U.S. patent number 3,854,049 [Application Number 05/423,115] was granted by the patent office on 1974-12-10 for compensation for patient thickness variations in differential x-ray transmission imaging.
This patent grant is currently assigned to Wisconsin Alumni Research Foundation. Invention is credited to Frederick Kelcz, Charles A. Mistretta.
United States Patent |
3,854,049 |
Mistretta , et al. |
December 10, 1974 |
COMPENSATION FOR PATIENT THICKNESS VARIATIONS IN DIFFERENTIAL X-RAY
TRANSMISSION IMAGING
Abstract
Differential x-ray images can be produced by sub-tracting two
different x-ray images which are produced by using two different
quasi-monoenergetic x-ray spectra. Two different x-ray filters may
be employed alternately to produce such spectra. For example, two
different filters containing iodine and cerium may be employed
alternately to produce two different x-ray spectra having peaks at
different energies. The x-rays to be filtered may be derived from
an ordinary x-ray tube which produces a continuous spectrum of
x-rays over a wide band of energies. The purpose of producing the
differential x-ray images is to subtract or cancel out the portions
of the x-ray images which are due to the ordinary tissues of the
patient, particularly the soft tissues, so that the presence of
certain contrast substances in the patient's body will be
emphasized or enhanced. Such contrast substances include iodine,
xenon or barium, introduced into the bloodstream, the lungs or the
food canal of the patient. The ordinary soft tissues of the patient
transmit x-rays of different energies to different extents. While
the two different x-ray images can be balanced for any particular
thickness of the patient, variations in such thickness over the
field of view tend to upset such balance. Thus, without
compensation for patient thickness variations, cancellation of the
x-ray images due to the ordinary tissues of the patient can be
achieved at only one value of patient thickness. In accordance with
the present invention, such patient thickness compensation is
achieved by adjusting the composition and density of the two x-ray
filters, and by adjusting the high voltage supplied to the x-ray
tube, so that the ratio of the two different x-ray images produced
by soft tissues remains nearly constant over a wide range of
patient thickness. Thus, substantial cancellation of the soft
tissue images can be achieved over a wide range of variations in
the patient thickness. The x-ray filtration is adjusted so that one
of the x-ray filters produces a spectrum of transmitted x-rays
having two peaks at energies below and above the energy of the peak
produced by the other x-ray filter. Thus, the sum of the images
produced by the two peaks tends to remain in a constant
relationship to the image produced by the other spectrum, despite
variations in patient thickness. The supply voltage to the x-ray
tube is varied as an inverse function of the average patient
thickness. Thus, the voltage is reduced when the average patient
thickness increases. The adjustment of the supply voltage to the
x-ray tube changes the two different quasi-monoenergetic x-ray
spectra so as to optimize the compensation for patient thickness
variations.
Inventors: |
Mistretta; Charles A. (Madison,
WI), Kelcz; Frederick (Madison, WI) |
Assignee: |
Wisconsin Alumni Research
Foundation (Madison, WI)
|
Family
ID: |
23677730 |
Appl.
No.: |
05/423,115 |
Filed: |
December 10, 1973 |
Current U.S.
Class: |
378/62; 378/108;
378/157; 378/98.11 |
Current CPC
Class: |
A61B
6/482 (20130101); H05G 1/60 (20130101); A61B
6/4035 (20130101); A61B 6/504 (20130101); A61B
6/00 (20130101); A61B 6/50 (20130101); A61B
6/481 (20130101) |
Current International
Class: |
A61B
6/03 (20060101); A61B 6/00 (20060101); H05G
1/60 (20060101); H05G 1/00 (20060101); G03b
041/16 () |
Field of
Search: |
;250/510,402,320,321,322,323 |
References Cited
[Referenced By]
U.S. Patent Documents
Primary Examiner: Lawrence; James W.
Assistant Examiner: Church; C. E.
Attorney, Agent or Firm: Burmeister, Palmatier &
Hanby
Claims
We claim:
1. A method of producing differential x-ray images to visualize a
contrast material in a patient,
comprising the steps of producing a first x-ray image using an
x-ray source and first filter,
producing a second x-ray image using said x-ray source and a second
filter,
said filters producing substantially different x-ray spectra,
producing a differential image corresponding to the difference
between said first and second x-ray images,
said contrast material being visualized in said differential
image,
and adjusting the voltage of said x-ray source as in inverse
function of the average patient thickness to minimize the effect of
variations in patient thickness upon said differential image.
2. A method according to claim 1,
in which the voltage of said x-ray source is adjusted to a range
which produces a flat peak of the ratio between said first and
second x-ray images when plotted against patient thickness.
3. A method of producing differential x-ray images,
comprising the steps of producing a first x-ray image using a first
x-ray spectrum,
producing a second x-ray image using a second x-ray spectrum,
and producing a differential image corresponding to the difference
between said first and second x-ray images,
said second spectrum having a spectral element of a predetermined
energy,
said first x-ray spectrum comprising a low energy spectral element
and a high energy spectral element,
said low energy spectral element having an energy below said
predetermined energy,
said high energy spectral element having an energy above said
predetermined energy.
4. A method according to claim 3,
in which said first and second x-ray spectra are balanced to
provide substantially the same image elements in said first and
second x-ray images for soft tissue whereby said image elements
will cancel out in said differential image.
5. A method according to claim 4,
in which said spectral elements are adjusted to produce substantial
balance between said first and second spectra over a substantial
range of tissue thickness.
6. A method according to claim 3,
in which said first and second spectra are produced by using first
and second x-ray filters in conjunction with an x-ray source.
7. A method according to claim 6,
in which said first and second filters contain iodine and cerium,
respectively.
