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FIELD OF THE INVENTION
The invention relates to a method and an apparatus for forming an image of
the ocular fundus.
BACKGROUND OF THE INVENTION
Apparatus to study and photograph the ocular fundus have been on the market
for many years under the term fundus camera and have become indispensable
aids in ophthalmic diagnostics. They are used, for instance, to study the
cardiovascular system using the so-called fluorescein angiography, to
diagnose intraocular tumors, to detect vessels damaged by diabetes, to
detect detached retinae, etc. In any case, what matters is to obtain an
image in which even the finest vessel structures are still discernible.
In the known fundus cameras, the ocular fundus is illuminated by means of a
light beam projected on the pupil of the eye, and the light reflected from
the eye's retina is directed to an observation microscope for image
formation. These devices have meanwhile reached their limits in terms of
further improvement possibilities, in particular since the patient's
strain-bearing capacity cannot be increased further.
For diagnostic reasons, however, it is desirable to obtain images of the
ocular fundus having a still better resolution and improved contrast,
while at the same time reducing the strain on the patient as far as
possible.
An attempt to progress in this direction is described in U.S. Pat. No.
4,213,678. In the device described in this patent, a collimated laser beam
focused through the eye is used for the illumination of a single point on
the retina, the beam being deflected in a manner resulting in a sequential
point-by-point scanning of the retina in the form of a line-scanning
pattern. The light reflected from the retina and passing through the full
pupil of the eye is directed to a photoelectric receiver. Its signal
output is synchronized with the scanning motion of the laser beam and
serves to generate an image on a television monitor.
While in contrast to a fundus camera this known device reduces the strain
to which the patient is exposed, it cannot provide a substantial increase
in the resolution since the laser beam, due to diffraction, illuminates a
relatively large area of the retina (10 .mu.m in diameter, approximately).
This disadvantage cannot be eliminated by improving the external optical
imaging system, because the imaging media of the optical apparatus of the
eye are invariably affected by optical aberrations. Nor does a magnified
representation of a partial area of the retina, to which in this known
device the selected scanning field has to correspond, result in a
substantial improvement of the achievable resolution. Because of the
limited point resolution of the resonant galvanometer scanner used, this
known arrangement does not permit the simultaneous generation and
representation of both an overview image and an image part of high
resolution.
SUMMARY OF THE INVENTION
It is an object of the invention to provide a method and an apparatus for
forming an image of the ocular fundus which provides a high-resolution
overview image. It is a further object of the invention to provide such a
method and apparatus which will also permit the generation of image parts
which are freely selectable with regard to their location, afford the same
high resolution, have a selectable image contrast and are adapted to be
called up and displayed readily and simultaneously with the overview
image, without requiring an intervention in the scanning mechanism.
In the method of the invention, an image of the ocular fundus is formed on
a television monitor of television means having a predetermined number of
image points corresponding to the applicable television standard. The
method includes the steps of: scanning a laser light beam across the
fundus of the eye in the form of a scanning raster; detecting the light
reflected from the fundus with a photoelectric receiver; transmitting the
output of the receiver to the television means and forming a television
image of the fundus; and, during the scanning of the fundus image field,
obtaining signals representative of a number of image points larger than
that corresponding to the television standard in at least a part of the
fundus, said fundus part being selected to be sufficiently large so as to
permit every point of the image raster to be assigned a separate image
signal value when displayed on the television screen.
The apparatus for forming an image of the ocular fundus includes: laser
means for forming a laser light beam; scanning means for deflecting the
light beam in the form of a line scanning raster on the fundus of the eye;
receiver means for receiving the reflected scanning raster from the
fundus; television means connected to the receiver for receiving the
signal output thereof and forming an image of the fundus from the output
for viewing by an observer; synchronizing means for synchronizing the
raster of the television means to the scanning raster of the scanning
means; and, electro-optical means for detecting aberrations of the eye and
modulating the same on the laser light beam whereby the image of the
fundus formed by the television means is corrected for the aberrations.
With the method and apparatus of the invention, signals representative of a
greater number of image points than corresponds to the television standard
at least in a partial area are obtained during the scanning of the image
field. By designing this partial area sufficiently large, each point of
the raster is assigned a separate image signal value when displayed on the
television screen. This enables the partial area to be reproduced at the
full information density representable in the television image.
