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BACKGROUND OF THE INVENTION
1. Field of the Invention
This invention relates to an ophthalmological measurement method and
apparatus, and more particularly to an ophthalmological measurement method
and apparatus in which the eye fundus is illuminated by a laser beam
having a predetermined diameter and motion of a laser speckle pattern
formed by laser light scattered and reflected from the eye fundus is
detected at an observation point as fluctuations in the speckle light
intensity to produce a speckle signal which is evaluated for
ophthalmological measurement.
2. Description of the Prior Art
Various conventional methods are used for ophthalmological measurement.
These conventional methods include illuminating the eye fundus with a
laser beam, detecting the light scattered by the eye fundus and analyzing
and evaluating this light. There are, for example, laser Doppler methods
for measuring blood flow in retinal and other tissue described in
"Investigative Ophthalmology," vol. 11 No. 11, page 936 (November 1972)
and "Science," vol.186 (November 1974) page 830, and in Japanese
Unexamined Patent Publication Nos. 55-75668, 55-75669, 55-75670, 52-142885
(corresponding to GB 13132/76 and U.S. Pat. No. 4,166,695), 56-125033
(corresponding to GB 79/37799), 58-118730 (corresponding to U.S. Pat. No.
4,402,601) and U.S. Pat. No. 4,142,796. However, these laser Doppler
methods involve the use of a high precision optical system, are
complicated to use and provide results which lack repeatability and
reliability, all of which make practical application difficult.
It is known that when a laser beam strikes an object which causes diffusion
or scattering of the beam, the light scattering from the object gives rise
to a speckle pattern caused by interference between reflected rays of the
coherent light. The laser speckle method utilizes this to evaluate the
state of tissues in the eye fundus. Examples of this method are described
in Japanese Unexamined Patent Publication Nos. 62-275431 (U.S. Pat. No.
4,734,107 and EPC 234869), 63-238843 (EPC 284248) and 63-242220 (EPC
285314).
These publications describe the use of a detecting aperture to extract
time-base fluctuations in the intensity of speckles formed at an optical
Fourier Transform plane with respect to the eye fundus, or at the
Fraunhofer refraction plane, or at an image plane (or a magnified image
plane) that is conjugate with respect to the eye fundus, and the blood
flow state is determined by an evaluation of the speckle signal thus
obtained.
A major obstacle to the clinical application of the above systems has been
their susceptibility to the effects of movement, such as movement of the
subject's eye, vibration and the like. This frequently causes unwanted
movement of speckle patterns on the detection plane, thus throwing the
detecting aperture and laser beam out of alignment during measurement. One
way to overcome this is described in the laser-Doppler method of Japanese
Patent Publication No. 56-125033. This involves the mechanical scanning
the eye fundus image on the detection plane and using differences between
the light reflectance of the walls of a blood vessel and that of other
areas of tissue to distinguish blood vessels, and correcting for
positional deviation. A drawback of this method is that it requires a
mechanism for the mechanical scanning of the eye fundus image, which makes
the apparatus too large and complex to be practical.
Another method, described in Applied Optics, Vol. 27, No. 6, page 1113
(Mar. 15, 1988) and in Japanese Patent Publication No. 63-288133 (U.S.
Pat. No. 014,994), shows the feasibility of an image scanning arrangement
which allows blood vessels to be distinguished and tracked automatically.
However, the method is based on the wavelength dependency of reflected
light and relies for its implementation on a plurality of laser beams of
different wavelengths which are projected in sequence. Again, this makes
the apparatus complex, impractical and costly. A further drawback is that
when corneal reflection is used to detect eye movement, the detection
precision is not high enough for the purposes of correcting for movement
by the blood vessel.
Conventional tracking methods involving the detection of eye movement
include one in which the corneal surface is illuminated by a laser beam
and movement of the reflected light is used to detect and track such eye
movement another method uses differences between two images of the eye
fundus obtained by a TV camera or other such imaging means.