8. A method according to claim 6,
in which said x-ray source includes a high voltage x-ray tube,
the supply voltage to said x-ray tube being varied as an inverse
function of the average tissue thinkness to maintain a substantial
balance between said first and second spectra for soft tissue over
a substantial range of tissue thickness.
9. A method according to claim 6,
including the additional step of changing the density of said
filters as a direct function of the average tissue thickness of the
patient.
10. A method according to claim 6,
including the steps of adjusting the density of said filters as a
direct function of the average tissue thickness of the patient
while adjusting the x-ray energy of said source as an inverse
function of said average tissue thickness to achieve effective
compensation for variations in said tissue thickness.
11. A method according to claim 6,
including the steps of adjusting the density of said filters while
adjusting the x-ray energy of said source to achieve effective
compensation for variations in the tissue thickness of the
patient.
12. A method according to claim 6,
including the steps of adjusting the density of said filters while
adjusting the x-ray energy of said x-ray source to produce a flat
spot in the characteristic curve of TB/TA plotted as a function of
patient tissue thickness at a desired average value of patient
tissue thickness,
where TB is the total of the x-ray photons transmitted through the
first filter and the patient while TA is the total of the x-ray
photons transmitted through the second filter and the patient,
whereby effective compensation is achieved for variations in the
patient tissue thickness.
13. A method according to claim 12,
in which said first and second filters contain iodine and cerium,
respectively.
14. A method according to claim 13,
in which said filters have iodine and cerium concentrations on the
order of 100 to 400 milligrams per square centimeter.
15. A method according to claim 12,
in which said flat spot is in the form of a minimum along said
characteristic curve.
16. A method according to claim 12,
in which said x-ray source includes an x-ray tube supplied with a
high voltage,
said voltage being adjusted as an inverse function of the patient
tissue thickness.
17. A method according to claim 16,
in which said first and second filters contain iodine and cerium,
respectively.
18. A method according to claim 17,
in which the iodine and cerium concentrations of the filters are on
the order of 100 to 400 milligrams per square centimeter.
19. A method according to claim 18,
in which said high voltage is on the order of 50 to 70 kilovolts
peak.
20. A method according to claim 17,
in which said high voltage is on the order of 50 to 70 kilovolts
peak.
21. Apparatus for producing differential x-ray images in which
image elements due to the soft tissue of a patient are at least
partially cancelled out,
comprising x-ray source means for selectively producing first and
second x-ray spectra,
means for producing first and second x-ray images of the patient
using said first and second x-ray spectra,
and means for producing a differential image corresponding to the
difference between said first and second x-ray images,
said second x-ray spectrum having an x-ray spectral element of
predetermined energy,
said first x-ray spectrum having low and high energy spectral
elements,
said low energy element having an energy less than said
predetermined energy,
said high energy element having an energy greater than said
predetermined energy.
22. Apparatus according to claim 21,
in which said spectral elements have magnitudes such as to produce
substantial balance between said first and second x-ray images for
soft patient tissue.
23. A method according to claim 22,
in which said spectral elements are of magnitudes to maintain the
balance between said first and second x-ray images for soft tissue
over a wide range of tissue thickness variations.
24. Apparatus according to claim 21,
in which said x-ray source means comprises an x-ray source and
first and second selectively usable filter means for producing said
first and second x-ray spectra.
25. Apparatus according to claim 24,
in which said first and second filter means are movable selectively
into operative relation with said x-ray source.
26. Apparatus according to claim 24,
in which said first and second filter means contain iodine and
cerium, respectively.
27. Apparatus according to claim 24,
in which said first and second filter means contain iodine and
cerium, respectively, in concentrations on the order of 100 to 400
milligrams per square centimeter.
28. Apparatus according to claim 27,
in which said x-ray source includes a high voltage x-ray tube and
means for supplying said tube with a high voltage on the order of
50 to 70 kilovolts peak.
29. An apparatus according to claim 24,
in which said first and second filter means contain iodine and
cerium, respectively,
said x-ray source including a high voltage x-ray tube and means for
supplying said tube with a high voltage on the order of 50 to 70
kilovolts peak.
30. Apparatus according to claim 24,
in which said x-ray source comprises a high voltage x-ray tube,
and means for varying the high voltage supplied to said tube.
Description
This invention relates to improvements in differential x-ray
imaging techniques. A differential x-ray image can be produced by
deriving the difference between two different x-ray images,
produced under different conditions as to x-ray energy, or as to
other factors. For example, the two different x-ray images can be
produced by using two different monoenergetic or
quasi-monoenergetic x-ray spectra. Two different monoenergetic
x-ray beams at different energies will produce different x-ray
images, because the two different beams will be transmitted and
absorbed differently by the tissues and other substances in the
patient's body. By subtracting the two images, a differential x-ray
image can be produced which will often emphasize or enhance the
visibility of particular tissues or substances in the patient's
body. Thus, for example, the visibility of x-ray contrast materials
can be enhanced. Examples of such materials are iodine, xenon and
barium, which can be introduced into the bloodstream, the lungs or
the food canal of the patient.
Instead of using two different monoenergetic x-ray beams, it is
generally more convenient to employ two different
quasi-monoenergetic x-ray spectra, which may have peaks at
different energy levels. The two different x-ray spectra will
produce different x-ray images which can be subtracted to produce a
differential x-ray image. By using this technique, the visibility
of contrast media or particular tissues can often be enhanced.
The two different quasi-monoenergetic x-ray spectra can be produced
by using two different x-ray filters alternately. The x-rays to be
filtered may be derived from an ordinary x-ray tube which produces
a continuous spectrum of x-rays over a wide band of energies. The
filters produce selective absorption of the x-rays so as to modify
the continuous spectrum to produce peaks at different energy
levels.
