Note: Descriptions are shown in the official language in which they were submitted.
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X-Ray Arrangement for Graphic Display of an Object Under Examination and
Use of the X-Ray Arrangement
Description:
This invention relates to an x-ray arrangement for graphic display of an
object
under examination that contains at least one radiopaque chemical element by
means of
x-ray radiation, a use of the x-ray arrangement as well as an imaging x-ray
contrast
process on an object under examination, for example a mammal, especially a
human.
The medical diagnosis with the aid of x-ray radiation is a technically highly-
developed field for diagnosis of diseases, for example for early detection,
for
radiographic identification, for characterization and for location of tumors,
vascular
diseases and other pathological changes of the human body. The technique is
very
efficient and exhibits high availability.
To produce x-ray radiation, x-ray tubes, for example with W-, Mo- or Rh-
rotating anodes and Al-, Cu-, Ti-, Mo- and Rh filters, are available. With
suitable
filtration, a portion of the bremsstrahlung is filtered out, such that in
advantageous
cases, essentially the characteristic radiation emerges from the x-ray tubes.
As detectors, either conventional x-ray films, digital plates or digital flat-
bed
detectors are used. In computer tomographs, a detector line or several
detector lines are
used. Also, several detectors can be connected in parallel. For direct
conversion of x-
ray radiation into electric signals, semiconductor detectors that consist of
cadmium-
telluride (CT), cadmium-zinc-telluride (CZT), amorphous salts or amorphous or
crystalline silicon are used (M. J. Yaffe, J. A. Rowlands, "X-Ray Detectors
for Digital
Radiography," Med. Biol., 42(l) (1997) 1-39).
An example of the design of such detectors is indicated in US 5,434,417 A. To
also make possible an energy sensitivity of the detector, the latter is formed
from
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several layers. X-Ray radiation with different energy is drawn through in
different
depths in this detector and produces an electric signal in the respective
layer by
photoelectric effect, which can be read out according to the layer and thus
according to
the energy of the x-ray photons, immediately identifiable as a current
impulse.
Computer tomography (CT) has already been used for a long time as a routine
process in regular clinical practice. With CT, sectional images through the
body are
obtained, with which a better spatial resolution is achieved than with the
conventional
projection radiography. Although the density resolution of the CT is also
clearly higher
than the density resolution of the conventional x-ray technology, contrast
media are still
required for reliable detection of many pathological changes. The latter
improve the
quality of the morphological information. In this case, on the one hand,
functional
processes are visualized in the body through the contrast medium (excretion,
perfusion,
permeability), and on the other hand, the morphology is emphasized by the
provision of
contrasts (different contrast medium concentrations in various tissues).
In many cases, the conventional x-ray technology could not be used, since the
contrast of the tissue to be examined was not adequate. For this purpose, x-
ray contrast
media were developed that produce a high radiographic density in the tissue in
which
they accumulate. Typically, iodine, bromine, and elements of atomic numbers
34, 42,
44 - 52, 54 - 60, 62 - 79, 82 and 83 are proposed as opacifying elements as
well as the
chelate compounds of the elements with atomic numbers 56 - 60, 62 - 79, 82 and
83.
As iodine compounds, for example, meglumine-Na- or lysine-diatrizoate,
iothalamate,
ioxithalamate, iopromide, iohexol, iomeprol, iopamidol, ioversol, iobitridol,
iopentol,
iotrolan, iodixanol and ioxilan (INN) can be used (EP 0 885 616 A1).
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In some cases, despite the administration of x-ray contrast media, no adequate
tissue contrast could be achieved. To achieve an additional increase in the
contrast,
digital subtraction angiography (DSA) was introduced, in which pre- and post-
contrast
images (logarithmic) are subtracted from one another. A subtraction method for
use in
mammography is disclosed in EP 0 885 616 Al: For projection mammography, it is
proposed there first to record a pre-contrast mammogram, then the patient is
to be
quickly injected i.v. with a commonly used urographic x-ray contrast medium,
and a
post-contrast mammogram is to be recorded about 30 seconds to 1 minute after
the end
of the injection. The data that are obtained from the two images are then
correlated with
one another, preferably subtracted from one another.
New developments in the field of CT relate to the excitation side, for
example,
the use of synchrotron radiation in CT (F. A. Dilmanian, "Computed Tomography
with
Monochromatic X-Rays," Am. J. Physiol. Imaging, 314 (1992) 175-193). Good x-
ray
images are obtained, for example, by means of "K-Edge Subtraction CT" (F. A.
Dilmanian, op. cit., page 179), whereby the strong increase of the absorption
coefficient
in the binding energy of the K-electron of an atom is used. The element iodine
has a K-
edge at an energy of 33.17 keV. Unfortunately, this process works only with
the aid of
the synchrotron radiation that is available to large storage rings, such as,
for example,
with DESY, since only this radiation has the monochromasia and intensity that
are
advantageous for the process. Conventional x-ray tubes do not yield any
monochromatic radiation but rather a continuous spectrum. They are therefore
not
readily suitable for such measurements of difference.