However, such methods involve detection of eye surface movement and are
only able to provide a low level of intraocular tracking precision.
Moreover, eye fundus images obtained via a TV camera usually suffer from a
poor S/N ratio because the amount of available light is insufficient,
further the apparatus required to detect movement based on differences
between two images is large and complex.
On the other hand, the speckle pattern moves as the scattering object
moves, so that it is proposed to detect its movement as a fluctuation in
the light intensity at the observation point to obtain the difference of
the traveling speed of the object depending on the signal intensity.
To discriminate the blood vessel and measure the diameter of the blood
vessel, there has been proposed a method in which the eye fundus is
photographed using a fundus camera to measure the diameter of the blood
vessel on the basis of the photographed eye fundus or a method in which a
television camera is used to take a picture of the eye fundus and the eye
fundus image is subjected to an image processing (for example, image
sampling, A/D converting, sharping, masking, filtering) to determine the
diameter of the blood vessel.
Such conventional methods require a long time to obtain measurement results
because the eye fundus must be photographed, thus making it impossible to
make real time measurement of the diameter of the blood vessel.
Furthermore, the eye fundus image taken by the television camera is
usually underexposed and has a poor S/N ratio. This necessitates
complicated image processing and results in a bulky and expensive
apparatus.
SUMMARY OF THE INVENTION
It is therefore an object of the invention to provided an improved
ophthalmological measurement method and apparatus employing the laser
speckle phenomenon which is simple and straightforward in construction and
is able to detect eye movement and automatically track the movement in the
eye fundus with good accuracy.
It is another object of the invention to provided an improved
ophthalmological measurement method and apparatus employing the laser
speckle phenomenon which is simple and straightforward in construction and
is able to accurately measure the diameter of the blood vessel.
The invention provides an ophthalmological measurement method and apparatus
in which the eye fundus is illuminated by a laser beam projected by
projecting means and having a predetermined diameter. Motion of a laser
speckle pattern formed by laser light scattered and reflected from the eye
fundus is detected at an observation point as fluctuations in the speckle
light intensity to produce a speckle signal which is evaluated by
evaluating means for ophthalmological measurement. In this arrangement the
speckle signal is evaluated to discriminate differences in travel
velocities of blood cells in the eye fundus to identify a blood vessel of
the eye fundus.
Any movement of the identified blood vessel of the eye fundus is detected,
and the position of the region illuminated by the laser beam and the
position of the observation point are adjusted by positioning means by an
amount corresponding to the amount of blood vessel movement to track the
blood vessel automatically. Furthermore, both edges of the identified
blood vessel are located to determine the diameter thereof.
In such an arrangement, the laser beam of predetermined diameter is
projected into the eye fundus by a laser beam projector and the movement
of a speckle pattern formed by diffused light scattered by blood cells
within the eye tissue passes through a light receiving system and is
detected by a photosensor as fluctuation in speckle light intensity. The
speckle signal mirrors the travel speed of the blood cells in the eye
tissues. The size of speckles on the photosensor and the scanning speed of
the photosensor are optimally set. The speckle light intensity will
fluctuate more rapidly when cell velocities are high, and the averaging
effect of the photosensor's storage time will result in a smaller output.
Conversely, a low cell travel speed will decrease the lowering of the
output from the photosensor. The differences in signal intensity thus
produced are used to distinguish blood vessels. Movable mirrors are driven
by an amount corresponding to shifts in the position of the blood vessel
caused, for example, by eye movement, so that the position of the region
illuminated by the laser beam and the observation position are controlled
to automatically track the blood vessel. Furthermore, both edges of the
identified blood vessel are located to determine the diameter of the blood
vessel. Thus, the invention provides an improved ophthalmological
measurement method and apparatus which is able to detect eye movement and
automatically track the movement in the eye fundus or measure the diameter
of the blood vessel with a simplified structure and with good accuracy.
