|Publication number||USRE42782 E1|
|Application number||US 10/876,200|
|Publication date||Oct 4, 2011|
|Filing date||Jun 23, 2004|
|Priority date||Oct 7, 1998|
|Also published as||CA2346704A1, CA2346704C, DE69943202D1, EP1119284A1, EP1119284A4, EP1119284B1, WO2000019885A1, WO2000019885A9|
|Publication number||10876200, 876200, US RE42782 E1, US RE42782E1, US-E1-RE42782, USRE42782 E1, USRE42782E1|
|Inventors||Vasyl Molebny, Ioannis Pallikaris, Youssef Wakil, Sergiy Molebny|
|Original Assignee||Tracey Technologies, Llc|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (26), Non-Patent Citations (23), Referenced by (2), Classifications (8), Legal Events (3)|
|External Links: USPTO, USPTO Assignment, Espacenet|
This is a continuation-in-part of PCT Application No. PCT/US99/23327, with an international filing date of Oct. 7, 1999, in which the US is a designated country and which claims priority to Ukranian Patent Application No. 98105286as a priority document, with a priority filing date of Oct. 7, 1998.
The present invention relates to medical ophthalmological equipment, more specifically, it relates to devices for measuring the refraction of the eye as a function of spatial pupil coordinates.
A method and a device for mapping the total refraction non-homogeneity of an eye are set forth in prior co-pending PCT Application No. PCT/US99/23327 that includes directing into the eye a narrow laser beam, its axis being parallel to the visual axis of the eye under investigation, scanning the beam over the eye aperture, receiving a portion of light scattered by the retina, analyzing the position of the laser spot projected on the retina, and reconstructing from the data a map of the total refraction of the eye. A discussion of the details of the total refraction non-homogeneity determination have been incorporated herein and set forth below.
In many applications, information on the contribution of other refractive components of the eye may be helpful or necessary, as, for example, for subsequent corrective surgery.
For the purpose of measuring the surface shape of a cornea, a method is known of projecting a regular structure or regular patterns, such as a pattern of concentric disks onto the cornea, analyzing the reflected light and reconstructing from the analyzed data the shape and therefore the refraction distribution caused by the cornea. It has been discovered by Applicants that marrying the techniques for analyzing retina-scattered light and for analyzing cornea reflected light may give very useful information on the contribution to the total eye refraction of such other refractive components of the cornea, and/or the eye lens, that has not heretofore been successfully accomplished.
Measuring devices are known, for the study of the refraction component of the optical system of the eye, which depend on spatial pupil coordinates. These include M. S. Smimov's Smirnov's device for measuring the wave aberration , Van den Brink's device for measuring the transverse aberration , N. M. Sergienko's device for measuring the physiological astigmatism , and a spatially resolved refractometer . The above devices, based on Scheiner's principle, involve point-by-point investigation over utilizing a number of optical techniques. However, in using all such devices the direct participation of the patient is needed in the preliminary aligning of the eye and in the aberration measurements.
Major disadvantages of the above measuring devices are their low accuracy and productivity, a prolonged measurement process resulting in the patient's fatigue, variations in accommodation, and eye movements while taking measurements, thereby increasing the aberration measurement errors.
More advanced measuring devices are known, which do not require the patient to act as a link in the “measurement chain”. These include a device for measuring the aberration by the Foucault's knife method , a device for measuring the wave aberration using Hartmann-Shack sensors [6-8], including measurements that incorporate adaptive optics completely compensating the wave aberration .
A common disadvantage of the measuring devices with a Hartmann-Shack sensor is the fixed field of view of the raster photoelectric analyzer of transverse aberrations due to the mechanically rigid construction of the lens raster and the invariable mutual spatial arrangement of the photosensitive elements of the charged coupled device or CCD camera. This results in a fixed configuration of grid sites at the pupil plane in which aberrations are measured, with no flexibility of reconfiguring these grid sites for more detailed measurements in separate zones of the pupil depending on their aberration properties.
Other disadvantages of existing devices include: they do not incorporate means for providing an accurate reproducible “linkage” of the patient's eye to the spatial co-ordinates of the measuring device; they do not incorporate a means for adjusting the accommodation of the patient's eye that is necessary for studying the dependence of aberrations on the accommodation characteristics; they are not capable of taking measurements on a dilated pupil without using medicines.
