Abstract:
An instrument has been designed to study the pulsatile motion of the eye by analysis of a beam of light reflected from the corneal surface. A laser light beam probe of small spot size and low divergence strikes the cornea apex and the reflected movement is recorded by a sensor. Analysis of the beam movement reveals the energy in the eye pulse without the necessity of physically touching the eye. The value of the intraocular pressure is determined from the calculated power spectrum. The sensitivity, accuracy and efficiency of the light beam makes possible studying both eyes concurrently and comparison of the pulse parameters of onset, amplitude and duration reveals any delay in circulation to an eye.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application is related to U.S. Provisional Patent Application Ser. No. 60/920,209 filed on Mar. 26, 2007, the entire disclosure of which is incorporated by reference as if set forth in its entirety herein. 
    
    
     FIELD OF THE INVENTION 
     The present invention relates generally to ophthalmic clinical screening for the detection of eye and cerebrovascular disease by means of non-contact measurement of the ocular pulse amplitude of a subject&#39;s eyes. 
     BACKGROUND OF THE INVENTION 
     Determination of the intraocular pressure (IOP) is a component of the standard complete eye examination, a necessary test in the diagnosis and management of glaucoma and an indicator of underlying cerebrovascular diseases. A direct measure of the IOP requires surgically opening the eye and placing a device within it. This is not clinically useful on a large scale. Clinically, indirect methods are required to determine the IOP. The instruments used for this determination are called tonometers. Tonometers evaluate the IOP by seeing how much force it takes to distort the eye and inferring what intraocular pressure is required to resist the applied force, and declaring that to be the IOP. 
     The force is applied to the eye by a plunger, an illuminated prism, a blast of air, a pneumatic pressure probe or piezoelectric crystals. All of these methods suffer from multiple problems. First, distorting the shape of the eye changes the pressure in the eye reducing the accuracy of the test. Second, all of these instruments have a large mechanical component introducing friction into the measurement. The damping effect of friction decreases the sensitivity of the test. Third, the force required for a tonometer to indent or flatten the eye surface in measuring the IOP is large compared to the energy contained within the ocular pulse wave. Therefore, the information carried in the ocular pulse wave and created in the eye following each heart beat remains hidden. Fourth, the lack of ability to produce detailed ocular pulse wave measurements eliminates the opportunity to test both eyes concurrently and to compare this data between the eyes. Fifth, applying force to the eye subjects the cornea to trauma. 
     SUMMARY OF THE INVENTION 
     Our invention relates generally to ophthalmic clinical screening for the detection of eye and cerebrovascular disease by a non-contact measurement of the ocular pulse amplitude waveform 
     The invention, an ocular pulse analyzer, probes the eye with a light beam and evaluates the eye&#39;s expansile response to an ocular pulse by analyzing movement of the reflected light beam. Valuable information about the ocular pulse is recorded in real time. The eye is not touched. The eye is not distorted. There is no friction interfering with the data and there is no trauma to the eye. The time of onset of the ocular pulse after the heart beat is easily measured. The amplitude of the ocular pulse, the duration of the ocular pulse and the shape of the ocular pulse are recorded to be compared with normal values for each parameter. Since both eyes are tested concurrently, they are compared with each other. 
     An embodiment of the ocular pulse analyzer consists of a pulsed laser diode followed by a laser beam expander/collimator and a laser beam shaper which form a low divergence small spot incident beam that strikes the corneal surface at a large incident angle. The reflected beam emerges at an equal and opposite incident angle, striking the analyzer CCD array forming a pixelated image of the beam spot. During a single ocular pulse up to 128 points are obtained by calculating the center-of-gravity of each reflected beam spot profile. The distance of each data point from the CCD array systolic reference point is proportional to the distance the front eye surface has moved. The collection of these points defines the ocular pulse amplitude waveform for each heart pulse. Sensors are used to generate triggers that pulse the laser and synchronize the analyzer control and data acquisition. Fast Fourier Transforms (FFT) are used to analyze the data producing both the ocular pulse waveform and its power spectrum. The zero-frequency power spectrum area is a measure of the energy in the mean intraocular pressure (IOP). Non-zero frequency power spectra areas measure energy that may be related to pressure changes during the ocular pulse and may be indicators of eye disease. Left-right eye data comparison allows rapid detection of eye and cerebrovascular disease. Two of these optical analyzers may be mounted so that both of the subject&#39;s eyes can be examined simultaneously. 