An alternative possibility is described in DE 101 18 792 A l: Here, to record
projection mammograms, a process is proposed in which x-ray radiation sources
with
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two x-ray anodes made of different materials are used. To record the
mammogram, first
an x-ray contrast medium is administered to the patient. Then, a first
projection image
is recorded with use of the first of the two x-ray anodes and then a second
projection
image is recorded with use of a second x-ray anode. By the superposition of
each
individual pixel from the first image with each individual corresponding pixel
from the
second image, a correlation image is then created. The characteristic
radiation of the
two x-ray anodes is matched to the absorption spectrum of the x-ray contrast
medium:
the emission energy of the first x-ray anode lies slightly below the
absorption energy of
the opacifying element in the x-ray contrast medium, and the emission energy
of the
second x-ray anode lies slightly above the absorption energy of the opacifying
element.
A drawback of this process consists in the fact that conventional x-ray tubes
must be
replaced with only one x-ray anode from bi-anode tubes.
In addition, for transmission radiography, an emission radiography is also
described:
Thus, in WO 2004/041060 A2, a device for non-invasive in-vivo determination
of a chemical element in the prostate of a human that has a probe, an
irradiation system
with which the chemical element can be excited to produce the emission of
radiation, a
radiation detector within the probe with which the emitted radiation can be
imaged, as
well as a signal recording, processing and display system, with which the
amount of the
chemical element in the prostate can be reproduced at various spots
corresponding to
the imaging of the emitted radiation, is described. The emitted radiation
essentially
consists of fluorescence radiation. In the case of the study of the prostate,
preferably the
distribution of Zn in the tissue is determined.
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In addition, in DE 36 08 965 Al, a process for determining the proportion of
various chemical elements in one layer of an area of examination by means of
gamma
or x-ray radiation is described. In this case, the Compton and the Rayleigh
scattered
radiation are detected separately. The course of the differential scatter
coefficients
determined from the measured values is influenced by the proportions of
various
chemical elements contained in the individual pixels. The proportion of these
chemical
elements can therefore be determined from them. To this end, out of a number
of
directions, a primary beam is drawn through the area of examination, and the
radiation
that exits from the area of examination under various angles is detected by a
detector
arrangement in various positions outside of the area of examination, after
which the
differential scatter coefficient for various pulse transfers is determined
from the
measured values for each pixel of the layer that are obtained in this case.
Also, Quanwen, Yu et al., "Preliminary Experiment of Fluorescent X-Ray
Computed Tomography to Detect Dual Agents for Biological Study" in: J.
Synchrotron
Rad. (2001), 8, 1030-1034 proposed the use of the x-ray fluorescence method
for
determining very low concentrations of non-radioactive substances in
biomedical
studies. By means of these methods, images can be obtained with which multiple
agents can be detected simultaneously with use of the fluorescence-Ka line in
a single
study to detect quantitatively, for example, the flow of blood in the brain
and the density
of the brain cells. In the presented study, images generated by means of these
methods
were compared to images obtained by means of x-ray transmission tomography.
The x-ray fluorescence or x-ray scattered light method described in the
publications indicated above has the drawback, however, that a visualization
in small
details of an object under examination is not easily possible because of
difficulties in the
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imaging. Rather, only highly resolved visualizations are obtained, such that
smaller
details are very difficult to visualize graphically.
The problem of this invention is therefore to avoid the above-mentioned
drawbacks and to find in particular arrangements and processes with which
images can
be produced with different radiopaque chemical elements. Further, the x-ray
images are
also to be able to be recorded in a simple, easy way, without high costs
resulting. The
technology is to be available on a broad basis. Also, smaller lesions in the
object under
examination are to be able to be made visible with the smallest possible
radiation dose
with high site resolution. Artifacts of movement are to be avoided.
This problem is solved by the x-ray arrangement for imaging an object under
examination that contains at least one radiopaque chemical element by means of
x-ray
radiation according to claim 1, the use of this x-ray arrangement according to
claim 11,
and the imaging x-ray contrast process according to claim 25. Preferred
embodiments
of the invention are indicated in the subclaims.
If the terms "emission" and "emitting" are used in the description of the
invention and in the claims below, they are to be understood to include, on
the one
hand, x-ray fluorescence, i.e., the emission of radiation after an excitation
of the
irradiated matter by means of electromagnetic radiation, and, on the other
hand,
preferably Rayleigh radiation. In the latter case, the radiation is emitted
again without
pulse transfer from the irradiation matter, whereby, however, no excitation of
orbital
electrons in atoms of this matter into excited states takes place as in the
case of
fluorescence because of the irradiation.