One important purpose of producing the differential x-ray images is
to subtract or cancel out the portions of the x-ray images which
are due to the ordinary tissues of the patient, particularly the
soft tissues, so that the visibility of certain distinctive tissues
or contrast substances in the patient's body will be enhanced. To
produce good cancellation, the portions of the two different images
which are due to the soft tissues must be balanced. However, even
the ordinary soft tissues of the patient transmit x-rays of
different energies to different extents. While the two different
x-ray images can be balanced for any particular thickness of the
patient, variations in such thickness over the field of view tend
to upset such balance. Thus, without compensation for patient
thickness variations, cancellation of the x-ray images due to the
ordinary tissues of the patient can be achieved at only one value
of patient thickness.
One object of the present invention is to provide effective
compensation for variations in the thickness of the patient over
the field of view, so that cancellation of the image portions due
to ordinary soft tissue can be accomplished to a high degree over a
range of patient thicknesses.
A further object is to provide effective compensation for patient
thickness variations without resorting to a complex electronic
compensation system.
In accordance with the present invention, compensation for patient
thickness variations is achieved by adjusting the nature of the
monoenergetic or quasi-monoenergetic x-ray spectra. It has been
found that the quasi-monoenergetic spectra can be adjusted by
adjusting the x-ray filtration and the supply voltage of the x-ray
source. The adjustment of the voltage changes the x-ray spectrum
produced by the x-ray tube. It is prefered to adjust the x-ray
filtration so that one of the x-ray filters transmits a spectrum
having two x-ray peaks at energies below and above the energy of
the peak produced by the other x-ray filter. Thus, the image
produced by the first x-ray filter consitutes the sum of the image
components produced by the two peaks. Such sum tends to remain in a
constant relationship to the image produced by the second spectrum,
despite variations in patient thickness. The peaks of the x-ray
spectra can be adjusted by varying the composition and density of
the x-ray filters. It is prefered to vary the supply voltage to the
x-ray tube as an inverse function of the average patient thickness.
Thus, the voltage is reduced when the average patient thickness
increases. The adjustment of the supply voltage changes the x-ray
spectrum produced by the x-ray tube, and thereby modifies the two
different quasi-monoenergetic x-ray spectra transmitted by the two
x-ray filters. By adjusting the supply voltage, it is possible to
optimize the compensation for patient thickness variations.
Many different materials may be employed in the x-ray filters. For
example, one filter may contain iodine, while the other contains
cerium. This combination of filters is particularly valuable for
enhancing the visibility of iodine in the patient's body. The
iodine filter produces a spectrum having first and second peaks at
low and high energy levels. The cerium filter transmits a spectrum
having a peak at an intermediate energy, higher than the low energy
peak but lower than the high energy peak produced by the iodine
filter. By adjusting the supply voltage, the relative magnitudes of
the high and low energy peaks can be adjusted.
Further objects, advantages and features of the present invention
will appear from the following description, taken with the
accompanying drawings, in which:
FIG. 1 is a diagrammatic illustration of an x-ray system to be
described as an illustrative embodiment of the present
invention.
FIG. 2 is a set of graphs illustrating the variation in the x-ray
attenuation coefficients, as a function of x-ray energy, for iodine
and for water, which accounts for most of the attenuation produced
by soft tissue.
FIG. 3 is a set of graphs illustrating the variation in the x-ray
attenuation coefficients for iodine and for cerium.
FIG. 4 is a set of graphs illustrating the wide band x-ray spectrum
produced by an x-ray tube, and also the quasi-monoenergetic x-ray
spectra which may be produced by using iodine and cerium filters in
conjunction with the x-ray tube.
FIG. 5 is a set of graphs illustrating the effects of variations in
the patient tissue thickness, with and without compensation.
FIG. 6 is a set of graphs illustrating the manner in which
compensation can be achieved for different ranges of patient
thickness.
FIG. 7 is a graph illustrating the manner in which the x-ray supply
voltage may be varied to achieve compensation for different values
of patient thickness.
FIGS. 8-12 are additional graphs illustrating the effects of
changing the filtration densities and the supply voltage.
As just indicated, FIG. 1 illustrates a system 10 for producing
differential x-ray images, with compensation for variations in the
thickness of the patient or subject 12 to be x-rayed. In general,
the thickness of the portion of the patient to be x-rayed may vary
in a more or less irregular manner over the field of view. While
the cross sectional shape of the patient 12 is shown as a simple
oval, it will be understood that the actual shape is more or less
irregular.
The x-ray system 10 comprises an x-ray source assembly 14 for
producing monoenergetic or quasi-monoenergetic x-ray spectra. In
this case, the source assembly 14 is adapted to produce
quasi-monoenergetic x-ray spectra. Thus, the x-ray source assembly
14 comprises an ordinary x-ray tube or source 16 which may be
energized by a variable high voltage supply 18. An x-ray source of
this type produces a continuous x-ray spectrum over a wide band of
energies, as represented by a graph 20 in FIG. 4. The maximum
energy of the band of x-rays is determined by the maximum voltage
applied to the x-ray tube 16, in this case sixty-five kilovolts
peak (KVp).
The x-ray beam from the x-ray source 16 is directed through the
patient 12 to an image detector 22. To produce quasi-monoenergetic
x-ray spectra, it is prefered to provide a plurality of selectively
usable x-ray filters, two such filters 24a and 24b being shown. The
x-ray filters 24a and b may contain various materials capable of
absorbing x-rays in a selective manner, so that the continuous band
of x-rays produced by the x-ray tube 16 will be converted into
x-ray spectra having one or more peaks at various energy levels. A
wide variety of filtering materials may be employed. For example,
the filters 24a and b may utilize iodine and cerium. FIG. 4
includes graphs 26a and b representing quasi-monoenergetic spectra
which may be produced by the use of the iodine and cerium filters.
It will be seen that the spectrum 26a produced by the iodine filter
24a comprises a low energy peak 26aL and a high energy peak 26aH.
