fiber optics translate the collected Doppler light without dispersion while preserving phase relationships, and absolute dimensions are determined from the image tracking electronics. A processor then computes volumetric blood flow which it compares with normative data.

Description:
BACKGROUND OF THE INVENTION 
     The present invention relates to instrumentation for measuring blood flow in vessels of the retina by Doppler velocimetry. 
     The general theory of laser Doppler velocimetry, as applied to the measurement of a flowing fluid such as blood inside a blood vessel, is a well known application of flow measurement technology. Briefly, monochromatic light aimed at the vessel and into the flowing blood is reflected by the blood cells as diffuse light with a frequency distribution corresponding to the components of velocity of the individual scatterers. By analyzing the frequency distribution of the reflected light at two fixed receivers with a known separation angle, the velocity or, ideally, the velocity profile of the flowing blood can be deduced. 
     When one attempts to apply this approach to detect blood flow rates in vessels of the retina, however, practical obstacles are encountered. First, individual retinal vessels have a diameter under several hundred microns, so that in order to perform a reliable measurement it is necessary to aim a beam of laser light of diameter approximately equal to the diameter of the vessel. Smaller beam diameters introduce the risk of missing the centerline flow measurement, while larger beam diameters result in a lower signal to background ratio. 
     Second, the Doppler analysis requires collection of the reflected light from two distinct directions having a specified angular separation. This light collection must be done outside the eye. The optical paths therefore will vary depending on the curvatures of the eye involved, and the collected light will include extraneous light due to reflection at various surfaces of the eye. 
     Third, it is necessary to perform this aiming and to collect a sufficiently strong return signal, despite relatively fast and large scale movements of the eyeball. 
     When it is considered that a small diameter beam must be used to maintain an acceptable signal to noise ratio, and that the level of reflected light from the fundus that can be collected outside the eye is highly attenuated, the foregoing obstacles are seen to impose severe limits on the quality of collected light available for Doppler analysis. 
     These difficulties have heretofore limited the clinical applicability of laser Doppler velocimetry to carefully controlled and rather cumbersome analytical investigations. Typically, the procedure is done by fitting a rectifying lens directly on the cornea, and then, with the illumination and collection optics manually positioned on a target vessel, recording short time segments of the collected spectra. A large number of such recordings are then analyzed and segments are pieced together to obtain an analytically derived synthetic recording representing the flow during one or more entire heartbeat intervals. The analysis and ultimate synthesis or identification of a representative one- or two-second Doppler spectrum is done some time after the recording, so that blood flow information is not quickly provided. 
     One approach to simplifying the processing of the recorded Doppler spectra is to develop algorithms for initially selecting only those recorded spectrum segments which meet certain criteria representative of the expected flow functions. Highly noisy or anomalous recording segments are discarded, thus limiting the amount of remaining data that must be processed. This approach, while clearly eliminating records resulting, for example, when the beam misses a vessel entirely, may screen out some valid flow information and render the system blind to clinically significant details. Analysis of Doppler records would be simplified if the instrumentation could be aimed with sufficient stability to record a continuous record having a duration of a full heartbeat interval or longer. More meaningful measurements of blood flow could also be obtained if the stability were sufficient to allow aiming a Doppler illumination spot on a central region of a blood vessel and on smaller vessels. 
     SUMMARY OF THE INVENTION 
     It is an object of the invention to provide a retinal Doppler velocimeter of enhanced utility and performance. 
     It is another object of the invention to provide a retinal Doppler velocimeter which is reliably positioned and maintained on a retinal vessel. 
     It is another object of the invention to provide a retinal Doppler velocimeter which provides continuous or real time vessel flow information. 
     These and other desirable features are achieved in accordance with one embodiment of the invention by providing an optical beam steering system for controllably steering a beam directed at the retina, and by projecting a Doppler illumination beam through the steering system in a forward direction while forming an image of the retina along an optical path that passes through the steering system in a reverse direction. A tracking system detects motion of the image and develops control signals to produce compensating motions of the steering system so that the Doppler illumination beam remains centered on a thin blood vessel. With the illumination thus stabilized, a set of collection optics collects the light reflected from a retinal vessel along two distinct directions and an analyzer determines the spectrum of collected light, and preferably also computes or displays at least one of the peak or minimum velocity, the time-averaged centerline velocity, or the corresponding volumetric flow rate. 