It is possible to operate in the online mode and to select the image part
by means of a so-called light pen, for example. For this purpose, it is
possible for instance to have initially an overview image produced in a
first scanning operation and to select the part to be emphasized on the
basis of this overview image. In a subsequent second scanning operation,
the scanning of this selected part yields signals representative of a
greater number of image points than corresponds to the television
standard.
It is also possible to obtain during the scanning of the entire image field
signals representative of a greater number of image points than
corresponds to the television standard. These signals are suitably
digitalized and stored in memory. In the actual evaluation operation, the
stored signals are represented as an overview image from which an image
part is selectable. Since the number of image point signals available for
this image part in the storage medium is equal to the number representable
in the television image, it is possible to display a magnified image part
of the same high resolution as the overview image without the need for
another scanning operation.
In both modes of operation, no intervention in the deflection mechanism
occurs. The image point signals are available in sufficient numbers for
the selected image part to contain all representable information items.
Accordingly, this image part is enlarged out of the overview image by
means of an electronic magnifying glass, the resolution in the image part
being improved in accordance with the selected coefficient of
magnification.
Irrespective of the mode of operation chosen, it is particularly
advantageous after selection of the image part to increase the intensity
of the scanning laser beam within the frame delimiting the image part.
This results in an improvement of the image contrast in the image part
without appreciably increasing the overall light exposure of the patient's
eye. Accordingly, the image part is viewed by means of what may be
referred to as an intensity magnifying glass.
It will be apparent from the foregoing that this new method enables a
selected part of the ocular fundus to be viewed with improved resolution
and higher image contrast, whereby the strain to which the patient is
exposed is even less than would be the case with a conventional fundus
camera.
The method of this invention permits particularly the display of the
overview image and of the selected image part on separate monitors
simultaneously, thus clearly indicating to the viewer the accurate
position of the image part in the field at all times.
The apparatus to implement the method of this invention is to be designed
in a manner permitting during the scanning of the image field the
obtainment of signals representative of a greater number of image points
than are representable in a television image. When applying the C.C.I.R.
television standard, for example, this means that significantly more than
800 image points have to be resolved in each scan line. In conventional
eye observation, the aim is to obtain a usable deflection angle of
30.degree. at which the field of view on the retina covers an area of
about 8 mm in diameter. It is obvious that with the apparatus known from
U.S. Pat. No. 4,213,678 the required high resolution is in principle not
attainable because the errors caused by the aberrations of the optical
imaging system of the eye do not allow the laser beam to be brought to a
spot diameter on the retina of less than 10 .mu.m, approximately.
The apparatus of this invention remedies this situation by providing a
polygonal scanner for fast and linear scanning of the image field in at
least one direction of the coordinates, and by disposing an active image
element between the laser light source and the scanners, the image element
cooperating with a sensor to form a closed loop circuit for optical
focusing of the image under adaptive control.
The active image element is advantageously designed as an active mirror as
known per se from literature describing other uses. The sensor in the
control loop is suitably a wave-front sensor. A control loop essentially
comprised of these elements permits the image to be focused under adaptive
control, i.e., it enables image deteriorations caused by the aberrations
of the optical imaging system and the transmitting media of the eye to be
compensated for.
In the apparatus of this invention, the illuminating laser beam is
generally widened to a diameter of between 3 mm and 4 mm, in exceptional
cases even still wider, and by compensating for all existing aberrations,
it is possible to focus the laser beam on a spot of a minimal diameter of
between 2 .mu.m and 3 .mu.m on the retina. This permits the resolution of
more than 5,000 image points per scan line, that is, it is possible for
example to resolve and represent individual receptors in the fovea.
Since the use of optical image focusing under adaptive control produces
data on the wave front of the imaging laser beam, the apparatus of this
invention enables the refractive index profile within the eye to be
reconstructed, permitting for the first time an automatic determination of
the refraction at high accuracy.
In the new apparatus, the deflection of the illuminating laser beam in at
least one direction of the coordinates is performed by means of a
polygonal scanner, the light beam being deflected by means of a rapidly
rotating polygonal mirror. The polygonal mirror is suitably used for
scanning in direction x, i.e., in the direction of the television lines.
For scanning in direction y, a polygonal scanner may likewise be used,
however it is also possible to achieve the deflection using a galvanometer
mirror.