Refraction can also be measured using a spatially resolved objective autorefractometer as disclosed in U.S. Pat. No. 5,258,791 . This device provides spatially resolved refraction data using a closed measuring loop which includes a reference pattern and a measurement beam. In this device, an origin of coordinates of the detector coincides with the center of the fovea image and the detector functions as a zero-position sensor.
The spatially resolved objective autorefractometer disclosed in U.S. Pat. No. 5,258,791 preferably using laser ray tracing, has a number of substantial problems relating to performance in the following basic and auxiliary functions: preliminary alignment of the optical axis of the device relative to the visual axis of the eye; accommodation monitoring of the patient's eye; allocation of points within the pupil at which refraction is measured; and measurement of the angle of laser beam incidence into the patients eye. Respective to the above basic and auxiliary functions, the above drawbacks are inherent to the device disclosed in U.S. Pat. No. 5,258,791.
Preliminary alignment of the optical axis of prior art devices relative to the visual axis of the eye may be problematic for at least the following reasons: first, the visual axis of the eye is assumed to be the line passing through the geometric center of the pupil and the fovea. However, it is known that the geometric center of the pupil does not always coincide with the visual axis due to the misalignment of the pupil opening and the optical axis of the cornea and the crystalline lens. In addition, the pupil may not be symmetrical.
Second, in prior art devices in which alignment is dependant on fixation of the patient's gaze at a focal point, changes in the position of the point at which the patients gaze is fixed results in angular movement of the patient's eye which disturbs the previous alignment. Consequently, both points (on the pupil and on the retina) through which the center line passes do not have a definite location.
Third, the focal points in devices without ametropia compensation can clearly be observed only with an emmetropic or normal eye. When the patient's eye is ametropic, such devices will see a diffused laser beam spot whose width increases with the ametropy. It is obvious that under such conditions the gaze cannot be fixed accurately in a certain direction, which is another factor preventing an accurate alignment. Another drawback of prior art refractometers is that the fovea and the photosensitive surface of the photodetector are optically coupled by the lenses only in the emmetropic or normal eye. In the event of an ametropic eye, the decentering or defocusing of the fovea image on the above-mentioned surface of the photoelectric detector causes additional refraction measurement errors which are not compensated for. The present invention is designed to compensate for this.
Fourth, a sufficiently bright laser radiation may irritate the fovea to such a degree that the eye begins to narrow its pupil reflexly. Therefore, before performance of the eye centering procedure, medicines paralyzing the ciliar body muscles are likely to be required, which changes the refractive properties of the eye as compared with its normal natural state.
The need for accommodation monitoring of the patient's eye has not been satisfied in prior art devices. As a consequence, the patient's eye can be accommodating at any distance. It is known that the refractive properties of the eye depend on the accommodation distance. Because the accommodation is unknown to the operator, it is impossible to correlate the refraction map and the eye accommodation.
It has become apparent to the present applicants that a spatially resolved refractometer should preferably include a device for adjusting to the patient's eye accommodation.
Prior art devices using electromechanical actuators greatly reduce the possibility of ensuring a high-speed scanning of the pupil and the possibility of shortening the duration of the ocular refraction measurement process.
In prior art devices using a disc or movable aperture bearing planar surface to control laser targeting, the aperture occupies only a small portion of the zone in which the laser beam intersects the planar surface. Thus, only that portion of the laser beam which is equal to the ratio of the area of one refraction measurement zone on the pupil to the entire pupil area passes through the aperture. Such a vignetting of the laser beam results in an uneconomical use of laser radiation and should be considered a major drawback of such designs.
Drawbacks in the measurement of the angle of laser beam incidence at which it crosses the necessary measurement zone of the pupil and the center of the fovea are inherent to designs which do not provide a sufficiently high refraction measurement speed. In such designs, the time of measuring the refraction at 10 measurement points of the pupil is up to one minute. During this period the patient's eye can move up to 100 times and change its angular position due to natural tremor, “jumps” and drift.