     The steering system contains optical elements arranged so that the forward and reverse optical paths are separated. 
     In a preferred embodiment, the Doppler collection optics are located behind the steering system to provide an unobstructed area between the instrument and the eye, and are positioned so the angles at which the light is collected bear a fixed angular offset. 
     The steering system includes a pair of two sided mirror elements, each pivotable about one of two mutually orthogonal steering axes, and an optical relay system which places a face of each mirror element in a conjugate relation to a face of the other mirror element. 
     In another or further preferred embodiment of the system, the blood vessel is imaged as a tracking target transversely onto a linear CCD array, which provides a direct measure of the vessel diameter. A processor computes the vessel&#39;s volumetric flow rate as a function of centerline blood flow velocity and vessel diameter. In a further embodiment according to this aspect of the invention, the processor may include a stored table of normal flow rates as a function of vessel size and the subject&#39;s age, and may provide a diagnostic output based on a comparison of the detected and the normal flow. In another embodiment, the processor may store diagnostic programs for summing the flow over several vessels and detecting discrepancies indicative of flow pathologies. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     These and other features will be understood by reference to the following description of illustrative embodiments of the invention, together with the drawings, wherein 
     FIG. 1 illustrates one embodiment of the invention; 
     FIG. 2 illustrates another embodiment of the invention; 
     FIG. 3 illustrates the Doppler collection optics of the apparatus of FIG. 1 or 2; 
     FIG. 3A illustrates alternate optics of the steering assembly; 
     FIG. 3B illustrates alternate optics of the Doppler collection apparatus; 
     FIG. 4 illustrates the processing of Doppler signals; 
     FIG. 5 illustrates a representative Doppler spectrum; 
     FIG. 6 illustrates the Doppler spectrum processed to identify flow rate; 
     FIGS. 7, 7(a), (b), and (c) illustrate the instantaneous maxima of Fourier spectra plotted over time, and the Doppler signal processing; 
     FIG. 8 illustrates tracking and Doppler illumination of a retinal vessel; 
     FIG. 9 illustrates the CCD image signal for determining vessel diameter; and 
     FIG. 10 illustrates processing of a diagnostic Doppler measurement system. 
    
    
     DETAILED DESCRIPTION 
     FIG. 1 illustrates a stabilized retinal laser Doppler system 10 in accordance with one embodiment of the present invention. System 10 includes a steering assembly 20, a tracking assembly 30, a red laser source 35 for illuminating a retinal vessel, and a two-channel Doppler pick up and analysis assembly 40. The steering assembly, which includes x- and y- axis deflection mirrors, is controlled by electrical signals to direct the optical path 100 of the beam from the red laser to a desired retinal vessel, and the beam is maintained centered on the targeted vessel by a tracking system which monitors the position of the image of the same or a nearby retinal blood vessel which has been imaged back through the same steering system into an electronic sensor array, e.g., a CCD 38. Change in position of the image on the CCD array is detected and used to develop control signals that reposition the steering mirrors to prevent motion of the image. The techniques for deriving a well defined tracking signal and controlling the mirrors with a sufficient speed and accuracy to aim the beam from laser 35 on a retinal target are described in U.S. Pat. No. 4,856,891 commonly owned by the assignee of the present invention, and the text of that patent is hereby incorporated herein by reference for purposes of a complete description and full disclosure. 
     The red laser light scattered from a targeted retinal vessel at the back of the eye is imaged by the eye objective optics back toward the steering system and is deflected by a pair of mirrors 41, 42 each having a diameter of approximately one millimeter and spaced approximately six millimeters apart. These mirrors reflect collected light into respective channels of the Doppler analysis unit. The mirrors each intercept about 1.4° of arc, and their spacing, corrected for the 3× magnification of the objective assembly, corresponds to a fixed divergence angle within the eye which allows calculation of absolute flow velocity values, when given the axial length of the subject&#39;s eye. 