BRIEF DESCRIPTION OF THE DRAWINGS
The objects and features of the present invention will become more apparent
from a consideration of the following detailed description taken in
conjunction with the accompanying drawings in which:
FIG. 1 is a diagram showing the structure of a first embodiment of an
apparatus according to the present invention;
FIG. 2 is a diagram for explaining the structure of a ring slit;
FIG. 3 is a characteristic curve showing the characteristics of a filter;
FIGS. 4 and 5 show observed images of the eye fundus;
FIG. 6 is a block diagram of a signal processor used in the embodiment;
FIG. 7 shows the waveform of the signal output of an absolute value
circuit;
FIG. 8 shows the waveform of the signal output of amplifier with limiter of
the embodiment;
FIG. 9 is a flow chart of the control process for finding a blood vessel;
FIG. 10 is a flow chart of the control process for tracking a blood vessel;
FIG. 11 is a flow chart of the control process for central position
correction;
FIGS. 12a to 12f are diagrams showing the relationship between speckle size
and CCD pixel size, and output signals;
FIGS. 13a and 13b are graphs showing speckle pattern travel speed and the
waveform of a CCD output signal;
FIG. 14 is a schematic view of another embodiment of the apparatus of the
invention;
FIG. 15 shows details of a movable mirror;
FIGS. 16 and 17 are schematic views of a signal processor;
FIGS. 18a to 18d show waveforms of CCD output signals;
FIG. 19 shows the arrangement of an image rotator; and
FIGS. 20 and 21 show an arrangement for oscillating an image on the CCD.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
The invention will now be described in detail with reference to embodiments
shown in the drawings.
The invention is particularly used for an ophthalmological measurement
apparatus in which the eye fundus is illuminated by a laser beam having a
prescribed diameter and motion of a laser speckle pattern formed by laser
light scattered and reflected from the eye fundus is detected at an
observation point as fluctuations in the speckle light intensity to
produce a speckle signal which is evaluated to measure a blood flow state
in tissues in the eye fundus. Therefore, the embodiments described below
are those which are applied to the ophthalmological measurement apparatus
including a basic optical arrangement of an eye fundus camera to measure
the blood flow state in the eye fundus tissue. The invention is, however,
not limited to such embodiments and may be applied to other type of
ophthalmological apparatus.
With reference to FIG. 1, a laser beam from a red-light He-Ne (wavelength:
632.8 nm) laser light source 1, for example, passes through a condenser
lens 2 and a light quantity adjustment filter 3 for adjusting the beam
intensity, and is then collimated by a collimator lens 4. Two apertures 5
and 6 are provided within the path of the beam for selectively adjusting
the size and shape of the region of an eye fundus 16b of a subject's eye
16 which is illuminated by the laser beam. The laser beam passes from the
two aperatures 5 and 6 through lens 7 and is reflected by swingable mirror
8.
The laser beam passes through a condenser lens 9 and is reflected by a
mirror 10 provided in a transparent portion of an annular aperture 11a
formed in a ring slit 11 arranged in an eye fundus camera illuminating
projector, as shown in FIG. 2 (in which the non-transparent portion is
indicated by shading). Such an arrangement enables the laser beam to
direct along the same optical path to the eye fundus as that followed by
the beam of light projected into the eye fundus to provide illumination
for photography and observation. The laser beam thus passes through relay
lenses 12 and 13, is reflected by a ring mirror 14 and, via an objective
lens 15, passes through the cornea 16a of the eye under examination 16 to
the eye fundus 16b where the blood vessel of interest is irradiated with
the laser beam for measurement and tracking.
The swingable mirror 8 is provided in the optical laser beam illumination
system to deflect the laser beam spot in the eye fundus 16b. Prior to the
start of measurement, this deflection is performed via an output section
46 using means such as a trackball 17. The swingable mirror 8 can be
controlled by an ordinary method such as a coagulator arrangement which
allows independent control of the angle of mirror deflection in the x and
y directions relative to the optical axis.