The magnitude of the low energy peak 26aL is greater than the
magnitude of the high energy peak 26aH. Similarly, the spectrum 26b
produced by the cerium filter includes a low energy peak 26bL and a
high energy peak 26bH. In this case, the high energy peak 26bH is
so small as to be insignificant. The magnitude of the low energy
peak 26bL for the cerium filter is comparable to the magnitude of
the low energy peak 26aL for the iodine filter. It will be seen
that the low energy peak 26aL for the iodine filter occurs at a
lower energy than the low energy peak 26bL for the cerium filter.
On the other hand, the high energy peak 26aH for the iodine filter
occurs at a higher energy than that of the low energy peak 26bL for
the cerium filter.
The x-ray filters 24a and b are arranged to be used alternately, so
as to produce two different x-ray images. As shown in FIG. 1, the
filters 24a and b are movable alternately into the beam of x-rays
from the x-ray tube 16. Thus, the filters 24a and b may be mounted
on a movable indexing member, such as the illustrated rotatable
disc 28. A mechanical device, such as the illustrated motor 30, may
be employed to rotate the disc 28.
The two different x-ray images produced by the use of the filters
24a and b are detected by the image detector 22 and are subtracted
by an image subtraction device 32. The differential images thus
produced may be displayed in some suitable fashion, in this case by
a television monitor 34.
Any known or suitable system may be employed for subtracting the
x-ray images. For example, the image detector 10 may be arranged to
produce positive and negative optical images corresponding to the
two different x-ray images. The positive and negative images are
then combined to produce the desired subtraction.
The prefered system is to convert the two different x-ray images
into television images which are subtracted electronically to
produce a differential television image, to be displayed on the
television monitor 34. Thus, the image detector 22 preferably
comprises a television system for converting the x-ray images into
electronic television signals. The image subtraction device 32
comprises means for subtracting the successive television images.
Such means may utilize one or more image storage tubes, or other
storage devices. For example, a silicon target storage tube may be
employed, so that one of the two images can be written negatively,
while the other image is written positively. In this way, any
identical portions of the two images will be cancelled out. A
differential image will develop on the silicon screen of the
storage tube. This differential image can be displayed on the
television monitor 34.
While various image subtraction systems may be employed, it is
prefered to utilize the highly advantageous image subtraction
system disclosed in the co-pending patent application of Charles A.
Mistretta and Michael G. Ort, Ser. No. 369,824, filed June 14,
1973. Such system uses two stages of subtraction. In such system,
the x-ray images are converted into video signals, which are then
processed through two successive stages of video subtraction,
preferably utilizing two different types of video storage devices.
The second stage of video subtraction also involves integration of
the differential features, so that the contrast and visibility of
such features can be built up over a multiplicity of cycles. In the
prefered form of such subtraction system, an intensification screen
and a television camera are employed to convert the x-ray images
into first and second video image signals, which are successively
supplied to a video difference detector. Such detector preferably
utilizes a video storage tube capable of storing video images in
the form of electrical charges distributed over a dielectric layer
on a conductive target back plate.
The first and second video images are supplied sequentially to the
storage tube, which produces a first differential video signal
corresponding to the difference between the first and second video
signals. The first differential video signal is then supplied to a
second integrating and subtracting storage device, preferably
utilizing a second storage tube capable of storing video images in
the form of electrical charges on a mosaic of dielectric islands on
a conductive back plate. The first differential video signal is
written positively on the target of the second storage tube. The
first storage tube is then employed to develop a second
differential video signal corresponding to the difference between
the second and first video image signals. The second differential
video signal is written negatively on the target of the second
storage tube. In this way, the differential features of the first
and second differential video signals are integrated and enhanced,
while the identical or non-differential features of the first and
second differential video signals are combined subtractively so as
to cancel them from the target of the second storage tube.
To obtain the maximum enhancement of the differential features,
this cycle of subtraction and integration is repeated so that a
multiplicity of cycles are completed. The enhanced image on the
target of the second storage tube can be read, as desired, and
reproduced as a visible display on a television monitor. By this
system, differential features amounting to only a fraction of one
percent of the full contrast range of the successive x-ray images
can be enhanced to full contrast so that such differential features
will be clearly visible.
Reference may be had to such co-pending application for a more
detailed disclosure of such differential imaging system. Any other
known or suitable differential imaging system may be employed.
The iodine and cerium filters 24a and b are especially valuable for
visualizing small quantities of iodine in the patient's body.
Iodine is present naturally, particularly in the thyroid gland.
Moreover, iodine may be introduced into the patient's body through
the bloodstream or the food canal, to serve as a contrast medium or
substance.
FIG. 2 comprises a graph 36 representing the variation of the x-ray
attenuation coefficient for iodine, as a function of the x-ray
energy, expressed in kilo-electron volts (KEV). It will be seen
that the x-ray attenuation coefficient generally decreases with
increasing x-ray energy. However, the graph 36 has an abrupt
discontinuity, usually refered to as the k-edge 38, at which the
x-ray attenuation coefficient increases sharply. Thus, at an x-ray
energy E.sub.A above the k-edge energy, the attenuation coefficient
is abruptly and substantially greater than at an x-ray energy
E.sub.B, below the k-edge. The iodine and cerium filters make it
possible to utilize the k-edge discontinuity 38 to produce an
enhanced differential feature between the two x-ray images, so that
the visibility of such feature will be increased in the
differential x-ray image.
Thus, it will be seen from FIG. 4 that the low energy peak 26a L
for the iodine filter is at approximately 30,000 electron volts,
below the k-edge energy of about 33,000 electron volts. The low
energy peak 26bL for the cerium filter is at an energy of about
40,000 electron volts, above the k-edge energy. Thus, the x-ray
spectrum produced with the cerium filter will be attenuated much
more by iodine than will the spectrum produced with the iodine
filter.