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With the x-ray arrangement, x-ray radiation that is transmitted for graphic
display through the object under examination and that is emitted from the
latter is used.
To this end, the x-ray arrangement according to the invention has the
following features:
a. At least one x-ray radiation source that emits essentially polychromatic x-
ray
radiation,
b. A first detector or a first detector unit (a unit that consists of several
detectors that are connected and/or arranged in parallel), with which values
of a first intensity of the x-ray radiation that is transmitted through the
object
under examination can be determined,
c. A second detector or a second detector unit, with which values a second
intensity of the x-ray radiation that is emitted from the object under
examination can be determined,
d. At least one correlation unit, with which the first intensity values of the
transmitted x-ray radiation can be correlated with one another with the
second intensity values of the emitted x-ray radiation, pixel for pixel, and
e. At least one output unit for visualizing the object under examination from
the pixel signals that can be obtained by correlation of the first intensity
values with the second intensity values.
The transmitted x-ray radiation and the emitted x-ray radiation can be
detected
either simultaneously or in succession.
This x-ray arrangement can advantageously be used for the graphic display of
an
object under examination that preferably contains at least one radiopaque
chemical
element by means of x-ray radiation. The radiopaque chemical element is
preferably
introduced by an x-ray contrast medium into the object under examination, and,
for
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example, administered to this end to the object under examination, for example
a human
or an animal.
The opacifying chemical elements with low atomic numbers that occur naturally
in an object under examination have only a small yield of x-ray fluorescence,
such that
the imaging with use of these elements does not appear to be practicable. In
addition,
the energy of the x-ray fluorescence photons is low in this case, such that
even their
range of action in the body tissue is low. In particular, starting from the
element iodine
(Z = 53) with the emission lines 28.6 and 32.3 keV, fluorescence lines are
available that
leave the object under examination to a sufficient extent and thus can be
recorded by a
detector arranged outside of the object. In the case of a lower atomic number
of the
chemical element, an arrangement of the second detector can be selected in
which the
latter is arranged as close as possible to the area to be examined (ROI:
Region of
Interest).
The x-ray arrangement is used to perform the x-ray contrast process according
to
the invention. The process has the following process steps:
a. Preferably administration of at least one radiopaque chemical element,
b. Irradiation of the object under examination with essentially polychromatic
x-
ray radiation,
c. Determination of values of a first intensity of the x-ray radiation that is
transmitted through the object under examination,
d. Determination of values of a second intensity of the x-ray radiation that
is
emitted by the object under examination,
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e. Correlation of the first intensity values of the transmitted x-ray
radiation,
pixel for pixel, with the second intensity values of the emitted x-ray
radiation,
f. Visualization of the object under examination from pixel signals that are
obtained by correlation of the first intensity values with the second
intensity
values.
In contrast to the known processes, in which either only the x-ray
transmission
tomography (TXCT) is performed or only the x-ray fluorescence (FXCT) is
detected,
transmission and the emission are measured here simultaneously or in
succession, and
these two techniques are combined with one another according to this
invention,
whereby the images that are obtained respectively in this case are
superimposed by a
suitable correlation process. In this procedure, the respective advantages of
both
techniques are used:
The x-ray transmission tomography namely offers the advantage of a high
achievable temporal and spatial resolution, such that in principle, the
smallest lesions or
other details can also be resolved in a human body that has been examined.
However,
the contrast that is obtained frequently is not sufficient also to make these
details
visible. This applies in particular for tests of lesions in soft tissue. In
addition,
examinations of certain body regions with the TXCT process are also hampered
by the
skeleton.
In contrast, the x-ray fluorescence tomography offers the advantage of an
extraordinarily high-contrast visualization, since only certain chemical
elements emit
electromagnetic radiation with suitable excitation of these elements, so that
these
elements that are found in the area of examination (ROI) are suitable as
extremely
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sensitive measuring probes. The FXCT method, however, suffers from the
drawback of
a low spatial resolution, such that smaller lesions can no longer be
visualized.
Only by the correlation of the intensity values of the transmitted x-ray
radiation,
pixel for pixel, with the intensity values of the emitted x-ray radiation and
the
visualization of the object under examination from pixel signals that had been
obtained
by this correlation can a high-contrast and detailed image of the area under
examination
(ROI) be produced. The opacifying image portion has namely a low resolution.
By the
correlation of the respective values with one another, this deficiency,
however, is
remedied to a great extent, since the necessary detailed information
originates from the
intensity values of the radiation that is measured by means of TXCT.
The invention can be used in particular for examining humans. The invention is
suitable for producing radiographs for visualizing masses, vessels and
perfusion, for
example for visualizing the esophageal-gastrointestinal passage, for
bronchiography,
chloegraphy, angiography and cardiac angiography, for cerebral angiography and
for
perfusion measurements, for mammography as well as for lymphography. The focus
in
the application engineering of the invention lies in computer tomography (MS-
CT;
CT), and their fusion modalities (PET-CT (positron-emission tomography), SPECT
(Single-Photon-Emission-Computer Tomography), sonography and with other
methods
of optical imaging). In principle, the invention can also be used to study non-
living
materials, for example in the field of materials testing.