Systematic instrument errors have plagued prior aberration refractometeis. Due to an irregular distribution of the light irradiance within the light spot on the retina, unequal photosensitivity across the surface photoelectric detector, time instability of the gain of preamplifiers connected to the photoelectric detector elements, and the presence of unsuppressed glares and background illumination the photodetector, the photodetector does not register a “zero” position of the spot on the fovea without systematic errors. Further, as a result of its own aberrations, the optical system providing for eye ray tracing contributes an angular aberration to the laser beam position. The present instrument incorporates structural elements which compensate for such errors and thus increase the refraction measurement accuracy.
What is needed is an improved electro-optical ray tracing aberration refractometer which makes it possible to achieve the following goals: flexibility of allocation of measuring points within the pupil and to pupil and improvement in the effectiveness of using lasers by reducing the vignetting of the laser beam at the aperture diaphragm; reduction in the duration of measuring refraction across the entire pupil to around 10-20 milliseconds; ensuring optical coupling of the photosensitive surface of the photodetector with the fovea even for ametropic eyes as well as for accommodation monitoring at any given distance; synchronously measuring both the total eye refraction aberrations and the component caused by cornea refraction characteristics maintaining incremental accuracy by interposing additional impingement points between impact points with wide variation in refraction characteristics; reduction in instrument errors when measuring aberration refraction; enhancement of the accuracy and definitiveness of instrument positioning relative to the patient's eye; the potential for automation of the positioning and controllability of the working distance between the patient's eye and the device components; and enablement of instrument positioning without medically dilating the pupil. The present invention provides the aforementioned solutions and innovations.
It has also been found that, due to noncoincidence of the points of primary information from the attempted combination of total eye refraction information and the surface shape of a cornea, and the followed approximation, any maps reconstructed from such cornea topography data and separately obtained total eye refraction aberration data would contain significant errors when reconstructing the differences. A workable device and method for synchronous mapping of the total refraction non-homogeneity of the eye and of the refraction components caused by the cornea would be desirable.
To avoid these and other drawbacks and solve the problems and produce a workable device and method for synchronous mapping of both the total refraction non-homogeneity of the eye and its refractive components, the following inventive technology has been devised. A device for synchronous measuring aberration refraction of an eye and calculation of the component of the total aberration refraction caused by the cornea has been accomplished with a device comprising a polarized light source producing a probing beam along a path to the eye. A beam shifter is provided to rapidly shift the probing beam to a plurality of spaced-apart, parallel paths for impacting the eye at a plurality of points on the cornea. Nonpolarized backscattered light from the retina of the eye is directed through a polarizing beam splitter onto a first position-sensitive detector by which the total eye refraction is determined for each of the plurality of impingement points. Synchronized with each impingement point determination of the total eye refraction, there is a polarized reflection from the impingement point on the cornea directed through the polarizing beam splitter and to a second position-sensitive detector for determining the refraction component caused by the cornea. A comparator can synchronously compare the total refraction for each impingement point to the component of refraction caused by the cornea and by subtraction can determine the component of refraction caused by the parts of the eye other than the cornea. This data can be placed in memory and/or can be supplied to a program device for map reconstruction and to a representative display of the refraction characteristics of the eye and its components.
For additionally improved resolution, a program device may be provided for inserting an additional measuring point or impingement point located between neighboring measuring points that produce a different value higher than a specified maximum difference threshold.
Thus, the object of the present invention is to provide an improved polarized light device for measuring the aberration refraction of the eye. The aberration refractometer allows estimates of the ametropy, astigmatism characteristics, simultaneously allows synchronous determination of component parts of total aberration refraction of the eye contributed by the cornea and therefore the components contributed by other portions of the eye and thereby allowing visual acuity and increased accuracy of calculations of the part of the cornea to be removed by photorefractive keratectomy  if necessary to correct eye refraction non-homogeneity or aberration.
In a preferred embodiment, the aberration refractometer comprises a light radiation source, preferably laser light or other polarized light; a telescopic system; a two-coordinate deflector consisting of two single-coordinate deflectors; a deflection angle control unit; an aperture stop; a field stop; a collimating lens; an interferential polarizing beam splitter; a first position-sensitive photodetector with an objective lens for detecting only the position of light backscattered from the retina; a second photosensitive photodetector with an objective for detecting only the position of polarized light reflected from the cornea; a comparator for point-by-point comparison of the retinal backscatter position and the cornea reflection position; and a data processing and display unit consisting of a computer, an analog-to-digital converter and a preamplifier. Use of laser or polarized light as a probing beam in combination with the interferential beam splitter allows the polarized light to be separated from the nonpolarized light reflected back from the retina and prevents it being detected by the first photodetector and further allows polarized light reflected from the cornea to be detected by the second photodetector without detecting the backscattered light from the retina.