    
    
     
       BRIEF DESCRIPTION OF THE FIGURES 
       Understanding of the present invention will be facilitated by consideration of the following detailed description of an embodiment of the present invention taken in conjunction with the accompanying drawings, in which like numerals refer to like parts and in which: 
         FIG. 1  illustrates a functional block diagram of the device according to an aspect of the present invention; and 
         FIG. 2  illustrates a diagram of a folded optical system for the left eye for measuring the movement of the eye front surface during the ocular pulse for the apparatus of  FIG. 1 ; and 
         FIG. 3  is a cross section of the light beam shaping, collimator, and beam dump module  203  of  FIG. 2  for the apparatus of  FIG. 1 ; and 
         FIG. 4  is a layout of the patient and operator optics  107  and  108  of  FIG. 2  for the apparatus of  FIG. 1 ; and 
         FIG. 5  illustrates how the light beam responds to the movement of the eye front surface in response to the ocular pulse; and 
         FIG. 6  is a plot of the light beam optical magnification vs. the incident angle of the input laser beam illustrating the variation of magnification; and 
         FIG. 7  is a plot of the displacement of the light beam vs. the displacement of the eye surface in response to the ocular pulse; and 
         FIG. 8  is a plot of the reflected light beam energy at the optical detector as a function of the light beam incident angle; and 
         FIG. 9  is a plot of a simulated power spectrum illustrating the IOP as the zero frequency component with three higher frequency components. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     It is to be understood that the figures and the descriptions of the present invention have been simplified to illustrate elements that are relevant for a clear understanding of the present invention Several standard optical elements and other components are used to expand and collimate and form the laser beam and to control and process it. Those of ordinary skill in the arts will recognize that these and other optical elements and/or steps are well known in the art, but because they do not facilitate a better understanding of the present invention, a discussion of such elements and steps is not provided herein. The disclosure herein is directed to all such variations to such elements and methods known to those skilled in the art. 
       FIG. 1  is a block diagram of an embodiment of the ocular pulse analyzer. A solid state laser  101  produces a pulsed axially symmetric Gaussian beam of 670 nm wavelength and 1 milliwatt power, which is injected along the incident-beam-line  10  into the beam forming optics module  102 , which forms a small diameter low divergence exit beam of up to 500 micrometers diameter and less than 1.0 milliradians divergence, which strikes the apex of the cornea O at a 65 to 75 degree angle with respect to the eye visual axis. The incident beam is reflected onto the reflected-beam-line OR at an angle of −65 to −75 degrees where it strikes a CCD array  105  forming a pixelated image of the reflected beam spot. The motion of the beam spot pixelated image on CCD array  105  is proportional to the movement of the eye front surface  104  as it responds to the ocular pulse. The eye orbit cup  109  is a compliant cup that forms an air tight seal on the eye orbit so that the equilibrium pressure at the eye front surface can be varied which allows parametric studies of the eye. Cup  109  also is the interface to the patient&#39;s face. The optics target module  107  projects an infinity image of an eye fixation target on which the patient focuses his vision thus minimizing random eye movement. Accessory optics port module  108  which allows access for the external calibration of the optics system. Both the fixation module  107  and the accessory optics port module  108  are aligned to the apparatus centerline. The ocular pulse analyzer beam forming optics are contained in beam tube  103 , and the reflected beam components are contained in beam tube  106 . The fixation module  107 , the observation module  108  and the eye orbit cup  109  are attached to an instrument stand  110 . 
     The laser pulser is synchronized to the heart pulse by delayed triggers derived from the blood pulse sensors BPS. During each heart pulse up to 128 pixelated beam spot widths are digitized by the analog-to-digital-converter ADC and processed to provide measurements of the eye surface movement. The computer/display module CPU/Disp system processes these data providing the operator with the IOP and other parameters. These data are stored in the patient data base DB. The CPU/Disp receives the pulse and timing data and sends control data to the DAC which then controls the Laser Pulser and the cup control module Cup Cntrl. By using the real time control module RTCntrl the operator controls the measurement and data processing parameters. 