To perform an examination, the transmitted radiation is recorded by means of
the first detector, which is found in the beam path of the x-ray tubes
attenuated by the
object under examination. The emitted radiation is measured by means of the
second
detector, which is arranged outside of this beam path, preferably at an angle
of about
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90 to the beam path. This second detector can in principle also be arranged,
however,
in any other angular position to the x-ray beam, for example 45 or 135 to
the beam
starting from the x-ray radiation source, without, however, it being detected
by the beam
being drawn through the object under examination. If the x-ray tubes are found
in the
12 o'clock position, ordinary computer tomographs are equipped with a series
of
detectors at the opposite 6 o'clock position. The second detector can
preferably be
arranged in the 3 o'clock position and/or the 9 o'clock position. By means of
this
second detector, both x-ray fluorescence and x-ray scatter (Rayleigh scatter,
Compton
scatter) can be recorded.
For selective detection of images with the second detector with use of the
emitted x-ray radiation, the energy of the emitted x-ray radiation can be
measured in
resolved form. In the presence of a pre-established emitting chemical element
in the
object under examination, it is especially advantageous to discriminate the x-
ray
radiation that is recorded by the second detector and that originates from
this opacifying
element from another emitted x-ray radiation, for example from scattered
radiation
(Compton radiation, Rayleigh radiation) and fluorescence radiation originating
from
other chemical elements. It is thus made possible to make certain areas (ROI)
very
selectively visible by, for example, building up the concentrations of
opacifying
chemical elements in certain organs of a human body, such that an especially
high
contrast of the tissue that is made visible compared to the surrounding tissue
is
produced. Also, the structure in a graphic display that is produced from the
skeleton is
less prominent in such a case compared to the visualization of the tissue, so
that the
skeleton leaves the graphic display virtually undisturbed.
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To detect and characterize the emission radiation, preferably an energy-
dispersive detector is used. It is also possible, however, to use simpler
detectors for this
purpose, and to ensure the characterization of the emission by x-ray-optical
modules
(filter combination, monochromators).
In addition, this principle can be applied in the same way to the measurement
of
the intensity values of the transmitted x-ray radiation with the first
detector. Also, in
this case, a selective visualization of the areas in the object under
examination (ROI) is
achieved, in which the opacifying chemical elements are concentrated.
Therefore, soft tissues, for example in the human, can also be visualized in
high
contrast with the invention. By coordinating the energy or the energy interval
of the
transmitted and emitted x-ray radiation that is recorded by the detectors with
the type of
opacifying chemical element, an efficient increase in contrast compared to
conventional
processes can be achieved.
To generate the x-ray radiation, a normal, commercially available x-ray tube
with a continuous spectrum can be used, for example a tube with an Mo, W or Rh
anode. Depending on the type of opacifying chemical element that is contained
in the
object under examination, a voltage is applied that makes possible an emission
of the
continuous radiation in the range of up to, for example, over 100 keV.
In principle, the x-ray radiation source can be operated without filtering the
emitted radiation, such that polychromatic radiation occurs in the entire
spectral range
on the object under examination. To reduce the radiation exposure of the
object under
examination, however, it is also possible to filter out such x-ray radiation
from the
spectrum of the polychromatic x-ray radiation source, whose energy is not
necessary or
is not advantageous for the detection. To this end, for example, an Al or a Cu
filter is
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used, which filters out energy in the range of < 20 keV (soft radiation).
Defined as a
continuous spectrum is thus an x-ray emission in a range of > 0 keV,
preferably > 15
keV, especially preferably > 17 keV, and quite especially preferably > 20 keV,
up to, for
example, 100 keV, whereby no spectral range within these limits compared to
others is
emphasized or excluded. The upper limit of the emission spectrum is determined
by the
voltage that is applied to the x-ray anode. The low-energy range of the
radiation is
preferably filtered out to eliminate the dose-relevant radiation for the human
body.
Normally, the object under examination is examined with polychromatic x-ray
radiation with a suitable detector. Optionally, an energy-dispersive detector
can also be
used to deten;nine the energy of the incident photons.
As energy-dispersive detectors and detector units, in principle two designs
are
available:
a. Energy-dispersive detectors according to the type of Cd(Zn)Te detectors, as
described in the introduction of the description. With such a series of
detectors, x-ray spectra of the emitted x-ray radiation can be measured pixel
for pixel.
b. Simple x-ray detectors are used. A discriminator, which in the simplest
case
consists of a suitable filter combination, is arranged in front of the
detector.