     Observation light for the system is provided by a yellow helium-neon laser 51, which is directed through a beam expander 52, past deflecting mirrors 53a, 53b, and through an attenuator 54 to provide a broad field beam which is folded into the illumination path by beamsplitting mirror 55. The beam illuminates a ±10° field of the fundus. 
     Yellow observation light reflected from the fundus is returned through the objective assembly consisting of lenses 61, 63 and an image rotator 65, and passes through the steering assembly to an eyepiece or viewing assembly 67 where it provides a visible image field that moves synchronously with the tracking image and with the targeted vessel and Doppler illumination spot. The viewing assembly may include a camera. The function of the image rotator 65, described more fully in the aforesaid U.S. Patent, is simply to rotate a tracking image, such as the image of a retinal vessel, into a fixed orthogonal frame on the CCD. This allows the tracker to lock onto an obliquely oriented vessel and apply its fixed-frame orthogonal steering corrections. The image rotator thus provides additional convenience in setting up the instrument, and removes the need for image field transform computations in the tracking system. 
     In this illustrated embodiment, a green helium-neon laser 31 is provided for the tracking system. Laser 31 provides a separate beam which is folded into the same optical path 100 as that followed by the Doppler illumination beam by a turning mirror 32 and a beam splitting mirror 33, so that the green beam is also steered by the steering assembly 20. An attenuator 101 in the path limits the intensity of the steered beam. The green tracking beam has a small diameter, e.g., under several millimeters, and thus beneficially limits the illuminated area of the eye. The wavelength separation of the three described light sources allows appropriately placed filters or dichroic beam splitters to eliminate interference from each of the different sources on the viewing or sensing units associated with the other sources. For example, the beamsplitter 37 which reflects the return tracking image to the CCD 38 may be a dichroic beamsplitter which reflects substantially all the green light toward the CCD, while passing substantially all the yellow light to the observation optics 67. Further concrete examples of appropriate spectral separation paths are more fully described in the aforesaid U.S. patent. 
     The steering system 20 includes two steering mirrors 21, 22 each arranged to pivot about one of two orthogonal axes lying in a common plane P which is conjugate to the eye fundus. A galvanometer control 21a, 22a attached to a pivot shaft moves each mirror so that it is precisely turned to a direction within an angular range of approximately ±1020 . Each mirror has first and second sides, denoted the A (or inside) and the B (or outside faces) herein, and according to a Principal aspect to the invention these mirrors are arranged to maintain optical separation of the input and output light paths. 
     This separation is achieved by an optical relay system which translates the outside faces of the mirrors to each other, preferably including lenses or focusing mirrors which place the turning axis of the one mirror conjugate a plane containing to the turning axis of the other mirror with a 1:1 magnification. Such conjugation optics are more fully described in the co-pending United States Patent application Ser. No. 522,376 of Yakov Reznichenko and Michael Milbocker entitled Bidirectional Light Steering Apparatus, filed on May 11, 1990 and commonly owned by the assignee of the present invention. Said patent application is hereby incorporated herein by reference. 
     For ease of illustration, however, the intermediate lenses or curved reflective surfaces are omitted from the drawing, and the optical relay system is shown simply by three flat mirrors 23a, 23b, 23c which translate a beam impinging on the B face of one steering mirror to the B face of the other steering mirror. As further described in the aforesaid patent application, the steering mirrors may be thin plates which are metallized on one side, but are preferably front-surface mirrors metallized on both sides. This construction more effectively eliminates ghosting and internal reflection from the steering system. 
     The return image along axis RI from the subject&#39;s eye is reflected from the &#34;A&#34; face of mirror 22 to the &#34;A&#34; face of mirror 21, and thus passes through the steering system with the same angular deflection as a reverse-steered beam 200 passing to the tracking and observation optics, so that the light input beam 100 and the return image beam 200 always follow substantially fixed directions to and from the tracking/observation optics. A pair of diaphragms 24a, 24b located in a fundus conjugate plane screen out corneal and other reflections. The diaphragm opening is approximately ten millimeters. 