FIG. 3 comprises graphs which compare the x-ray absorption
coefficients for iodine and cerium. It will be seen that FIG. 3
again includes the graph 36 for iodine, showing the k-edge
discontinuity 38. FIG. 3 also includes a graph 40 for cerium. Here
again, the x-ray absorption coefficient generally decreases with
the increasing x-ray energy. However, there is a sharp
discontinuity or k-edge 42 at a particular energy level, at which
the x-ray absorption coefficient increases sharply. The k-edge 42
for cerium is at a substantially higher energy than the k-edge 38
for iodine.
The spectrum 26a for the iodine filter in FIG. 4 includes a sharp
drop 44 which corresponds to the k-edge 38 for iodine. Similarly,
the spectrum 26b for the cerium filter in FIG. 4 includes a sharp
drop 46 which corresponds to the k-edge 42 for cerium.
FIG. 2 also includes a graph 47 which represents the variation of
the x-ray attenuation coefficient for water, as a function of x-ray
energy. It will be seen that attenuation coefficient decreases
gradually with increasing x-ray energy. The absorption coefficient
for water corresponds generally to the absorption coefficient for
soft tissue. It will be seen that the absorption coefficient for
the energy E.sub.A above the k-edge for iodine is slightly less
than the absorption coefficient for the energy E.sub.B, below the
k-edge 38. It is possible to compensate for this slight difference
for any particular value of patient thickness. This can be done,
for example, by adjusting the relative densities of the iodine and
cerium filters 24a and b so that the intensity of the x-rays
transmitted through the iodine filter 24a and the pertinent portion
of the patient's body is the same as the intensity of the x-rays
transmitted through the cerium filter at the pertinent portion of
the patient's body.
However, in accordance with the present invention, it has been
found that effective compensation can be provided for variations in
the thickness of the patient over a wide range, by adjusting the
filters and the supply voltage to the x-ray tube, so as to produce
particularly advantageous quasi-monoenergetic x-ray spectra.
As previously indicated, one of the filters, in this case the
iodine filter 24a, is adjusted to produce an x-ray spectrum having
two quasi-monoenergetic peaks, while the other filter, in this case
the cerium filter 26b, is adjusted to produce only one significant
quasi-monoenergetic peak. The two peaks produced by the first
filter are below and above the energy level of the peak produced by
the second filter. Thus, as to x-rays in the high energy peak for
the first filter, the absorption coefficient for soft body tissues
is less than for x-rays in the peak produced by the second
filter.
As previously indicated, spectra of this type are illustrated in
FIG. 4, in which the spectrum 26a for the iodine filter has a low
energy peak 26a L and high energy peak or bump 26a H. At an
intermediate energy, the spectrum 26b produced by the cerium filter
has a single significant peak 26bL. There may be a small high
energy peak or bump 26bH for the cerium filter, but this small bump
is so insignificant as to be negligible.
Quasi-monoenergetic spectra of this type produce compensation for
variations in the thickness of the patient, because the high energy
peak 26aH. of the first filter gradually predominates over the
intermediate energy peak 26bL of the second filter as the patient
thickness increases, while the low energy peak 26aL of the first
filter gradually predominates over the peak 26bL for the second
filter, with decreasing patient thickness. Due to this action, it
is possible to achieve a close balance between the x-ray
intensities transmitted through the two filters and the patient's
body for a considerable range of patient thicknesses.
It will be evident from FIG. 4 that as the patient thickness
increases, the low energy iodine peak 26aL is reduced with respect
to the main cerium peak 26bL, but the high energy peak or bump 26aH
of the iodine spectrum grows relative to the cerium peak 26bH. The
sum of the high and low energy iodine portions 26aL and 26aH of the
spectrum 26a thus remains approximately equal to the transmitted
intensity of the cerium spectrum. The ratio of the transmitted
intensities can be tuned to unity over a fairly broad range of
patient thicknesses, as illustrated in FIG. 5, which compares the
ratios of the transmitted intensities for the case of monoenergetic
x-ray lines situated at the positions of the main iodine and cerium
bumps 26aL and 26bL and for the case of three monoenergetic x-ray
lines approximating the spectra provided by the iodine and cerium
filters.
Thus, FIG. 5 comprises a first graph 50, shown in a broken line,
which plots the ratio of I.sub.ABOVE to I.sub.BELOW, as a function
of tissue thickness, expressed in centimeters. I.sub.ABOVE is the
intensity of the x-rays transmitted through the tissue at an x-ray
energy of 36,000 electron volts (keV). This energy is above the
k-edge for iodine. I.sub.BELOW is the transmitted x-ray intensity
through the body tissue for an x-ray energy of 30 keV, below the
k-edge for iodine.
It will be seen that the graph 50 rises steadily with increasing
tissue thickness, and passes through unity for only one value of
tissue thickness, approximately 15 centimeters.
FIG. 5 includes a second graph 52, shown in a full line, in which
I.sub.BELOW is produced by two x-ray spectral lines, at 30 keV and
at 50 keV. The latter line, at 50 keV, corresponds to the high
energy peak 26aH of the spectrum 26a produced by the iodine filter
24a. It will be evident that the graph 52 has a broad plateau 52a,
centered at an average value of patient thickness which can be
chosen by adjusting the filters and the supply voltage. In this
case, the plateau 52a is centered at an average patient thickness
of about 15 centimeters. This plateau provides a broad region of
relative insensitivity to patient thickness. Thus, the patient
thickness can vary over the field of view without materially
affecting the ability of the differential x-ray system to cancel
out the portions of the x-ray images due to ordinary soft
tissue.
The exact location of the plateau 52a of FIG. 5 can be changed by
varying the filters and the voltage supplied to the x-ray tube.