For energy selection, however, monochromators that are adjusted to, for
example, the x-ray fluorescence of the administered contrast medium, can
also be used.
c. It is also definitely technically possible, however, to adapt the detector
directly to the contrast medium. Thus, Gd(Zn)Te detectors or Dy(Zn)Te
detectors can be used.
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In all cases, the detector is positioned as much as possible so that a minimum
of
the Compton scatter is measured.
To determine the values of the intensity as well as the energy of the x-ray
radiation emitted by the object under examination, the detected photons are
divided into
at least two different energy ranges, which contain, for example, the Ka and
the KB
emission lines. To increase the element specificity, a Compton correction
optionally
can be performed. As the examples further indicated below show, this is not
always
necessary, however.
If a native x-ray contrast is ignored, an x-ray contrast medium can be
administered to the object under examination, for example a human, to perform
the
process according to the invention. The x-ray contrast medium can be
administered, for
example, enterally or parenterally, especially by i.v., i.m. or subcutaneous
injection or
infusion. Then, the x-ray image is made. Those contrast media that exhibit
high
attenuation coefficients in the selected spectral area per se are suitable.
Contrast media
whose absorbing element has the K-edge of the absorption spectrum in the
selected
spectral range are also especially suitable. Such x-ray contrast media contain
opacifying chemical elements with an atomic number of 35 or greater than 35 -
in this
case, for example, this is a contrast medium that contains bromine - with an
atomic
number of 47 or greater than 47 - in this case, this is a contrast medium that
contains
iodine -- , with an atomic number of 57 or greater than 57 - in this case this
is a contrast
medium that contains lanthanides, especially a contrast medium that contains
gadolinium - or with an atomic number of 83 - in this case this is a contrast
medium
that contains bismuth. Therefore, x-ray contrast media that contain opacifying
chemical
elements with an atomic number of 35 (bromine) to 83 (bismuth) are suitable.
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Especially suitable are contrast media with opacifying chemical elements with
an
atomic number of 53 (iodine) - 83 (bismuth). Also suitable are x-ray contrast
media
with opacifying chemical elements with an atomic number of 57 or greater than
57
(lanthanides) - 83 (bismuth) and especially preferably media with opacifying
chemical
elements with an atomic number of 57-70 (lanthanides: La, Ce, Pr, Nd, Pm, Sm,
Eu,
Gd, Th, Dy, Ho, Er, Tm, and Yb).
Suitable iodine-containing x-ray contrast media are, for example, compounds
that contain triiodine aromatic compounds, such as, for example,
amidotrizoate, iohexol,
iopamidol, iopanoic acid, iopodinic acid, iopromide, iopronic acid, iopydone,
iothalamic
acid, iopentol, ioversol, ioxaglat, iotrolan, iodixanol, iotroxic acid,
ioxaglic acid and
ioxithalamic acid and iosimenol (fNN). Trade names for x-ray contrast media
that
contain iodine are Urografiri (Schering), Gastrografin (Schering), Biliscopin
(Schering), Ultravist (Schering) and Isovist (Schering).
Also suitable as x-ray contrast media are metal complexes, for example
Gd-DTPA (Magnevist (Schering)), Gd-DOTA (Gadoterate, Dotarem), Gd-HP-DO3A
(Gadoteridol, Prohance (Bracco)), Gd-EOB-DTPA (Gadoxetat, Primavist), Gd-
BOPTA (Gadobenat, MultiHance), Gd-DTPA-BMA (Gadodiamide, Omniscari
(Amersham Health)), Dy-DTPA-BMA, Gd-DTPA-polylysine, and Gd-DTPA-cascade
polymers, i.a., whereby DTPA = diethylenetriaminepentaacetic acid, DOTA =
1,4,7,10-
tetraazacyclododecane, HP-D03A = 10-(hydroxypropyl)-1,4,7,10-
tetraazacyclododecane-1,4,7-triacetic acid), EOB-DTPA = 3,6,9-triaza-3,6,9-
tris(carboxymethyl-4-(4-ethoxybenzyl)undecanedicarboxylic acid, BOPTA = (4-
carboxy-5,8,11-tris(carboxymethyl)-1-phenyl-2-oxa-5,8,11-triazatridecan-
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13-oic, benic acid), DTPA-BMA = diethylenetriaminepentaacetate-
bis(methylamide),
DTPA-polylysine = diethylenetriaminepentaacetate-polylysine, DTPA-cascade
polymers.
The x-ray contrast media can be administered enterally and parenterally. In
the
case of parenteral administration, the intravenous (i.v.) administration is
preferably
selected. Preferred dosages are doses up to 0.75 g of 1/kg of body weight in
the iodine-
containing non-ionic contrast media. This corresponds to approximately 6 mmol
of 1/kg
of body weight. In addition, the dose can preferably be increased to 1.5 g of
1/kg of
body weight (corresponding to approximately 12 mmol of 1/kg of body weight)
and in
exceptional cases up to 2 (corresponding to approximately 16 I) or 5 g of 1/kg
of body
weight (corresponding to approximately 39 mmol of 1/kg of body weight). In the
case
of the lanthanide complexes, the preferred dose is approximately 0.1 mmol/kg
of body
weight. Doses up to 0.3 mmol/kg of body weight or up to 1 mmol/kg of body
weight
are suitable and, in addition, also preferred.