The instrument of the present invention is able to reduce the time needed for measuring the refraction, eliminate light beam energy losses at the aperture stop and create a flexible system for locating the measurement points on the pupil by providing the following: the telescopic system is positioned in the probing beam path after the two-coordinate first deflector at a distance corresponding to the coincidence of the entrance pupil of the telescopic system and the gap or zone between the single-coordinate deflectors, the aperture slop or diaphragm is placed between the lenses of the telescopic system at the point of coincidence of their foci, and the field stop or diaphragm is positioned in the plane of the exit pupil of the telescopic system and, at the same time, at the location of the front focus point of the collimating lens situated in front of the interferential polarizing beam splitter at such a distance from the patient's eye which is approximately equal to the focal distance of the collimating lens.
To ensure a constant optical coupling of the photosensitive surface of the first photodetector and the retina for both emmetropic and ametropic eyes, a group of lenses with variable optical power is installed between the interferential polarizing beam splitter and the eye, said group of lenses having the function to adjustably form the retina image of an ametropic eye at infinity regardless of the emmetropic or ametropic condition of the eye. The photosensitive surface of the first photodetector is conjugated with the front focal plane of the objective lens, being inserted following the interferential polarizing beam splitter on the path of the light scattered by the retina.
To provide for fixation of the patient's line of sight along the optical axis of the instrument and to compensate for accommodation of the eye at the required distance while keeping constant optical conjugation of the patient's eye with the photosensitive surface of the detector, a second beam splitter or an optical axis rotation mirror, as well as a plate with a gaze fixing test pattern or a test-target for sight fixation are optically coupled with the photosensitive surface of the photodetector and are located between the photodetector and the objective lens. A second optical group of lenses, with variable negative optical power and which function to form an image for the patient's eye of the test-target at a distance corresponding to the preset accommodation, is positioned between the second beam splitter or optical axis bending mirror and the interferential polarizing beam splitter. When an optical axis bending mirror is used, it is mounted on a movable base making it possible to displace the mirror so as to enable the light radiation scattered by the retina to reach the photodetector during the measurement of the patient's eye characteristics.
In one embodiment, to account for systematic refraction measurement errors, a second mirror, for bending or redirecting the optical axis of the probing laser beam is inserted in the laser beam path after the last optical element before entering the patient's eye. Following the second mirror, an optical calibration unit for simulating an eye is inserted. The optical calibration unit includes an axially movable or stationary retina simulator whose optical characteristics are equivalent to those of the human retina. The second optical axis bending mirror is installed on a movable base so that it can be moved into the probing laser beam path during measurement with the optical calibration unit and moved out when measuring the patients eye refraction.
To align the instrument relative to the patient's eye as well as to enhance accuracy and enable automation of the aligning process, the instrument is provided with a third beam splitter to insert a channel for eye alignment verification of the instrument and the patient's eye. In a preferred embodiment, the co-axial verification channel comprises one or more point of light sources and a TV or electro-optic detecting device, together serving to display the pupil and/or eye image and providing a permission channel to measure eye characteristics when the optical axis of the instrument and the visual axis of the patient's eye coincide. To enable the instrument to be used without dilating the pupil with a medicine, a laser radiation source and/or infrared light sources are incorporated into the coaxial verification alignment mechanism. It is contemplated that the instrument can be used to make refraction measurements under conditions simulating either day or night light conditions.
In an alternate embodiments it is further contemplated that the alignment verification can be done under the control of the operator of the device or can be automated. In one embodiment, the co-axial verification channel provides either visual or acoustic notification that coincidence between the instrument optical axis and the patient's visual axis is proximate or near to “on target” status. Once this status is attained, the instrument is “armed” electronically. Once full coincidence is attained, the measurement controller automatically causes spatially defined parallel light beams, preferably laser beams, to be rapidly fired and enter the eye through the input channel. Light reflecting from the retina is directed to the retinal spot detecting channel for spatial and intensity characterization. This process can permit upwards of at least 5 replicate measurements over 65 spatial locations to be taken within 15 milliseconds without the need for the patient to actively participate in the targeting and alignment process.