Suprisingly, it has been discovered that the center of the plateau
moves to greater patient thicknesses as the supply voltage to the
x-ray tube is lowered. Thus, the supply voltage needs to be
adjusted as an inverse function of the average patient thickness.
The supply voltage needs to be decreased as the average patient
thickness is increased, and vice versa. Thus, for example, it has
been found that for average patient thicknesses of 20 centimeters
or more, it is highly desirable to use a supply voltage of 50,000
volts peak (kVp), with high tube current to produce the desired
intensity of the x-rays, rather than raising the tube voltage.
Thus the supply voltage to the x-ray tube is an important factor in
achieving effective compensation for patient thickness variations.
The effect of changing the supply voltage is illustrated in FIG. 6,
which comprises a series of graphs representing the ratio of cerium
to iodine transmissions, as a function of patient thickness, for
different supply voltages. The density of the cerium filtration was
200 milligrams per square centimeter.
Specifically, FIG. 6 comprises three graphs 54a, 54b and 54c,
plotted for supply voltages of 50 kVp, 60 kVp and 70kVp. Each graph
shows a plateau, which is inverted, in this case, because the ratio
being plotted is the inverse of the ratio plotted in FIG. 5. The
plateau of the curve 54a, plotted for 50 kVp, is centered at a
patient thickness of about 22 centimeters, while the curves 54b and
54c, for supply voltages of 60 and 70 kVp, are centered at patient
thickness values of about 11 centimeters and 6 centimeters,
respectively. The transmission ratio at the center of each plateau
is not unity, but compensation for this factor can be made by
adjusting the electronic gain in the differential x-ray imaging
system. Thus, different values of electronic gain can be used in
the television system for the images produced with the use of the
iodine and cerium filters, to bring about the optimum cancellation
of the portions of the images representing ordinary soft tissue.
The plateaus of the curves shown in FIG. 6 then provide wide
regions in which the effectiveness of the cancellation is
insensitive to variations in patient thickness.
FIG. 7 comprises a graph 56 in which the desirable x-ray supply
voltage is plotted against patient thickness. This graph highlights
the desirability of reducing the supply voltage to the x-ray tube
as the average patient thickness is increased. By thus reducing the
supply voltage, it is possible to optimize the compensation for
variations in the patient thickness.
If the average patient thickness is decreased, the supply voltage
should be increased to optimize the compensation for patient
thickness variations. Thus, the supply voltage should be adjusted
as an inverse function of the average patient thickness.
It is possible to change the composition and density of the x-ray
filters so as to change the quasi-monoenergetic spectra produced by
the use of the filters. For a particular average patient thickness,
it is possible to adjust the filters so as to optimize the
compensation for variations in the patient thickness.
As indicated by the legends in FIG. 4, the curve 26a represents a
quasi-monoenergetic spectrum produced by the use of an iodine
filter having a density or concentration of 0.200 grams of iodine
per square centimeter. The quasi-monoenergetic spectrum represented
by the curve 26b was produced by a cerium filter having a density
or concentration of 0.238 grams of cerium per square
centimeter.
It is readily possible to vary the concentration of the iodine,
cerium or other material in the filter. For example, this can be
done by constructing each filter so as to include a tank or other
receptacle which can be filled with a liquid solution containing
the iodine, cerium or other material in a dissolved state. The
concentration of the solution can readily be varied. The tank or
receptacle may be made of a plastic material or some other material
which produces very little attenuation of x-rays.
The curves of FIG. 6 were produced with cerium and iodine filters.
The cerium filter has a density or concentration of 200 milligrams
of cerium per square centimeter. The iodine concentration of iodine
filter was comparable and was adjusted to produce a ratio of unity
between the iodine and cerium transmissions for zero patient
thickness. The patient thickness was then varied for several
different values of supply voltage to the x-ray tube, so as to
produce the three curves 54a, 54b and 54c.
In FIG. 6, a percentage figure is indicated by a legend for each
curve, representing the percentage of the differential x-ray signal
for an iodine concentration of 1 milligram per square centimeter in
the patient. The percentage signal applies to the broad minimum or
inverted peak in the curve. Each percentage figure represents the
maximum signal that can be obtained for that particular supply
voltage to the x-ray tube. For a supply voltage of 50 kVp, the
maximum differential signal was approximately 1 percent. For a
supply voltage of 60 kVp, the maximum differential signal was
approximately 0.95 percent. For a supply voltage of 70 kVp, the
maximum differential signal was approximately 0.8 percent. It will
thus be evident that the percentage value of the maximum
differential signal decreases with increasing supply voltage to the
x-ray tube. The graphs of FIG. 5 were obtained for monoenergetic
spectra which approximate the quasi-monoenergetic spectra produced
by iodine and cerium filters, as represented by FIg. 4. The curves
50 and 52 represent the results of calculations based on the use of
monoenergetic x-rays. Thus, the curve 50 represents the ratio of
I.sub.ABOVE to I.sub.BELOW as a function of tissue thickness, when
the x-ray energy E.sub.ABOVE is 36 keV, while the x-ray energy
E.sub.BELOW is 30 keV. The x-ray energy E.sub.ABOVE of 36 keV is
above the k-edge for iodine, while the x-ray energy E.sub.ABOVE of
30 keV is below the k-edge for iodine. The x-ray intensities
I.sub.ABOVE and I.sub.BELOW represent the transmitted x-ray
intensities for the x-ray energies E.sub.ABOVE and E.sub.BELOW.
The curve 52 of FIg. 5 represents the ratio of I.sub.ABOVE to
I.sub.BELOW for the same x-ray spectra, except that the x-ray
component E.sub.BELOW at 30 keV is supplemented by an x-ray
component at 50 keV having an intensity of approximately 14 percent
of the intensity of the 30 keV component.