       FIG. 2  illustrates an embodiment of a compact left eye optic head  200  with mirror folded beam lines allowing high angle light reflection on the cornea. The incident laser beam from laser  101  is transformed by the beam forming module  102  which is comprised of a 2.5× inverted Kepler&#39;s telescope that expands and collimates the beam. Kepler&#39;s Telescope consists of a diffraction limited entrance lens  201  such as an infinity corrected 5× microscope objective of 36 mm focal length and numerical aperture 0.1, which forms a beam waist of about 14 micrometers diameter at image plane  201   a , followed by an exit lens  202  such as a 90 mm focal length doublet located about one focal length from plane  201   a  which forms a well collimated exit beam of at most 10 mm diameter and a total beam divergence of less than 0.5 mrad at the entrance of beam shaper, collimator and dump module  203  which attenuates and collimates the beam using apertures of decreasing diameter. At the beam shaper, collimator and dump module  203  exit the beam has a diameter no larger than 500 micrometers, a divergence of less than 1.0 mrad and a power of less than 1 microwatt. The beam then strikes turning mirror  204 , which reflects the incident beam so that it strikes the cornea at an angle of from 65 to 75 degrees to the normal of the eye where it is reflected at an equal and opposite angle striking turning mirror  205 , which reflects the beam onto the normal to CCD array  105 . The beam then passes through the quadrature error detectors  206  which detect reflected beam centering and out of range reflected beam pulses, finally impinging onto CCD array  105  which forms a pixelated image of the reflected beam spot. To provide noise and stray light shielding; the beam forming module  102  is mounted in beam tube  207 , the reflected beam tube components  206  and  105  in beam tube  208  and the compact optic head  200  in instrument case  209 , which also mounts the orbit cup  109  and mounts the fixation module  107  and the accessory optics port module  108 . As may be understood by those possessing an ordinary skill in the pertinent arts instrument case  209  may be mounted on a standard slit lamp bench or other ophthalmic examination stand. 
       FIG. 3  details an embodiment of a beam shaper, collimator and beam dump module  203  which simultaneously forms the probe beam that is required for observing the unperturbed movement of the eye front surface. The expanded laser beam enters module  203  through an aperture in wall  302  where the laser beam diameter is reduced by striking the walls which are conical and may be made of graphite, the beam then passes through four alternating radial light traps  306  and four decreasing apertures  304  further reducing the beam diameter, divergence and power. Light traps  306  collect large angle light rays, and scattered rays from the wall by multiple reflections in the radial traps further reducing stray light. After passing through a beam defining exit aperture  308  the beam power may be reduced by a factor greater than 1000, may have a reduction in beam divergence to less than 1 milliradian and a diameter of less than 500 micrometers. As may be understood by those possessing an ordinary skill in the pertinent arts the apertures may be produced by laser or other machining means and may be coated with an optically absorbent coating by sputtering or other means. The internal parts may be assembled with spacers into a stack that is self aligning. 
       FIG. 4  illustrates the eye fixation module  107  and accessory optics port module  108 . An incandescent light source  402  illuminates a ground glass  404 , which illuminates an optical fixation target  406 . Light from target  406  passes through beam splitter  408 , onto lens  410  which forms an image of the fixation target at infinity, which the patient fixes his eye on, minimizing random eye motion. Light from the eye surface  412  is reflected 90 degrees by beam splitter  408  onto lens  414  which forms an image of eye surface  412  on at the exit of the accessory optics port  416 , which provides access for external calibration of the optics system. 