The emission lines of gadolinium are approximately 43.0 and 48.7 keV, i.e.,
far
above the emission lines of iodine, which are approximately 28.6 and 32.3 keV.
Instead
of the gadolinium atoms, the metal complexes can also contain, for example,
all other
lanthanides, such as lanthanum, dysprosium or ytterbium.
Digital detectors have already been offered for some time by various
manufacturers (for example: The BBI Newsletter, February 1999, page 34; H. G.
Chotas, J. T. Dobbins, C. E. Ravin, "Principles of Digital Radiography with
Large-Area,
Electronically Readable Detectors: A Review of the Basics," Radiol., 210
(1999) 595-
599). They often consist of amorphous silicon or other semiconductor
materials. In the
x-ray arrangement according to the invention, i.a., the following detectors
are suitable:
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detectors with phosphorus plates (for example from Fuji Chemical Industries,
Konica),
with amorphous silicon (for example from GE Medical, Philips Medical, Siemens
Medical), with salts (for example from Philips Medical, Toshiba), with
gadolinium
hyposulfite (for example from Kodak), with cadmium telluride (CT) or cadmium-
zinc-
telluride-(CZT) semiconductors, with yttrium oxyorthosilicate, with lutetium
oxyorthosilicate, with sodium iodide or bismuth germanate. Especially good
results are
achieved with the so-called C(Z)T detectors, i.e., detectors that consist of a
cadmium-
(zinc)-telluride-(C(Z)T) semiconductor.
The design of an energy-dispersive detector, which is formed from a
semiconductor, is described in detail in US 5,434,417 A. In this case,
segmented
semiconductor strips that are irradiated from the front with the x-ray
radiation are
provided. The radiation is drawn through the semiconductor material until it
interacts
with the semiconductor material. The penetration depth depends on the energy
of the x-
ray photons. In the case of greater energy of the x-ray photons, the radiation
is drawn
through more deeply -- until it interacts with the detector material and
generates a
current impulse by a photoelectric effect -- than with lower energy of the x-
ray photons.
The current impulses can be discharged in the individual segments of the
detector by
means of applied electric contacts. The current impulses are processed with an
input
amplifier.
On the one hand, the detector can be designed in the form of a flat-bed
detector.
In this embodiment, all pixels are detected simultaneously and passed on to
the
correlation unit for evaluation. In this case, the detector consists of a
large-area
arrangement of individual detector sensors, preferably in a matrix that has
rows and
columns of such sensors.
CA 02610964 2007-12-06
18
In addition, a detector unit that is used to determine the emitted x-ray
radiation
and optionally to record an emission image and that for this purpose is
designed with an
x-ray-optical module for energy selection can also be provided.
Instead of the flat-bed detector, line detectors or a matrix of several
detectors
that are suitable for picking up an individual pixel can also be used. In the
case of more
recent detectors, the x-ray radiation from the object under examination is
simultaneously sent via an x-ray fiber optic light guide. A number of such
fiber optic
light guides are combined in a surface detector.
In addition, the detector can be designed for picking up an individual pixel
and
can be movable so as to pick up all pixels. In this embodiment, the detector
can detect
only energy-dependent intensities in an individual pixel during the
measurement. The
intensities of the individual pixels are detected in succession, for example
by lines, and
are passed on to the correlation unit for further processing.
In addition, the detector can also have an array of detector sensors designed
for
picking up a pixel in each case and can be movable so as to pick up all
pixels.
According to this invention, both a line of detector sensors and another
arrangement, for
example a matrix-like arrangement, of detectors sensors, are defined as arrays
of
detector sensors. In this embodiment, the detector detects the intensity
values in the
individual pixels by lines or optionally also by blocks. To pick up all
intensity values,
the detector is preferably moved perpendicular to the main axis of the array
during the
measurement. The intensity values that are determined during the measurement
are
forwarded to the correlation unit.
For graphic display of, for example, the distribution of opacifying chemical
elements in the object under examination, it is advantageous to detect the
radiation
CA 02610964 2007-12-06
19
intensities with, in each case, the same weighting, that are emitted by the
respective
space elements. In addition, for this purpose, it is also advantageous to load
the
respective space elements with, in each case, the same radiation intensity
from the x-ray
radiation source. In practice, these premises turn out to be only
approximately correct,
since, on the one hand, the irradiated x-ray radiation is attenuated by
absorption to
differing extents, depending on how much of the distance from the radiation
lies in the
object under examination, and, on the other hand, the radiation that is
emitted by the
space elements in the object under examination is attenuated by self-
absorption to
differing extents, depending on how much of the distance between them and the
detector lies in the object under examination.