To minimize the discrepancy that has to be corrected arising from
differences in laser beam deflection angles in the x and y directions, the
angle at which the laser beam is reflected by the swingable mirror 8 is
made as small as space will permit. The swingable mirror 8 is disposed at
a position that is substantially a conjugate of the cornea 16a or pupil of
the eye. This assures that the laser beam can be moved over the eye fundus
without any major change in the position of beam incidence on the cornea.
The laser beam is provided on the same optical path as the photography and
observation light beam. This arrangement is highly convenient since it
enables the location within the eye fundus 16b at which the laser beam is
being projected by the swingable mirror 8 to be brought within the field
of view for photography or observation by using mechanisms for swinging
and tilting the eye fundus camera vertically and horizontally and the eye
fixation means.
This measurement and tracking region is also illuminated by an illuminating
projector of the fundus camera to facilitate observation. The system for
providing the illumination for observation includes an observation light
source 18, a condenser lens 19, a condenser lens 21, a filter 22 and a
mirror 23 disposed on the same light path as a photographic light source
20.
The filter 22 disposed between the condenser lens 21 and the mirror 23 is a
wavelength separation filter having the type of characteristics shown in
FIG. 3 to filter out red components from the observation and photographic
light. A filter is selected that has spectral characteristics appropriate
to the wavelength of the laser beam source that is employed.
Speckle light produced by the scattering of the laser beam in the eye
fundus and reflected observation and photographic light passes through the
objective lens 15, the ring mirror 14, a focusing lens 24, an imaging lens
25 or 26 and a relay lens 29, is reflected by a movable mirror 30 and
passes through a relay lens 31 and is thereby formed into an image at a
ring mirror 32. The light reflected by the ring mirror 32 passes through a
relay lens 33 and is divided by a wavelength separation mirror 34.
Cylindrical imaging lenses 42a and 42b form speckle light reflected by the
wavelength separation mirror 34 into an image on a scanning type sensor
CCD 43. The wavelength separation mirror 34 is set at an angle of about 45
degrees relative to the optical axis and as the wavelength separation
mirror 34 has the same kind of spectral characteristics as wavelength
separation filter 22, shown in FIG. 3, it reflects most of the speckle
light produced by the red He-Ne laser beam.
Light that is transmitted by the wavelength separation mirror 34 passes
through an imaging lens 35 and forms an image at a reticle 36. The
examiner can view this image through an eyepiece 37. The eyepiece 37 can
be adjusted to compensate for individual differences in visual acuity. The
reticle 36 is used as a reference for such adjustments.
With reference to FIG. 4, the lines of the reticle 36 which intersect at
right-angles can be differentiated, and the intersecting portion coincides
with the center of an aperture 32a in the ring mirror 32. The reticle 36
can be rotated about the intersecting portion. Rotation of the reticle 36
to align it with a blood vessel 16c, as shown in FIG. 4, produces a
synchronous rotation of the cylindrical imaging lenses 42a and 42b and the
CCD 43, automatically orienting the CCD 43 perpendicularly to the image of
the blood vessel. FIG. 5 illustrates the eye fundus image that will thus
be formed on the face of the CCD 43. In the drawing, la denotes the area
illuminated by the laser beam.
Because the diameter of speckles, the boiling motion of the speckle pattern
and the sensitivity of the CCD 43, the cylindrical imaging lenses 42a and
42b are set so that the image of the eye fundus is formed on the CCD 43
with a lower magnification when it is in a direction parallel to the blood
vessel 16c than when it is orthogonal to the blood vessel. As shown in
FIG. 5, CCD 43 is provided at a position at which the image of the
aperture 32a of the ring mirror 32 does not cross the face of the CCD 43,
and the CCD 43 is arranged perpendicularly to the blood vessel 16c of
interest.