In either case, it is necessary to obtain from the laser-beam deflecting
elements synchronizing pulses which serve to synchronize the raster of the
television monitor used for image generation with the scanning raster.
By the utilization of simple means, the apparatus for forming an image of
the ocular fundus constructed in accordance with this invention may have
its function extended by being employed as an apparatus for measuring the
blood flow and the oxygen saturation of the blood. Such measurements are
very valuable to the examining physician since the diagnosis arrived at on
the basis of the image of the ocular fundus can be corroborated by such
measurements without the need to apply another device.
The blood flow is measured by means of a laser Doppler velocimeter the
construction of which will become apparent from the subsequent description
of the figures. This device permits determination of the spatial blood
flow distribution in the ocular fundus which is of particular importance
for diabetes examinations.
The measurement of the oxygen saturation of the blood is a measure of the
spatial distribution of oxygen saturation in the retina which is of
importance for diabetes and cardiovascular system examinations.
Devices for measuring blood flow and oxygen saturation of the blood are
known per se. The present invention provides a novel and particularly
valuable possibility to perform these measurements with the same device
that serves for the formation of a high-resolution and high-contrast image
of the ocular fundus.
BRIEF DESCRIPTION OF THE DRAWING
The invention will now be described with reference to the drawing wherein:
FIG. 1 is an embodiment of the apparatus of the invention for forming an
image of the ocular fundus;
FIG. 2 is an exemplary image of the ocular fundus generated by the
apparatus according to FIG. 1;
FIG. 3 is a schematic of the active mirror utilized in the apparatus of
FIG. 1;
FIG. 4 is a perspective view of the active mirror;
FIG. 5 is a separate schematic showing the configuration of the wave-front
sensor utilized in the apparatus of FIG. 1;
FIG. 6 is an embodiment of a laser Doppler velocimeter for measuring the
flow of blood; and,
FIG. 7 is a graph showing the dependence of the extinction coefficient of
hemoglobin and oxyhemoglobin on wave length.
DESCRIPTION OF THE PREFERRED EMBODIMENTS OF THE INVENTION
Referring now to FIG. 1, reference numeral 1 identifies a laser utilized as
a light source which can, for example, be configured as a 2 mW helium-neon
laser which emits light at a wavelength of 633 nm. The light emitted by
the laser passes through an electrically controllable shutter 2 which can,
for example, be a Pockels cell and a polarization filter 3. The laser beam
is widened by optical system 4 illustrated schematically and is directed
to an active mirror 6 via a semi-transparent mirror 5. The light reflected
from this mirror 6 passes through an optical system 7 and impinges upon a
polygonal mirror 8 which is rotated by means of a motor 9 in the direction
of the arrow.
If the polygonal mirror 8 is for example made up of twenty facets, then its
rotational speed is approximately 12,000 revolutions per minute. During
its rotation, the mirror 8 deflects the laser beam linearly in line
direction x. At a line repetition rate of 4 KHz, a useable deflection
angle of 30.degree. is obtained with a 20% dead time caused by the change
of facets. System 10 shown schematically generates synchronizing pulses
which identify the beginning and end of each scan line. These pulses are
directed to a control unit 11.
Such a polygonal scanner is disclosed, for example, in the journal entitled
"Analytical and Quantitative Cytology", Vol. 3, No. 1, March 1981, pages
55 to 66, especially pages 57 and 63.
The optical deflecting plane is imaged upon a linear galvanometer scanner
12 via an objective 70. The galvanometer scanner 12 deflects the laser
beam in the vertical direction in a saw-tooth form. An arrangement 13 is
provided which drives the scanner 12 and which is supplied with control
pulses via the control unit 11. The control unit 11 couples all control
signals to the mirror reference signal of the system 10 in a phase-stable
manner. This is necessary because the high inertia of the rotating
polygonal mirror 8 does not permit the rotational speed of the latter to
be controlled to a definite cycle.
The illuminating laser beam is imaged on the pupil of the eye 16 via a
further objective 14, with the beam being transmitted into the eye via a
semi-transparent mirror 15. The illuminating beam has a reduced diameter
in the pupillary plane 17 of the eye 16 since the pupillary plane 17 is
conjugated to the deflection planes of the scanners 8 and 12.