The x-ray component at 36 keV corresponds generally to the cerium
peak 26bL of FIG. 4, which is above the k-edge of iodine, while the
x-ray component of 30 keV corresponds generally to the iodine peak
26aL, which is below the k-edge for iodine. The x-ray component at
50 keV corresponds generally to the high energy bump 26aH for
iodine, as represented in FIG. 4.
FIGS. 8, 9 and 10 represent the results which are obtained by using
different values of filter density or thickness, for different
values of the supply voltage to the x-ray tube. In each graph, the
ratio of TB to TA is plotted as a function of variations in the
tissue thickness of the patient. TB is the transmission of x-rays
through the patient with the use of the iodine filter, which has
its major spectral peak below the k-edge for iodine. Such peak is
designated 26aL in FIG. 4. TA is the transmission of the x-rays
through the patient with the use of the cerium filter, having its
major peak 26bL above the k-edge of iodine.
FIG. 8 comprises four curves 68a, 68b, 68c and 68d, representing
the results obtained by using four different sets of filters at an
x-ray supply voltage of 50 KVp. For the curve 68a, the cerium
filter had a density of 0.100 grams of cerium per square
centimeter, while the iodine filter had a density of 0.101 grams of
iodine per square centimeter. For an iodine concentration in the
patient of one milligram per square centimeter, the differential
signal has a percentage value of 0.76 percent at a patient
thickness of 15 centimeters.
For the curve 68b, the corresponding values are as follows: cerium
filter density 0.200; iodine filter density 0.195; percentage of
differential signal, 1.19 percent.
For the curve 68c, the corresponding values are as follows: cerium
filter density 0.300; iodine filter density 0.284; percentage of
differential signal, 1.42 percent.
For the curve 68d, the corresponding values are as follows: cerium
filter density 0.400; iodine filter density 0.370; percentage of
differential signal, 1.54 percent.
It will be understood that the symbol TA in FIGS. 8, 9 and 10
represents the total number of transmitted photons in the cerium
spectrum, having its major peak above the k-edge for iodine. The
symbol TB represents the total number of transmitted photons in the
iodine spectrum, having its major peak below the K-edge for iodine.
For each curve, the relative densities of the cerium and iodine
filters were adjusted to produce a ratio of unity between TB and
TA, for a patient tissue thickness of zero.
In each of the curves 68a-d of FIG. 8, the ratio of TB to TA
decreases with increasing tissue thickness toward a minimum which
occurs at approximately 20 centimeters or greater, depending upon
the density of filtration. These curves represent the situation for
an x-ray tube supply voltage of 50 KVp. For this relatively low
supply voltage, and for even lower voltages, any increase in the
density of filtration tends to increase the tissue thickness at
which the minimum occurs.
FIG. 9 comprises four curves 69a, 69b, 69c and 69d, which represent
the ratio of TB to TA as a function of patient tissue thickness,
for increasing values of filtration, at a supply voltage of 60 KVp.
For the curve 69a, the cerium filter had a density of 0.100 grams
per square centimeter, while the iodine filter had a density of
0.101 grams per square centimeter. The differential signal was 0.60
percent at a tissue thickness of 10 centimeters for an iodine
concentration in the patient of one milligram per square
centimeter.
The corresponding values for the curve 69b are as follows: cerium
filtration 0.200; iodine filtration 0.197; differential signal 0.95
percent.
For the curve 69c, the corresponding values are as follows: cerium
filtration 0.300; iodine filtration 0.288; differential signal 1.18
percent.
For the curve 69d, the corresponding values are as follows: cerium
filtration 0.400; iodine filtration 0.376; differential signal 1.34
percent.
It will be observed that for the supply voltage of 60 KVp,
represented by FIG. 9, the ratio of TB to TA decreases with
increasing tissue thickness until a minimum value is reached,
whereupon the ratio increases with further increases in the tissue
thickness. The increases in the filtration values do not increase
the tissue thickness at which the minimum occurs, to any great
extent. Thus, the position of the minimum remains at about 10.5
centimeters, despite the changes in the concentrations of cerium
and iodine in the filters. The curves of FIG. 9 indicate that a
supply voltage of 60 KVp is appropriate for an average tissue
thickness of about 10.5 centimeters, because this value of supply
voltage provides effective compensation for variations in patient
tissue thickness. Such compensation is optimized by a supply
voltage in this general neighborhood.
FIG. 10 comprises four curves 70a, 70b, 70c and 70d, representing
the ratio of TB to TA as a function of patient tissue thickness for
a supply voltage of 70 KVp, at four different filter concentrations
or densities. For the curve 70a, the cerium concentration was 0.100
grams per square centimeter. The iodine filter concentration was
0.102 grams per square centimeter. The differential signal was 0.46
percent for an iodine concentration of one milligram per square
centimeter in the patient.
The corresponding values for the curve 70 b are as follows: cerium
concentration 0.200; iodine concentration 0.201; differential
signal 0.70 percent.
The corresponding values for the curve 70c are as follows: cerium
concentration 0.200; iodine concentration 0.298; differential
signal 0.86 percent.
For the curve 70d, the corresponding values are as follows: cerium
concentration 0.400; iodine concentration 0.392; differential
signal 0.99 percent.
It will be observed that the curve 70a has a minimum at a tissue
thickness of about seven centimeters. The ratio of TB to TA
increases for smaller or larger values of tissue thickness.
With increasing filtration, the minimum value of the ratio tends to
occur at smaller values of tissue thickness. Thus, the minimum
value of the curve 70b is at a tissue thickness of about 6
centimeters. Thus, in general, at a supply voltage of 70 KVp or
higher, increasing the filtration tends to decrease the tissue
thickness at which the minimum occurs. Moreover, the value of the
ratio of TB to TA at the minimum point tends to increase with
increasing filtration. This is the opposite of the situation for
supply voltages of 50 or 60 KVp, at which the value of the ratio at
the minimum tends to decrease with increasing filtration.