For a more complete understanding of the present invention, including features and advantages, reference is now made to the detailed description of the invention along with the accompanying figures in which like numerals represent like elements and in which:
While the making and using of various embodiments and methods of the present invention are discussed in detail below, it should be appreciated that the present invention provides many applicable inventive concepts which may be employed in a variety of specific contexts. The specific embodiments discussed herein are merely illustrative of specific ways to make and use the invention and do not delimit the scope of the invention.
The total refractive eye aberration for a thin beam of light entering at impingement point 132 is determined by locating point 136 on the retina and determining the spatial position 138 of that illuminated point on the retina 136 relative to the fovea 134 aligned along central axis 121. This position may be indicated relative to central axis 121 by coordinates (dx1, dy1). Light impacting the retina is backscattered off the retina. The backscattering is not a reflection per se and therefore the entering beam 126b is depolarized by the retinal surface. The backscattered light, having its optical axis represented by path 126c, is non-polarized light. The light 126c is directed in beam splitter 128 along path 126d to a polarizing beam splitter 140. Beam splitter 140 is a polarizing beam splitter so that it reflects non-polarized light and allows polarized light to pass through it. Thus, beam 126d is directed along path 126e through a lens 142 that focuses it along path 126f to a first photo detector 144. As will be discussed more fully below with respect to
Returning now to
Synchronously with each of the polarized probing beams 126a and the plurality of additional beams 127a, the component of aberration caused by aberrations in the cornea surface 131 may simultaneously be determined. Because beams 126a and 127a are polarized light, they will partially reflect off of the cornea surface 131 as a polarized light beam 154a, in the case of probing beam 126a, at an impingement point 132 and as reflected polarized light 153a in the case of probing beam 127a and at impingement point 133. Because the light reflects from the cornea at an angle corresponding to the angular position of the anterior cornea surface 131 at the point of impingement 132, reflected beam 154a diverges from beam 126a, depending upon the refractive characteristics of the cornea 131 at the impingement point. Beam 154 is directed by beam splitter 128 to polarizing beam splitter 140 along path 154b. The beam 154b passes through polarizing beam splitter 140 because it is polarized light that was reflected from the cornea surface and travels along path 154c so that its position may be detected by a second photodetector 156. To facilitate determination of the reflective angle of beam 154 off of the cornea, the distance from the cornea to a semitransparent scattering screen 158 is a known quantity so that the offset distance 160 of the beam 154c impacting scattering screen 158 is indicative of the topography of the cornea surface 131.
The scattering screen 158 causes, the light beams 154c and 153c, to scatter, as schematically depicted with scattering diagrams 155 or 157. The position 160 or 159, with respect to the optical centerline 161, is imaged by a lens 162 onto the second photodetector 156. Once again, the second photodetector 156 may comprise an array of x and y photodetectors using a beam divider to determine the x-y position 160 for the reflected light from the cornea. This information is provided to a cornea cause refraction calculator 164.
The data from both total refraction calculator 146 and from he the cornea cause refraction calculator 164 is fed into a comparator 168 and also to memory 170. The comparator information produces data, including the total refraction for each point, the cornea cause refraction for each impingement, i.e., for each shifted probing beam and may also determine the component of the refraction aberration due to components of the eye other than the cornea. From this information, a map of refractive characteristics of the eye is reconstructed in a map reconstruction unit 172. The reconstruction map produced at 172 may be displayed at a display 174, such as a CRT screen or a color printout. All of the total refraction calculator 146, the cornea cause refraction calculator 164, the comparator 168, the memory 170, the map reconstruction unit 172 and the display 174 may be separately provided or alternatively may be included in a computer system and display screen and/or printer schematically represented by system dash lines 148 in
The semitransparent light scattering screen 158 may, for example, be milk glass or translucent fluorescent light cover material having a substantially homogeneous characteristics so that polarized light beams impacting at any point produce the same relative intensity and same relative diffusion by which the position of such light beams may be detected with position sensor 156. The second position sensor 156, although not depicted, may also be constructed similarly to position sensor 146 so that x component sensor array 88 and y sensor component array 89 are used in combination to get an x-y position sensor.