For photography purposes a swingable mirror 27 is pivoted about a point 27a
in the direction indicated by the arrow to raise it to a position 27',
whereby the observation and photographic light including speckle light
from the eye fundus that is reflected by the swingable mirror 27 and forms
an image which is photographed on photographic film 28. Thus, the system
can be used for observation and photography of the eye fundus like an
ordinary fundus camera. The ability to observe and photograph the eye
fundus when it is being illuminated by the laser beam is desirable, as it
enables the point of measurement to be directly confirmed and filmed.
In a system for receiving speckle light from the eye fundus and reflected
light for observation and photography, light passing through the aperture
32a of the ring mirror 32 forms an image of the eye fundus 16b at a
pinhole aperture 38. The light from the pinhole aperture 38 passes through
an interference filter 39 and, when measurement is started, is received by
a photomultiplier 40 which outputs a speckle signal to an analysis section
41. The interference filter 39 blocks light having a wavelength other than
the 632.8 nm red light produced by the He-Ne laser.
The swingable mirror 30 is provided in the system for receiving speckle
light from the eye fundus and light for observation and photography for
positional correction purposes so that the image of the blood vessel in
the eye fundus 16b is formed at the pinhole aperture 38 after passing
through the ring mirror 32. Prior to the start of measurement, this
adjustment is effected via the output section 46 using a means such as a
trackball 17.
As described above, the trackball 17 is also used for operating the
swingable mirror 8 prior to the measurement. A switch or other such means
may be provided to switch trackball control between the swingable mirror 8
and the swingable mirror 30. The swingable mirror 30 can be controlled by
any ordinary means which allows independent control of the angle of mirror
deflection in the x and y directions relative to the optical axis. This
applies also to the swingable mirror 8.
To minimize the discrepancy that has to be corrected arising from
differences in laser beam deflection angles in the x and y directions, the
angle at which the laser beam is reflected by the swingable mirror 30 is
made as small as space will permit.
By locating the swingable mirror 30 at a position that is substantially a
conjugate of the cornea 16a or pupil of the eye, the mirror 30 can be
deflected to move the eye fundus 16b image at the pinhole aperture 38
without the beam being blocked by the pupil or other portion of the eye.
In the light receiving system the imaging lens 25 is a wide angle type,
wide enough to provide a view which allows all of the image of the eye
fundus 16b to be checked. The imaging lens 26 is a narrow angle type with
a high magnification factor which provides a magnified image to facilitate
alignment of the blood vessel image in the area illuminated by the laser
beam with the pinhole aperture 38.
The imaging lenses 25 and 26 are arranged so that they can be switched
instantaneously without moving the optical axis. This variable power lens
arrangement facilitates accurate beam alignment with the required
measurement position.
The diameter of the ring mirror 32 is just large enough to allow the
passage of the light beam from the blood vessel 16c of interest, and the
ring mirror 32 is located at a position that is substantially a conjugate
of the eye fundus 16b. This assures that the examiner can align the system
accurately by manipulating the image of the blood vessel of interest so
that the image overlays the aperture of the ring mirror 32. FIG. 4 shows
the image that this will produce. As the wavelength separation mirror 34
passes a small amount of speckle light, it is possible for the examiner to
confirm the position of the illuminated area 1a.
When measurement is started, speckle light is received by the CCD 43 which
outputs a signal to a signal processor 44. The signal processor 44
produces a blood vessel discrimination signal which is converted to a
digital signal and output. If the blood vessel has moved because of
movement of the eyeball, for example, the amount of this movement is
detected from the digital blood vessel discrimination signal by an
arithmetic unit 45 which computes a correction amount for compensation for
the movement of the blood vessel. The computation result is output to the
output section 46 which uses feedback correction to control the swingable
mirror 30 and swingable mirror 8 so that the image of the eye fundus is
constantly maintained at the same position at the pinhole aperture 38 and
the laser beam continues to illuminate the same region in the eye fundus
16b.