The rays reflected on the retina 18 pass through a semi-transparent mirror
19 and are then collected by a non-spherical lens 20 having a high
aperture ratio. Thereafter, the rays pass through a further polarization
filter 21 and are detected by means of a receiver 22 in the region of a
plane conjugated to the pupillary plane 17 of the eye 16.
The signals delivered by the receiver 22 are amplified in amplifier 23 and
are fed to a television monitor 25 via a standard television signal
generator 24. The standard television signal generator 24 is controlled by
control unit 11 and delivers the synchronizing pulses necessary for
generating the image. In this way, an image of the ocular fundus 18
becomes visible on the monitor 25 since the scanning raster described by
the illuminating laser beam on the retina 18 corresponds to the television
raster.
If desired, the signals coming from the receiver 22 can also be digitalized
in the converter 26 after they are amplified in amplifier 23 and then
stored in a memory 27. This memory 27 can be configured as a refresh
memory of a digital image system which, after the end of the read-in
cycle, illustrates the image signals on the monitor 25 in such a manner
that they are phantom color coded. It is also possible to configure the
memory 27 as a long-term storage memory. In this instance, the image
signals are always recallable so that, for example, a comparison of images
generated at different points in time is possible. Also, a measured
evaluation of the image signals can thereby be made at any time.
The polarizers 3 and 21 are arranged in a crossed relationship to each
other. In this way, the corneal reflex is suppressed. In cooperation with
the electronic shutter 2, a definite part of the fundus selected from the
overview image on the monitor 25 can be emphasized by increasing its
intensity. FIG. 2 serves to provide a more detailed explanation, showing
an overview image which, for example, is generated by scanning the retina
18 of the eye 16 by means of the scanners 8 and 12. With the aid of a
light pen, a part 28 of the fundus, for example, can be selected from the
overall image on the monitor 25 in a known manner. The generator 24 then
generates the necessary synchronizing pulses which cause a higher
intensity of the laser image to be passed through the shutter 2 via the
switching arrangement 29 as long as the scanning beam moves within the
selected part 28. This intensity is therefore higher than the intensity
associated with a movement of the scanning beam outside of the part 28.
This part 28 thereby appears emphasized by means of a so-called intensity
magnifying glass and has a better image contrast than the surrounding
field. The part 28 is suitably reproduced in lieu of the overview image
and fills out the format of the monitor 25. Of course, it is also possible
to utilize two monitors so that both the overview image and the image of
the part 28 can be simultaneously reproduced and illustrated.
With this arrangement, it is possible to definitively locate damaged
vessels by means of the overview image or by means of the image of the
selected part. The electronic shutter 2 functions so rapidly that it is
also possible to photocoagulate the damaged vessels by means of a
momentary increase in the intensity of the laser light. In this method, it
is assured that the light intensity is only then increased when the
scanning beam has reached the coordinates of the vessels to be coagulated.
As shown in FIG. 1, the illuminating laser beam impinges upon a receiver 30
via the mirror 15 arranged in front of the eye 16. The receiver 30 serves
to control the applied power and automatically switches the laser beam off
via control unit 31 and the shutter 2 as soon as this power reaches
impermissibly high values. At the same time, this receiver 30 serves to
eliminate variations in intensity of the illuminating laser beam by means
of regulation via the shutter.
It has been shown that the images obtained by scanning and illustrated on
the monitor 25 have a more plastic appearance than do the images obtained
with a fundus camera of conventional construction. In order to be able to
select the optimal plasticity, it is desirable to utilize a
variable-frequency dye laser in lieu of the helium-neon laser 1. This
makes it possible to select the most suitable wavelength of the
illuminating projector.
An active image element 6 is arranged in the illuminating beam path of the
apparatus of FIG. 1 and serves to optically improve the image by adaptive
control within a closed loop circuit. This control circuitry includes a
schematically-illustrated wave-front sensor 32 which detects the optical
aberrations of the image of the retina 18 which is deflected upon the
sensor by the mirror 19. The aberrations are controlled via a control
circuit and the image element 6 to be described below.