     Turning briefly to FIG. 3, the Doppler signal reception assembly 40 of FIG. 1 is illustrated in greater detail. The pick-off mirrors 41, 42 deflect two portions of the reflected Doppler beam which define in this embodiment a precise angular separation corresponding to a 13.5° divergence outside the eye. The light is relayed to a fiber bundle 47 in each channel. Each bundle 47 serves to channel the light received at one end of the bundle without further divergence or attenuation, while preserving phase relationships, to a photomultiplier tube 49 (RCA 8645). A red laser line filter 48 (Melles Griot 632.8 nm) removes extraneous wavelengths. 
     If it is desired to retain a manual tracking or viewing port in the Doppler assembly, a construction such as shown in FIG. 3B may be employed. In this construction the mirrors 41, 42 may be larger, and a pair of pinhole diaphragms 44 define the Doppler beam separation angle. An optical relay assembly consisting of objective optics 46 and relay mirrors 46a, 46b in each channel relay the collected light to the respective fiber bundles 47, and an annular green filter (a Kodak Wratten filter #57A, not shown) placed in the optical path together with an eye objective 50 provides an additional or alternative way to view the targeted vessel during Doppler measurement. 
     In any case, the Doppler illumination beam from laser 35 is focused to a spot of a diameter approximately equal to the diameter of the targeted vessel on the retina, and the incident beam power is attenuated to approximately five microwatts, resulting in a biologically safe level of retinal irradiance. While the photomultiplier tubes necessary to detect return irradiation at these low levels would be driven to saturation by normally encountered stray reflectances, in the illustrated apparatus the photomultiplier tubes develop an acceptable signal due .in large part to the above described separation of the tracking and illumination signals in the steering system. 
     In a further embodiment shown in FIG. 2, the Doppler pick off and receiving assembly 40 is positioned on the opposite side of the steering system from the eye. In this embodiment, the angular relation between each pick-up and the input illumination beam is a constant, thus eliminating second order effects. A further construction difference resides in the replacement of the mirrors 41 or 42 (FIG. 1) and 46a, 46b, (FIG. 3B) with a longer fiber bundle 47 for each channel that extends directly into the return image path and conducts light to the photomultiplier tube. The bundles have a diameter of slightly over three millimeters, and translate the light from the retinal conjugate image plane without dispersion while preserving relative phase relationships. 
     FIG. 3A illustrates an alternative construction of the steering system 20 for use in the Doppler apparatus of FIGS. 1 or 2. In this embodiment, the x- and y- steering mirrors 21, 22 are identical to those of FIGS. 1 and 2, but the relay path between the outer faces of those mirrors consisting of mirrors 23a, 23b, 23c and associated relay lenses has been replaced by a pair of focusing mirrors 24a, 24b in a unity -magnification telecentric arrangement. This simplifies the layout and alignment of the steering assembly, reduces the number of reflective interfaces, and eliminates solid scattering media from the illumination path. 
     FIG. 4 illustrates the processing of collected light of the Doppler analyser. The reflected light includes light scattered from the surface of the blood vessel which serves as a reference frequency, as well as light which has Penetrated the vessel and is scattered from blood cells flowing within the vessel. These two types of light are combined on the photomultiplier tube 49 (FIG. 3), where they heterodyne to produce an electrical signal having beat frequency components corresponding to the individual velocities of the scattering cells. The electrical signal developed by each photomultiplier channel is fed to a spectral analysis system 110, which for each five millisecond interval provides an output in real time representative of the frequency components of the analysed signal, of which a representative trace is illustrated in FIG. 5. The trace is quite noisy, as it is derived from the individual motions of scattering objects which follow some general cross-sectional flow profiles within the vessel, but which also have components of motion due to thermal motion and fluid flow turbulance and irregularities. However, despite the extreme noisiness of the signal, the frequency trace does have an ascertainable upper or cut-off frequency 120 (FIG. 5) corresponding to the maximum or centerline flow velocity value of the target vessel. 