Light sources placed in front of the eye are used to align the visual axis of the eye with the optical axis of the instrument. Preferably a plurality of orthogonally placed light emitting diodes (LEDs) 102, for example emitting at a λ of 940 mm could be employed. Light produced by LEDs 102 is reflected off the cornea and imaged by camera 112. When the reflected light aligns with preset targeting parameters, the instrument is in the proper alignment and therefore in the permissive mode for firing of the spatially resolved parallel beams formed along channel 59.
The illuminated eye is then ultimately imaged by camera 112 as the image passes through the beam splitter prism 92 and is redirected at beam splitter 100 to pass through optical elements 104, 106, 108 and 110 to finally fall upon the CCD camera 112.
A retinal spot position detecting channel 99 is used to detect the position of reflected spots from the retina of eye 98 created by the input channel and includes a interferential polarization beam splitter 92 that directs non-polarized reflected light from the retina of eye 98 to a position sensor.
In one embodiment of a photodetection position sensor as shown in
Details of the embodiment depicted in
Light source 96 and condenser lenses 77 and 79 enable homogeneous irradiating of the linear arrays 88 and 89, thus checking their homogeneity at servicing. Light emitting diode 96 and condenser lenses 77, 79 form a wide beam for calibrating photodetectors 88 and 89. If any of the elements is out of tolerance, its output can be corrected at signal processing procedures.
A fixation target channel 85 preferably comprises a light source. In a preferred embodiment the light source is a green 565 λ LED 84. The light may be transmitted through lenses 74 and 75 and directed by prism 86 and through beam splitter 100 which has wavelength differentiating optical coatings. Fixation target is positioned on the optical element 106. The light beam from LED 84 passes through lenses 104 and 108 and fixation target 106 and is reflected off of the mirror 110. The fixation target light passes back through the lens 104 and is redirected by beam splitter 100 at 90 degrees out toward the eye for the patient to visualise the image as coming from the location of the surface 110 which image can be moved from near fixation to far fixation or adjustable anywhere in between and this may be used for changing the eye accommodation over a period of time and simultaneously taking a series of measurements including spatially resolved aberration refraction measurements as well as pictures on the CCD camera 112. This produces a time lapse imaging of the eye and measurements of the aberration refraction as it cycles through different fixation target distances. The different target fixation distances may be automatically moved or adjusted from near to far using electro mechanical adjustment means that may be synchronized with the measurements and/or images taken on a time lapse basis.
The instrument described herein was developed to provide a total aberration refractometer able to accurately and quickly provide a refractive map of either emmetropic or ametropic eyes without accommodation error.
An ametropia compensator is schematically depicted as a varifocal group of lenses 10 and 11, adjustable to compensate for the patient's eye ametropia. One of the lenses is mounted on a movable base connected to actuator drive 38. An accommodation controller is schematically depicted as lenses 16 and 17 that constitute a varifocal group of lenses for accommodation control of the patients eye.
An objective lens 18, at whose focal point the photosensitive surface of a position-sensitive photodetector 19 is located, is intended to form an image of the irradiated retina in the plane of the photosensitive surface of the position-sensitive photodetector. The photosensitive elements of the photodetector are connected through a preamplifier 22 and an analog-to-digit converter 23 to a computer 24. A beam coupler 39 is movably mounted between the objective lens 18 and the photodetector 19 to optically conjugate the plane of the test-target or plate 20 with the photosensitive surface of the first photodetector 19 as well as with the fovea surface. The plate 20 is needed to ensure the fixation of the patient's gaze. Located behind the plate 20 is a light source or radiator 21 serving to illuminate the plate.
Elements 25 through 30 comprise a microscope whose objective lens consists of lenses 25 and 27 together with mirror 26. A plate 29 with first coordinate-grid is preferably located at the back focal plane of a lens 27. A lens or a group of lenses 30, the front focal point of which coincides with the back focal point of the lens 27, comprises an eyepiece of the microscope. The beam splitter 28 serves to optically couple the retinal plane with the photosensitive plane of a TV camera 32 connected to the computer through a video signal conversion and input board, alternatively termed a frame grabber board, 33.
By means of a mirror 12 provided with an opening, the optical axis of the microscope is aligned with the optical axes of the ray tracing channel (elements 1-11) and the photoelectric arrangement for measuring the transverse ray aberration on the retina (elements 16-19).