In the illustrated embodiment, a membrane mirror is selected as the active
image element 6 and is known from the Journal of the Optical Society of
America, vol. 67 (1977), No. 3, March 1977, pages 399 to 406. As the
schematic illustration of FIG. 3 shows, such a mirror comprises a
transparent electrode 34 which is evaporated on a glass window and has a
voltage U.sub.o applied to it. A grounded mirror foil 35 is arranged at a
short distance (approximately 50 .mu.m) from this electrode. The mirror
foil 35 is a thin plastic foil approximately 0.5 .mu.m thick and
vapor-coated with aluminum. At short spacing of approximately 50 .mu.m
from the other side of the foil 35, an array of several electrodes 36 is
arranged which are individually controllable. The control of the
electrodes 36 is effected with a voltage U.sub.o .+-.U.sub.i which
develops a resulting electrostatic force that acts on the foil 35. In this
way, the mirror 6 corresponds to a multi-channel electrometer, with the
maximum deflection of the foil 35 lying in the area of magnitude of 1
.mu.m.
The control of mirror 6 is suitably effected on the basis of a modal
control process. For this purpose, basic forms of optical aberration such
as an astigmatism, spherical aberration, defocusing and coma are modulated
upon the illuminating laser beam in a plane conjugated to the pupillary
plane 17 of the eye 16. As shown in FIG. 4, this is effected by means of a
control unit 33. The control unit 33 includes an arrangement 33a which
generates the signals corresponding to the above-mentioned selected basic
forms of optical aberration. The control unit 33 includes a further
arrangement 33b for distributing the signals to the electrodes 36.
The image generated after switching in the mirror 6 is cast upon a
wave-front sensor 32 via the mirror 19. An example of the configuration of
this wave-front sensor 32 is shown in FIG. 5.
The light reflected from the eye 16 is imaged on a rotating matrix 36 by
means of an optical system 34. The optical system 34 includes a conjugated
aperture plane at the location indicated by reference numeral 35. The
interference pattern generated thereby is imaged on a diode array 38 via
the optical system 37. The diode array 38 measures the profile of the
wave-front. From this, signals are obtained via the processor 39 which
adjust the active mirror 6 via the arrangement 33 until the optimal
focusing parameters are determined which compensate for all optical
aberrations of the eye 16 to be examined.
When the optimal corrective condition is reached, the illuminating laser
beam, which is widened to a diameter of approximately 4 mm, can be focused
to a minimal focal magnitude of approximately 2.5 .mu.m in diameter on the
retina 18.
The data on the wave front of the imaging laser beam, which is obtained
with the above-described optical focusing of the image under adaptive
control, makes it possible to reconstruct the refraction profile within
the eye 16. The arrangement 40 is provided for this purpose. The
arrangement 40 is so configured that it shows the data obtained and/or
prints the same.
The corrective data determined for the mirror 6 is stored in the memory of
the control unit 33 and is always recallable therefrom in real time.
Before the eye 16 is observed by the physician, the aberrations are first
determined by the above-described optical focusing of the image under
adaptive control. For this purpose, a separate investigation is performed.
In this procedure, the optimal focusing parameters are determined at a
data rate of, for example, 100 lines per second in various directions. The
aberration data thereby determined for the eye 16 is stored in the data
memory of the computer 33. It is, for example, possible during an
image-forming time of 80 milliseconds to obtain a resolution of 800
subapertures using 63 corrective electrodes 36 within a selected fundus
section. This is made possible because of the speed of the active mirror 6
and the data processing.
During the actual investigative procedure, the corrective values for the
mirror 6 are recalled in real time from the memory of the computer 33 so
that geometric true scale fundus images can be generated which are not
falsified by the optical aberrations of the eye.
By scanning the retina 18 by means of an illuminating laser beam corrected
via the mirror 6, signals are obtained in a scan line which are
representative of a greater number of image points than can be shown on
the television screen of the monitor 25. In this way, more than 5000 image
points can be resolved in a line. The monitor 25 therefore shows an
overview image of signals which are averaged.
If, for example, part 28 of FIG. 2 is selected from the overview image for
a more precise examination, this part fills out the format of the monitor
25 so that it takes up the whole screen whereby the full resolution is
reached if the part 28 is selected to reflect an appropriate magnitude.
The standard signal generator 24 provides for such an adaptation.
With the apparatus of FIG. 1, it is thus possible to observe an image
section with improved image resolution without requiring any kind of
adjustment to the optical system of the apparatus. Accordingly, one can
characterize the foregoing as obtaining an image enlargement by means of
an electronic magnifying glass.