     In order to detect this maximum flow velocity value, the signal trace (FIG. 5) of the spectral analysis processor 110 for each diameter is digitized and passed to a processor 115 which determines the maximum frequency and displays the corresponding flow velocity. In the preferred embodiment, the cut-off frequency is identified by an integrator/differencer which constructs a new function from the output of the spectral analysis system 110 such that the new function has a maximum value at the cut-off. This processing is implemented in a software module, which for each frequency value u defines a &#34;window&#34; of width 2A about the value, and slides the window along the frequency scale. For each u, it subtracts the value of the frequency signal integrated over a fixed interval of width A to the right of υ, from the value of the frequency signal integrated to the left of υ. 
     The resulting function, which for each frequency υ o , is defined by ##EQU1## has a maximum precisely at the frequency where there is an extreme discontinuity in fluctuation of the signal value. FIG. 6 shows the function f(u) so defined, with the same frequency scale as illustrated in FIG. 5. Thus, it is seen that despite the jumpiness of the spectral output, the frequency corresponding to peak blood flow is readily detected. 
     FIGS. 7, 7A, 7B and 7C illustrate the basic signal processing of the above-described system. The line A of FIG. 7A shows an eight second signal trace consisting of the frequency maximum at each instant in time derived by the spectral analysis system 110 from the output of one photomultiplier tube 49 when the Doppler illumination spot is aimed at a retinal artery. The line of FIG. 7B shows the corresponding trace of the other photomultiplier. Each channel has different absolute frequency range, owing to their different light collection angles, but both show the distinctive periodic pulses associated with arterial flow due to the cardiac pumping cycles. The line of FIG. 7C shows the blood flow velocity equal to a constant, for a given eye and instrument configuration, times the difference between lines A and B. Specifically, line C represents the peak instantaneous centerline blood flow velocity, which, at a given instant, is directly proportional to the difference in peak or cutoff frequencies of the two signals, lines A and B. 
     FIG. 7 illustrates the overall operation of the spectral analysis system and processor of FIG. 4. Each PMT analog output is A/D converted and stored in a computer-accessible form, e.g., on a disc. A software Fourier transform module analyses each t-second block of signal values and computes its power spectrum. Each power spectrum (channels 1 and 2) is Processed by the frequency cut-off detection algorithrum described above in relation to FIG. 6. In the prototype instrument, the processor operated in real time to digitize and process eighty-nine five-millisecond samples per second for each channel, producing the highly detailed traces illustrated in the figures. 
     It further bears note that in FIGS. 4 and 7, the function of the spectral analysis system and the processor are not clearly separated, for the reason that in the preferred embodiment the spectral analysis and subsequent signal processing steps may be primarily performed by the processor, which may be a microcomputer equipped with numerical analysis software and with Fourier transform software. 
     In a preferred embodiment of the invention, further functions are performed in the processor on other opto-electronic signals to develop a number of specific indicators or pieces of clinical information as more fully described below, including volumetric blood flow outputs, vessel blockage or flow anomaly determinations, and normative comparison of circulation. 
     It should be noted that because the Doppler analysis module uses the light reflected from the outside of a blood vessel as a reference beam, frequency shift effects caused by motion of the eye or of the illumination spot cancel out, and the detected flow rate is substantially the same whether the focused Doppler beam is stationary or is moving along the long direction of the blood vessel, with or opposed to the flow direction. For this reason, the tracking system need not control motion in two dimensions, but may be a one-dimensional tracker with its tracking components configured in an orientation to correct only for motion of the eye in a direction transverse to the vessel at which the Doppler beam 100 is directed. This may be accomplished, for example, when using a tracker as shown in the aforesaid U.S. patent, by choosing a tracking target vessel which either is, or lies parallel to, the vessel on which Doppler measurements are to be taken, and by tracking motion transverse to that tracking target to develop steering correction signals. 
     It is also possible to use a two-axis tracking system to stabilize the Doppler beam and return light for analysis and imaging. 
     In one presently preferred embodiment of the invention, a single-axis tracker is employed, and output signals from the tracking CCD provide quantitative measurements to convert the Doppler output to absolute volumetric flow measurements. 