In summary, the effect of changing the densities or concentrations
of the filters depends upon the supply voltage to the x-ray tube.
At 70 KVp or higher, increasing the filtration tends to decrease
the tissue thickness at which the minimum value of the ratio of TB
to TA occurs. At 60 KVp, changing the concentrations of the filters
does not affect the position of the minimum point to any
substantial extent. At 50 KVp or lower, increasing the filtration
tends to increase the tissue thickness at which the minimum ratio
occurs.
For an average tissue thickness of about 10.5 centimeters, 60 KVp
is an appropriate supply voltage, because the compensation for
variations in the patient tissue thickness will be optimized. For
average values of patient tissue thickness ranging down to 5
centimeters or less, the voltage should be increased progressively
to 70 KVp or higher, to optimize such compensation. As the average
tissue thickness increases to 20 centimeters or higher, the voltage
should preferably be reduced progressively to 50 KVp or lower, so
as to optimize the compensation for variations in the tissue
thickness.
FIG. 11 and 12 will be helpful in explaining the effects
represented by FIGS. 8-10. It will be seen that FIGS. 11 and 12
reproduce the graph 36 of FIG. 3, representing the x-ray absorption
coefficient for iodine, plotted as a function of x-ray energy. The
x-ray absorption coefficient is designated m in FIGS. 11 and 12. As
before, the graph 36 includes the abrupt k-edge 38.
FIGS. 11 and 12 also include the graph 26a of FIG. 4, representing
the quasi-monoenergetic spectrum produced by the use of the iodine
filter 24a. The graph 26a is superimposed upon the graph 36 in each
case.
In FIG. 11, the spectrum graph 26a is drawn for an x-ray supply
voltage of 50 KVp, but in FIG. 12, the graph 26a is drawn for a
supply voltage of 70 KVp. In FIG. 11, m.sub.1 represents the value
of the x-ray attenuation coefficient for iodine at the x-ray energy
value corresponding to the low energy peak 26aL of the iodine
filter spectrum 26a. On the other hand, m.sub.3 represents the
value of the x-ray attenuation coefficient at the energy level
corresponding to the high energy peak or bump 26aH of the iodine
filter spectrum. It will be seen that m.sub.3 is substantially
greater than m.sub.1. As a result, at a supply voltage of 50 KVp,
the high energy bump 26aH loses to the low energy bump 26aL as the
filters are made thicker or denser. As a result of this decrease in
the relative strength of the high energy bump or peak 26aH, a
greater tissue thickness is required before the transmitted x-rays
due to the high energy bump 26aH catch up with the x-rays
transmitted by the cerium filter. The energy level of the x-rays
transmitted by the cerium filter is substantially lower than the
energy level represented by the high energy bump 26aH, so that a
greater percentage of the x-rays represented by the high energy
bump are able to penetrate the increased tissue thickness.
The effect is just the reverse at 70 KVp, as will be evident from
FIG. 12. It will be seen that the x-ray attentuation coefficient
m.sub.1 at the energy of the low energy bump or peak 26aL is
greater than the x-ray attenuation coefficient m.sub.3 at the x-ray
energy corresponding to the high energy bump 26aH. Thus, as the
filters are made thicker or more dense, the x-rays due to the high
energy bump 26aH actually gain in magnitude relative to the x-rays
due to the low energy bump 26aL. Thus, less tissue thickness is
required for the more pentrating x-rays due to the high energy bump
26aH to equal or exceed the x-rays transmitted through the patient
due to the cerium spectrum 26b, shown in FIG. 4. Thus, the minimum
along each curve in FIG. 10 occurs at a relatively small tissue
thickness. As the filters are made more dense, the minimum occurs
at a decreased tissue thickness. This represents the situation at
70 KVp.
In FIG. 8, representing the situation at 50 KVp, the minimum along
each curve occurs at a relatively great tissue thickness. As the
filters are made more dense, the minimum occurs at a greater tissue
thickness.
FIG. 9 represents an intermediate situation at 60 KVp. In this
case, the minimum along each curve occurs at an intermediate tissue
thickness. Changes in the tissue thickness do not affect the
location of the minimum to any great extent.
From the graphs of FIGS. 8, 9 and 10, it will be possible for those
skilled in the art to select an appropriate supply voltage and
appropriate filter densities, according to the average tissue
thickness involved, in order to achieve effective compensation for
variations in the tissue thickness. The supply voltage and the
amount of filtration should be selected to produce a curve having a
minimum at or near the average tissue thickness.
It will be evident that iodine and cerium filters may be employed
very advantageously for producing differential x-ray images due to
iodine in the patient's body. Such iodine may be present naturally,
as in the thyroid, or may be introduced into the patient's body as
a contrast agent. The iodine and cerium filters are also valuable
for producing differential x-ray images due to xenon gas, which may
be inhaled into the lungs of the patient. Ordinary xenon gas can be
employed, because the xenon does not have to be radioactive. The
xenon gas produces a k-edge which is at a somewhat higher energy
level than the k-edge for iodine, but at a lower energy level than
the k-edge for cerium. Thus, the iodine and cerium filters can be
employed to produce a differential x-ray image due to xenon gas in
the lungs of the patient. By this technique, the lungs can be
visualized with a greater degree of clarity.
Various other contrast media may be employed to produce
differential x-ray images. Barium is an example of another contrast
substance. For each contrast substance, the composition and density
of the x-ray filters 24a and 24b are selected so as to produce
quasi-monoenergetic x-ray spectra having peaks above and below the
k-edge for the particular contrast substance. By proper selection
of the x-ray supply voltage and the filter densities, effective
compensation can be achieved for variations in the tissue thickness
of the patient.
* * * * *