The arithmetic unit 45 further serves to distinguish the blood vessel parts
on the basis of the blood vessel discrimination signal and to calculate
the blood vessel diameter. After calculation the results are output to the
output section 46, which then displays the blood vessel diameter on a
display.
Observation and photography light (other than red component light) together
with the small amount of speckle light is transmitted by the wavelength
separation mirror 34 and forms an image of the eye fundus at the reticle
36 also during the measurement process, and can therefore be observed by
the examiner. The ability to thus observe the eye fundus during blood flow
measurement is highly effective for preventing errors, as it enables any
deviation from the area of interest to be observed.
The electrical system from the signal processor 44 onwards will now be
described. FIG. 6 is a schematic diagram of the signal processor. With
reference to the drawing, the signal processor 44 is constituted of a
drive circuit 56, a high-pass filter 51, an amplifier 52, an absolute
value circuit 53, an amplifier with limiter 54 and an A/D converter 55.
Drive pulses generated by the drive circuit 56 are input to a 1,024-pixel
linear CCD 43. The CCD 43 converts to the speckle light to obtain a
speckle signal which is passed through the high-pass filter 51 to extract
just the high frequency components. This high frequency component signal
is then amplified by the amplifier 52 and passed through the absolute
value circuit 53 to obtain an absolute value.
The output signal thus obtained from the absolute value circuit 53 is
illustrated in FIG. 7. The signal waveform shown is only that obtained
from the central area of the CCD, not the whole area of the CCD; this also
applies to FIGS. 8, 12 and 13. The signal is then input to the amplifier
with limiter 54 to extract a blood vessel discrimination signal by
selectively amplifying the required portions such as the portion A shown
between the dotted lines in FIG. 7, the other, unnecessary parts being cut
off by the limiter. The signal output by the amplifier 54 is illustrated
in FIG. 8. The blood vessel discrimination signal thus obtained is
converted to digital form by the A/D converter 55 and input to the
arithmetic unit 45.
In the arithmetic unit 45 the digital signal data is stored in memory.
Assuming that the CCD is composed of n pixels so that n data elements are
stored in memory, and an address is assigned to each data element read out
of memory, the address of the first data element to be read out would be 0
and that of the nth data element would be n-1. Data from the blood vessel
portion will have a high value and data from portions other than the blood
vessel will have a low value. To simplify the explanation, data obtained
from the blood vessel will be assigned a value of 1 and data obtained from
other locations will be assigned a value of 0. While read-out data from
the blood vessel will have a value of 1, the effect of the speckles will
be that data which is not from the blood vessel will sometimes have a
value of 1 and sometimes 0.
A method of reducing the effect of speckles will now be described with
reference to the blood vessel search procedure illustrated by the flow
chart of FIG. 9.
In block B1 the reference position is set for the blood vessel search.
During the first search the examiner will have aligned the system
beforehand so that the blood vessel crosses the center part of the CCD 43.
This means that during the first measurement data from the center of the
CCD will have a value of 1, signifying a blood vessel. However, starting
with the second search, movement of the blood vessel will gradually shift
it from the center of the CCD, making it necessary to move the reference
point to the center of the CCD. In block B1 step S1 it is determined
whether a search is a first search or a second or subsequent search. If it
is determined that it is a first search, in step S2 the reference point is
set to the center address of the CCD; if it is a second or subsequent
search, in step S3 the reference point is set to a position midway between
the edges obtained in the preceding step.
The edges are established by blocks B2 and B3. If the left edge is searched
for in block B2, the right edge will be searched for in block B3, and
vice-versa. In step S4 the data corresponding to the address is read out,
and in step S5 it is determined whether the data has a value of 1 or 0.