     FIG. 8 illustrates details of the Doppler imaging of such a device. A retinal vessel 120 which may have a diameter of under fifty to a few hundred micrometers, is illuminated by a green tracking beam 90 which has a round or rectangular cross section of approximately 0.5-1.0 mm diameter, and the Doppler illumination beam is focused to a spot 95 on the same vessel. The retinal region illuminated by the tracking beam 90 is imaged and aligned, via the steering system 20 as described above, as an image 90&#39; onto a CCD line array 130 which is oriented perpendicularly to the image 120&#39; of the vessel. A magnifying objective assembly of five to twenty five magnifications is used, such that the CCD lies entirely within the image 90&#39; of the tracking beam. For example, for a linear array consisting of a one by two hundred fifty-six pixel CCD of approximately twelve millimeters length, a twenty-five power objective assures that the image of a five hundred micrometer wide tracking beam will cover the CCD 130. Correspondingly, the image of a fifty to one hundred micrometer retinal blood vessel will cover approximately twenty-five to fifty pixel elements of the CCD. 
     FIG. 9 illustrates the sensed illumination values, along the length of the CCD, of tracking light reflected from the retina. The characteristic double-valley minimum in detected light intensity corresponds to the blood vessel image, with a central local maximum corresponding to the specular reflection from the top center of the vessel wall. The full width half maximum points of the illumination values, illustrated by the two arrows, correspond to the vessel diameter d. 
     In the above described preferred embodiment, the tracker not only polls the CCD at one millisecond intervals to determine position-correcting control signals from the steering mirrors, but processes the CCD output to determine the vessel diameter d by solving for the full width half maximum points. 
     In a further preferred embodiment, the processor further performs internal computations to combine the detected flow velocity and vessel diameter, and to compute an absolute volumetric flow rate. 
     The preferred processing for determining the volumetric flow rate proceeds as follows. First, the processor integrates the centerline flow velocity (FIG. 5) over a time interval, by numerical processing, and divides the integral to determine an average centerline blood speed &#34;acb&#34;. Next, the processor determines the vessel cross-sectional area A=π(d/2) 2  from the vessel diameter d. For sufficiently large vessels, (over about fifty micrometers) the assumption of Poiseuille flow holds, and the total volumetric flow is calculated to be A (acb)/2. 
     This capability of directly computing the volume of blood flow in a vessel during observation is advantageously augmented in several further embodiments of the invention to provide systemic measurements of total flow, branch flow anomalies, blockage or general sufficiency. In one such embodiment, the processor stores a memory table listing the range of normal blood flow as a function of vessel size. This information may be stored separately for arteries (recognizable by their distinctly pulsatile flow) and for veins (having a more uniform flow rate). 
     Once a vessel has been targeted by the tracker and its diameter d determined, the stored normal value indexed by the diameter d is retrieved and the normal value is compared to the computed flow value to determine whether there is an anomaly. 
     In another embodiment, the &#34;normal&#34; value need not be a predetermined universal value, but may be determined by measuring and storing the flow values for vessels of varying diameters in a patient&#39;s healthy eye; and the comparison is then made against the measured flow velocity or volume values of the other eye. It will be understood that the &#34;normal&#34; values need not be functions only of diameter, but may also be ordered or indexed as functions of the subject&#39;s sex or age, the subject&#39;s blood pressure, or other clinical parameter. 
     In a further variation of this embodiment, the processor may include means for summing the flow rates of each of a plurality of arteries, and for summing or subtracting the flow of a plurality of veins, thus providing indications of the blood flow for whole regions of the retina. An imbalance in the total flows into or out of a retinal region provides an indication of flow anomaly indicative of possible pathology. In other embodiments, a simple comparison to a threshhold flow value may indicate a particular pathology such as hemorrhaging, or a detached retina. General processing states for one or more of these further systems are illustrated in FIG. 10. 
     This completes a description of the invention and several representative embodiments thereof, together with subsidiary details and variations of construction. The invention being thus disclosed, modification and equivalents thereof will occur to those skilled in the art, and such modifications and equivalents are considered to lie within the scope of the invention, as determined by the claims appended hereto.