This problem occurs in all emission-spectroscopic methods. To solve the
problem, the second intensity values are first corrected taking into
consideration the
absorption of irradiated x-ray radiation andlor the self-absorption of the
emitted x-ray
radiation in the object under examination, and the first and second intensity
values are
correlated with one another only after this correction, pixel for pixel. Such
a correction
can be performed by means of numerical processes by the geometry of the object
under
examination and an at least approximately position-dependent x-ray opacity
being taken
into consideration. To determine the position-dependent x-ray opacity, the
images
generated from the first intensity values can be used. To determine the
position-
dependent absorption and self-absorption, the position-dependent x-ray opacity
that is
obtained from this measurement can be taken as the baseline in a first
approximation,
since the absorption coefficients for the irradiated x-ray radiation is
similar to that of the
emitted radiation.
CA 02610964 2007-12-06
Because of the self-absorption of the emitted radiation, it may further be
advantageous to move the position and angular position of the second detector
during
the measurement relative to the area of examination (ROI), for example on a
circular
segment pathway to offset structural inhomogeneities in the object under
examination
that have a varying absorbing action depending on the angle of observation and
observation point. In this case, the graphic displays would be obtained after
correction
of the self-absorption was done by taking averages.
The signal that originates from the input amplifier is then sent into at least
one
correlation unit, with which the intensity of the detected x-ray radiation
from a pixel of
the object under examination is correlated with the image of the emitted x-ray
radiation
(x-ray scatter and x-ray fluorescence) of the same pixel. The correlation unit
can be a
correspondingly programmed data-processing unit.
To correlate the intensity values of the photons of the two modalities
(transmission image and emission image), the latter are correlated with one
another a
pixel at a time, preferably subtracted from one another or divided by one
another. To
this end, in one case, a comparator can be used, and in the other case, a
division term
can be used for correlation that is performed pixel for pixel. Of course,
other
mathematical operations can also be performed for correlation of the intensity
values of
the transmitted and emitted x-ray radiation from an image.
To process the measured intensity values of a pixel, preferably the following
devices that can be implemented in a data-processing unit are provided,
namely:
dl. A first storage unit, with which the first intensity values of the
transmitted
x-ray radiation can be stored pixel for pixel,
CA 02610964 2007-12-06
21
d2. A second storage unit, with which the second intensity values of the
emitted
x-ray radiation can be stored pixel for pixel (e.g., with the elements I, Gd
and
Yb),
d3. A computing unit, which provides for a suitable correlation of the two
generated image data sets and thus generates or computes an image data set
from
the information of the transmission data set and the data from the x-ray
emission, preferably x-ray fluorescence.
As a result, it is possible to correlate the intensity values of all pixels in
transmission and emission with one another, whereby the emission image is
adapted via
the characteristic emission lines to the contrast medium that is used. If a
mixture that
consists of x-ray contrast media (e.g., Ultravist and Gadovist'11) or
substances that
contain both iodine and a lanthanide (such as Gd or Dy) are used, the emission
lines that
are characteristic in each case can be used for emission imaging, whereby the
measured
data sets are then correlated with one another by pixels and are used for
graphic display,
or whereby alternatively, the respective intensity values are correlated with
one another
pixel for pixel, and the data that are obtained are then used for graphic
display. To this
end, the data that are obtained are delivered a pixel at a time to an output
unit, which
contains, for example, a monitor (CRT or LCD display) or a plotter.
For a more detailed explanation of the invention, the following figures and
examples are used. To provide a direct illustration of how the invention
works, in no
case was any effort made to correct the measured x-ray spectra according to
the
absorption of the excitation beam and self-absorption. Here, in detail:
Fig. I shows an image of a test arrangement in a computer tomograph;
CA 02610964 2007-12-06
22
Fig. 2 shows a diagrammatic visualization of the imaging arrangement or of the
experimental set-up;
Fig. 3 shows a diagrammatic visualization of the test arrangement for
generating the first phantom measurements;
Fig. 4 shows emission spectra of the phantom of Fig. 3, filled with water
(Fig.
4a), Ultravist (Fig. 4b), Gadovist (Fig. 4c);
Fig. 5 shows emission spectra of the phantom of Fig. 3, filled with water
(Fig.
5a), Ultravist (Fig. 5b), Gadovist (Fig. 5c), whereby in each case a
PMMA disk that is 5 cm thick was arranged between the detector and the
phantom;
Fig. 6 shows the intensity of the emission based on the position/shift of the
phantom from Fig. 3 in selected energy bands (corresponding to the Ka
and KB lines (iodine: Fig. 6a, gadolinium: Fig. 6b, mixture that consists
of iodine and gadolinium: Fig. 6c);
Fig. 7 shows CT cross-sectional images (transmission images) of the phantom
filled with Gd, an iodine/Gd mixture, iodine, air and water.