       FIG. 5  shows a geometric model of how the reflective head functions. Initially the optic head is aligned to the eye surface at t 0  the time of the systolic ocular pulse, when the eye surface  104  has moved to its maximum outward extent  104 P 1 . At this time the optics projects the apertured beam spot image onto the corneal apex (0,0) and reflects it onto the CCD array  105  which forms a pixelated spot image that is collected by the data processing system. As the ocular pulse amplitude reduces, the eye surface  104  moves inward. The projection of the incident beam is not changed, but the angle of incidence to the perturbed eye surface is changed. The reflected light centroid is shifted on the CCD array  105 , and the relative position of the beam on the detector increases. A similar situation occurs when the ocular pulse pressure is raised, except that the reflected light shift moves in the opposite direction. The curvature of the cornea and the change in the location of the eye surface normal with respect to the incident beam provide a 50 times magnification of the eye surface movement at an incident angle of 75 degrees with an equal incident and reflected beam drift length. The measurement of the eye surface motion can be explained from the single ray trace deflection diagram given in  FIG. 5  which shows two positions of the eye surface  104  during a single ocular pulse. At time t 0  the eye surface  104  is at (0,0) where the incident beam strikes the cornea. As the ocular pulse amplitude decreases the eye front surface  104  moves radially inward along the eye normal axis VA, moving to the second eye surface  104  position  104 P 2 . To measure this movement we place a small diameter, low divergence light source at the point (X0, Y0) aligned so that the center of the output beam intersects the initial eye surface  104 P 1  at (0,0). The angle A 0  between the input ray and the eye surface normal axis is the angle of incidence of the input beam. Physically the laws of reflection apply and the incident light ray is reflected along line (0,0)-(−X0, Y0) where it strikes the detector at the pixel located at the point (−X0, Y0) where the beam intensity profile is collected and the profile center of gravity calculated. At time (t o +δt) the eye surface  104  moves towards the eye center along the eye normal until it reaches the second eye surface position  104 P 2  where the incident beam intersects the surface at the point (XN, YN) and is reflected to the array detector at point (XS, YS) along a new normal and a different angle of incidence where another beam profile is collected and a center of gravity calculated. The distance between the two reflected beam detector profiles is proportional to the displacement of the two eye surface locations. 
       FIGS. 6 through 8  show the dependence of magnification on the angle of incidence, the pixelated signal displacement as a function of the eye surface motion and the fraction of beam reflected power as a function of the angle of incidence.  FIG. 6  shows the dependence of the reflective head sensor magnification on the angle of incidence A 0 . There is a strong dependence of magnification on the angle of incidence. Facial obstructions limit the maximum angle of incidence to 75 degrees or less with a useful range of 65 to 75 degrees corresponding to a magnification range of about 30 to 50 times.  FIG. 7  is a plot of the pixelated beam spot image displacement as a function of the eye surface movement for a 12.5 mm radius spheroidal eye ball with an 8.5 mm corneal radius and an incident beam angle of 75 degrees and a magnification of 50×. Note that a 50 micrometer movement of the eye surface as it responds to the ocular pulse causes a 450 micrometer displacement on CCD array  105 .  FIG. 8  shows the dependence of the energy reflected from the eye as a function of the beam incident angle for both incoherent and coherent light beams. The model results show that a coherent beam reflects more beam energy than an incoherent beam.  FIG. 8  also shows that for an incident angle of 75 degrees that the reflectivity of the coherent beam is 32% and that of the incoherent beam is 20%. The coherent beam reflected power is 1.6 times as high as the incoherent beam. The refracted beam powers are 68% and 80%. Since we wish to lower the beam refracted into the eye an incident angle of 75 degrees is desirable. 
       FIG. 9  is a plot of a FFT algorithm power spectrum calculated using a simulated ocular pulse amplitude wave train. Note that the plot shows four spectral power peaks the zero frequency power intensity plus three higher frequency components. The area under the zero frequency peak directly gives the mean energy in the eye ball that generates the IOP. The higher frequency spectral components are natural oscillations of the eye generated as the ocular pulse ebbs and flows without external drive or instrumental effects. Variations in these components may give data on the health of the eye and the cerebrovascular system. 
     Those of ordinary skill in the art will recognize that many modifications and variations of the present invention may be implemented without departing from the spirit or scope of the invention. Thus, it is intended that the present invention cover the modification and variations of this invention provided they come within the scope of the appended claims and their equivalents.