Since, on the basis of block B1, it can be reliably assumed that the
reference address is over the blood vessel, in step S6 the address is
decremented and the point at which the data value first changes from 1 to
0 is determined as edge 1 of the blood vessel. In step S7 the position of
edge 1 of the blood vessel is detected and stored as blood vessel edge
data.
In steps S9 to S12 the same process is used to detect the position of edge
2 and stored the information as blood vessel edge data. In the present
method, the edges of the blood vessel are searched for starting from the
center of the blood vessel, so that the number of data elements that need
to be read out is greatly reduced together with the effect of speckles,
since it is only necessary to read out data corresponding to the portion
of the diameter of the blood vessel concerned. This enables the
determination process to be carried out quickly and reliably, as compared
with the method in which searches proceed sequentially from address 0 to
n-1 and the data has to be examined each time to ascertain whether it
represents speckles or the presence of a blood vessel.
The diameter of the thus identified blood vessel can be determined by
multiplying the width between both edges of the blood vessel with a
coefficient determined by the magnification of the light receiving system.
It is preferable to obtain the width of the identified blood vessel
several times and to derive therefrom an average value or the smallest of
the measured widths of the blood vessel for improvement in determining the
blood vessel diameter.
A plurality of positional information is required if the amount by which
the blood vessel has shifted is to be obtained just on the basis of blood
vessel edge information. Also, this information will be affected to some
extent by speckles It is therefore necessary to obtain information from at
least three edge searches in order to determine the movement of the blood
vessel. By comparing the difference between the (m)th and (m+1)th data
with the difference between the (m+1)th and (m+2)th data, it becomes
possible to check whether or not there has been movement of the blood
vessel in the period from the acquisition of the (m)th data to the (m+2)th
data. If it is determined that there has been movement, it is possible to
determine the amount of movement by, for example, obtaining the weighted
averages of the differences. Correction is not required if there has been
no movement, therefore, a method shall now be explained which consists of
taking the smallest of the differences as the amount of movement.
FIG. 10 is a flow chart of a process for determining the amount of blood
vessel movement in accordance with this method. In step T1 the data up to
the preceding two searches is stored prior to the data being updated. Step
T2 is a blood vessel search, the details of which are as described with
reference to the flow chart of FIG. 9. In this step, fresh blood vessel
edge data is incorporated. In step T3 it is determined whether or not
sufficient data has been prepared to enable the amount of movement to be
obtained. If there is not enough data the process returns to step T1; if
the data is sufficient the process advances to step T4. In step T4
differences C1, C2, D1, D2 between consecutive data sets are obtained for
both edges, and in step T5 the presence or absence of movement is
determined by determining whether or not the differences C1, C2, D1, D2
have the same sign, which is to say, whether or not the movement has been
in the same direction in each case.
If the signs are the same and it is determined that movement has taken
place in the same direction in each case, the process advances to step T6.
If the signs are different and it is therefore determined that movement
has not taken place in the same direction, the process moves to step T7.
In step T6 the minimum value among C1, C2, D1, D2 is taken as the amount
of movement, and after computing the amount of correction, taking into
consideration the magnification and other such optical system factors, the
necessary correction amount for returning the blood vessel to the initial
position is obtained and output.
Step T7 is for when the movement of the blood vessel is so small that it is
not detected from just one or two searches. In such a case, in step T7 the
discrepancies C02, C01, C00, D02, D01, D00 between the initial positions
(a0, b0) and each edge (a.sub.m, b.sub.m), (a.sub.m-1, b.sub.m-1),
(a.sub.m-2, b.sub.m-2) are obtained. The signs of C02, C01, C00, D02, D01,
D00 are determined in step T8. The signs all being the same will signify
that there has already been a shift to one side from the initial position,
and the process advances to step T9, while if there are differences among
the signs it will be unclear whether or not movement has taken place to
one side from the initial position, so the process will return to step T1.
In step T9 the minimum of the discrepancy values C02, C01, C00, D02, D01,
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