In a preferred embodiment, four light-emitting diodes (LEDs) 14 are installed in a cross-wise configuration in front of the patient's eye. Each LED is preferably located in the same plane as each other LED, at an equal distance from the optical axis and perpendicular with the axis. The microscope and the LEDs comprise a system for the visual and television positioning of the instrument relative to the patient's eye. The microscope is installed so that the front focal plane of lens 25 coincides with the plane, where imaginary or virtual images of the LEDs 14, mirrored by the anterior corneal surface, are located.
Before the total refraction measurement process is commenced, the instrument is positioned relative to the patient's eye and the instrument is calibrated using the optical calibration unit 34-36. Movably mounted between the lens 11 and the LEDs 14 is a mirror 13 which serves to join the optical axes of the instrument and the optical calibration unit 34-36. In one preferred embodiment of an optical calibration unit, it comprises a meniscus or cornea simulator 34, liquid medium or vitreous simulator 35, and retina simulator 36. The retina simulator 36 is preferably movably mounted so that it can be moved along the optical axis by means of actuator or drive 37.
The instant measuring instrument incorporates a computer 24 or like device for controlling the acousto-optic deflector 4, analog-to-digital converter 23, and actuators or drives 37 and 38. The computer 24 or like device or devices may perform additional duties including, for example, mathematical processing and data storage, calculation and display of aberration parameters and the ocular refraction characteristics as well as provide setting measurement modes and implementation of automatic instrument alignment.
The instrument for measurement of the total eye refraction, in its preferred embodiment, functions in the following way. The light beam emitted, for example by laser 1, is expanded, collimated and directed to the acousto-optic deflector 4, which changes its angular position in accordance with the corresponding computer program. The telescopic narrower 5 and 6 reduces the beam thickness to the requisite magnitude. The center of the stop or diaphragm 7 is a point of angular “swinging” of the beam exiting from the telescopic narrower. Due to its positioning in the front focal plane of the lens 6, the aperture stop or diaphragm AD has its image in the back focal plane of the lens 8 which is aligned with the eye pupil. Further, because the stop or diaphragm 7 is positioned in the front focal plane of the collimating lens 8, angular swinging of the laser beam with the angle vertex located on the stop or diaphragm 7 is converted into parallel shifting of its optical axis after passing the lens 8.
If the patient's eye is ametropic, the axial movement of the lens 10 (or 11) converts the telocentric beam into a beam which diverges (in the case of myopia), or converges (in the event of hyperopia), so that the image of the diaphragm 7 is optically conjugated with the retina. This also ensures parallelism of the rays reflected by the retina in the zone in front of the beam splitter 9, which is necessary for its normal functioning.
The light entering the eye 15 of the patient is polarized in the plane shown in
Lenses 16 and 17 and the objective lens 18 produce an image of the illuminated area of the retina in the plane of the first photodetector 19. In
In one more embodiment, presented in
Still another embodiment of the invention, schematically shown in
In the various embodiments of
If photodetector 19 is a four-quadrant photodiode, as, for example, that shown diagrammatically in
where β is the transverse magnification in the plane of the first photodetector as related to the plane of the retina, b is a coefficient depending on the size of the light spot in the plane of the photodetector, and U1, U2, U3 and U4 are the photoelectric signals coming from the corresponding photodetector channels.
If photodetector 19 is a lateral position sensing detector, as shown in
where β is the transverse magnification between the planes of photodetector and retina, U1, U2, U3 and U4 are the signals coming from the electrodes, 1, 2, 3 and 4 correspondingly, and a is a scaling coefficient depending on the electrical parameters of the lateral detector.
The principle of operation relating to the positioning the instrument in relation to the patient's eye is illustrated in
As can be seen from
Taking into account that the largest contribution to the optical power of the eye is made by the anterior surface of the cornea, the visual axis line is assumed to be the line passing through the fovea center and the vertex of center of curvature of the front surface of the cornea. If the radiator B1 is positioned in front of the patient's eye, then, due to reflection of the light from the anterior or front surface of the cornea, this surface functioning as a convex mirror, forms an imaginary or virtual image B′1 of the radiator, located symmetrically to the axis in accordance with the laws of geometric optics.