It is possible and advantageous to also make use of the intensity magnifier
at the same time; that is, the selected image part can be scanned with
greater intensity. In this way, a detailed illustration of the fundus
image with a high resolution and high image contrast is made possible for
the first time. The resolution is so large that individual receptors in
the fovea are resolved.
It is generally not necessary to correct the illuminating laser beam
optimally over the entire image field by means of the active mirror 6. Of
special interest are mostly small image fields, for example, the fovea or
also individual vessels.
For this reason, an overview image is generally generated first, and the
mirror 6 remains inactive. On the basis of this overview image, the
interesting parts are then selected on the monitor 25 and the optical
focusing of the image under adaptive control is carried out with respect
to these parts. The corrective values determined thereby are stored in the
computer 33.
In the subsequent enlarged illustration of the selected parts, the parts
are illustrated on the monitor 25 with a high image sharpness and high
contrast.
When investigating the abnormal eye, for example the eye which is extremely
myopic or also after a cataract operation, it may also be suitable to
correct the illuminating beam over the entire image field in order to
obtain a good overview image on the one hand and, on the other hand, to
permit the selection of any desired part.
In order to obtain a high resolution of the image with the apparatus of
FIG. 1 also in the vertical direction, it is possible to employ an image
window which is in a form of stripes and contains all image lines to be
realized by means of the scanner 12. This image window is then displaced
in the vertical direction and the entire image is put together from the
individual stripes in the memory 27.
The apparatus according to the invention is especially advantageously
configured so that in addition to providing high-precision image
illustration, it can also measure the spatial blood flow distribution in
the fundus as well as the spatial distribution of the oxygen saturation of
the blood in the retina.
A differential laser Doppler velocimeter 41 is utilized for measuring the
spatial blood flow distribution and is illustrated separately in FIG. 6.
The beam of the helium neon laser 1 is divided into two parallel beams of
the same intensity by the prism system 42. The beams pass through the dove
prism 43 and the objective 44 and are imaged on a common point of the
retina 18. An interference pattern occurs in the region where the volume
is measured. Erythrocytes which transverse this light matrix generate a
scatter signal having a modulation frequency which is proportional to the
stripe distance and to the particle velocity. This scatter signal is
directed via mirrors 45, 46 to a receiver 47 in front of which an
interference filter 48 and an aperture 49 are arranged. The measuring
signal delivered by the receiver 47 is filtered by a band-pass filter 50,
is digitalized in a converter 51 and temporarily stored in a
microprocessor system 52. The capacity spectrum is there computed with the
aid of a time-optimized Fast Fourier transformation from which the speed
of flow is determined and indicated at 53.
The velocimeter of FIG. 6 is not shown in FIG. 1 in order to avoid
complicating the schematic. Its location in FIG. 1 is indicated by arrow
41, that is, it practically replaces the optical system 4. The lens 44 in
FIG. 6 is defined by the lens of the eye 16.
The apparatus of FIG. 1 can also be operated with a helium-selenium laser
54 having a beam which is mirrored in via mirror 55. Such a laser can, for
example, have an output power of 100 milliwatts and extend over a great
many lines in the visible region of the spectrum. The yellow laser lines
in the region of the absorption of the hemoglobin and oxyhemoglobin
molecule are suitable for measuring the degree of oxygen saturation in the
blood vessels of the retina. Because of the different spectral absorption
characteristics of the oxyhemoglobin and of the deoxygenated hemoglobin,
the degree of oxygen saturation can be determined by measuring at the
wavelengths of 559 nm, 569 nm, and 586 nm. For this purpose, the light
reflected from the eye 16 is directed via a mirror 56 and two spectral
mirrors 57 and 58 to three receivers 59, 60 and 61. The spectral signals
generated by these units are transmitted to a computer 62 computing the
oxygen saturation.
FIG. 7 shows the wavelength dependence of the extinction coefficients of
hemoglobin and oxyhemoglobin. The measurement is made at a wavelength of
559 nm, whereas at the wavelengths of 569 nm and 586 nm isobesto points
exist which are used to form a standard. In order to compensate for the
normally present scatter, it can be advantageous to utilize two or three
further wavelengths generated by the laser 54. In this way, the
signal-to-noise ratio can be improved.
It is understood that the foregoing description is that of the preferred
embodiments of the invention and that various changes and modifications
may be made thereto without departing from the spirit and scope of the
invention as defined in the appended claims.
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