When several radiators, such as for example, B1 and B2, are positioned in front of the patient's eye symmetrically to the optical axis of the instrument (
Thus, to make the optical axis of the instrument and the visual axis of the eye coincide, two conditions must be satisfied: the patient's gaze is fixed on the point A and the images B″1 and B″2 are centrally positioned in relation to the axis of the objective lens 52. The positioning can be checked using the coordinate grid provided on the plate 29 (
The coincidence of the points B″1 and B″2 with the surface or plane 54 is indicative of setting the fixed working distance between the instrument and the eye which is the result of the focusing of the images B″1 and B″2 on the surface 54.
The point of gaze fixation is created by locating the mirror 39 (
Another embodiment of eye instrument alignment can be implemented using manually or automatically operated measurement of the pupil edges; forming a figure, approximately a circle. Its center does not coincide usually with the center of symmetry of four reflexes, two of which B″1 and B″2 are shown in
The calibration of the instant aberration refraction instrument may be effected using the optical calibration unit. The optical calibration unit can be made to incorporate known aberrations at the corresponding cornea simulator 34 measurement points. For example, the aberration may be determined by the computer using special optical design programs. If, for example, the front surface of the lens 34 is ellipsoidal, then the aberration refraction at all the points of the pupil is equal to zero.
When an ametropy compensator is used, nonparallel laser beams will enter the optical calibration unit. This will result in a standard aberration of defocusing; to compensate for this aberration, the retina simulator can be moved along the optical axis by means of the actuator 37 to the focus point. Thereby, the fovea can be optically conjugated with the retina simulator.
Systematic errors of measurements of the transverse aberration will be evidenced by the deviation of the measurement results from the estimated data. Such determinable systematic errors can be taken into account when measuring actual total ocular aberrations.
The calibration by comparison with the optical calibration unit is preferably performed automatically before measuring the ocular aberrations by locating the mirror 13 on the optical axis of the instrument.
Prior to the ray tracing of the patient's eye the mirrors 13 and 39 are withdrawn from the light path entering the eye and then the light passes to the photodetector. The aberration displacement of the image of the light spot on the fovea is measured at a set of points on the cornea corresponding to an ocular ray tracing grid chosen by the operator. An example of a grid or an allocation of measurement points on the pupil is shown in
The data on measurement of the transverse aberrations on the retina δx (ρ, φ) and δy (ρ, φ) are used for further calculations of the coefficients of the Zernike polynomials by means of the least squares method in order to approximate the function of the total wave aberration of the eye. The wave aberration function is then used to calculate the local total refraction at any point of the pupil. In addition, the approximation makes it possible to determine or reconstruct the nature of local aberration refraction in that small axial zone of the pupil, where it is impossible make accurate direct measurement of refraction.
In one experiment conducted using this instrument in which five replicate tests were performed and the results averaged, the laser beam total aberration on the retina at 65 points of the pupil was been performed in within 12 milliseconds with no more that 5 mW of light radiation entering the eye.
The extremely fast measurement permits the computer control program to cause a plurality or spatially resolved aberration measurements to be made in a very short period of time. The control program in one embodiment automatically activates a plurality of measurements coordinated with a series of adjusted accommodation fixation distances and automatic determination of proper eye alignment to receive a series of data measurements from the retinal spot position detecting channel. A series of refraction measurements for a dynamic eye refraction system is produced. Spatially resolved refraction measurements can be automatically programed and automatically made during a variety of dynamic changes such as varying accommodation or during normal functioning of the eye under a variety of predetermined conditions and internal or external changing conditions.
While this invention has been described with reference to illustrative embodiments, this description is not intended to be construed in a limiting sense. Various modifications and combinations of illustrative embodiments, as well as other embodiments of the invention, will be apparent to persons skilled in the art upon reference to the description. It is therefore intended that the appended claims encompass such modifications and enhancements.
Other alterations and modifications of the invention will likewise become apparent to those of ordinary skill in the art upon reading the present disclosure, and it is intended that the scope of the invention disclosed herein be limited only by the broadest interpretation of the appended claims to which the inventors are legally entitled.
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|International Classification||A61B3/10, A61B3/107, A61F9/007, A61B3/103|
|European Classification||A61B3/10F, A61B3/103|
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