In Fig. 1, a photographic visualization of a test arrangement in a computer
tomograph with a rubber ball 1, which is fastened to a rack 2, is shown. The
rubber ball
is arranged in the center of the computer tomograph. In various tests, the
rubber ball
was filled with air, water, as well as different contrast medium solutions.
The ball was
found between the CT tubes (above the rubber ball; not shown) and a line
detector
(below the table that is visible under the rubber ball, not visible).
At an angle of 90 to the connecting line between the CT tubes, the rubber
ball
and the detector, a measuring chamber 3 was positioned for detection of the x-
ray
CA 02610964 2007-12-06
23
fluorescence. With this experimental set-up, a tissue, tumor or the like
filled with
contrast medium was simulated as an object under examination, which is
examined in
the computer tomograph. To this end, the object was scanned in layers, and in
this case
the scatter spectra were measured.
The experimental set-up used in this test is shown in detail in Fig. 2. The
diagrammatic drawing shown there shows the ball 1, which is found as a phantom
in the
isocenter of the gantry 4. The CT tube 5 was arranged in the 12 o'clock
position and
remained fixed there. The measuring chamber 3, consisting of a detector 6 and
a lead
tube 7, was directed at an angle of 90 to the x-ray cone beam that protrudes
from the
CT tube to the phantom (ball) (in z-direction; see arrow).
To detect the x-ray radiation, a CZT detector 6 with a 3 mm x 3 mm x 2 mm
cadmium-zinc-telluride crystal and 100/400 m apertures was used (Amptek,
Inc.,
USA). The data recorded by the fluorescence detector were conveyed from the
detector
via an amplifier to a multichannel analyzer 9 and then fed to an Excel
(Microsoft)
spreadsheet, which was stored on a PC 10. The signal intensities SI = SI(E)
were thus
available in digital form as a function of the energy E.
In Fig. 3, a diagrammatic visualization of the test arrangement for generating
the
first phantom measurements is shown. A portion of the measuring chamber for
measuring the fluorescence 3 can be seen to the left in the visualization,
while the ball I
is shown in the center of the visualization. The individual sectional planes
that are
perpendicular in Fig. 3, from which the fluorescence goes into the measuring
chamber,
were produced by an x-ray fan beam incident from above. The dotted lines mark
the
respective positions of the CT tubes above the image cutaway. The horizontal
scale
CA 02610964 2007-12-06
24
indicates the shift of the fan beam and thus indicates the sectional planes
addressed in
each case (excited layer) in the ball.
A "zero measurement" was made at +45 mm and thus outside of the excitation
beam.
After each recording of a spectrum, the entire measuring structure was moved
10
mm further into the gantry (in the z-direction), and the new spectrum was
recorded.
Thus, various spectra were produced in layers based on the respective position
of the
ball in the beam or corresponding to the ball geometry.
With this measuring structure, the x-ray fluorescence thus could be measured
based on the topography of the phantom, whereby at z = -60 mm, the layer
closest to the
detector was irradiated, and at z = 0, the layer farthest from the detector
was irradiated
(at z=-60, the self-absorption of the emission is thus minimal and at z= 0, it
is
maximal; because of the spherical geometry, an absorption effect in the
irradiation is
made noticeable at higher contrast medium concentrations).
Example 1:
In a first measurement, the ball was filled with water and measured at 80 kV,
50
mA for each 80 s per position of the ball in the beam corresponding to Fig. 3
(parameters: detector: XR-100.CZT (aperture 0.1 mm), ball-detector distance:
18.0
cm; ball-CT tube distance: 32.0 cm).
In Fig. 4a, the scatter spectra of the water in the phantom are depicted for
the
various positions.
CA 02610964 2007-12-06
In a second measurement, the ball was filled with a solution of 50 mmol/1 of
iodine in water (Ultravist ) and measured at 80 kV, 50 mA for each 80 s per
position
(parameters: detector: XR-100.CZT (aperture 0.1 mm)).
The emission spectra obtained in the various positions are reproduced in Fig.
4b.
The Ka and KB lines of iodine (28.6 and 32.3 keV) can be seen clearly. From
the graph,
a dependency of the measured intensity of the x-ray fluorescence on the
geometry of the
phantom is clear. The larger the irradiated layer of the phantom was, the
higher the
measured intensity.
In a third measurement, the ball was filled with a solution of 50 mmol/1 of
gadolinium in water (Gadovist ) and measured at 80 kV, 50 mA for each 80 s per
position (parameters: detector: XR-100.CZT (aperture 0.1 mm)).
The emission spectra obtained in the various positions are reproduced in Fig.
4c.
The Ka and KR lines of gadolinium (43.0 and 48.7 keV) can be seen clearly. It
was
shown that the intensity of the measured emission radiation especially in the
range of
the K lines depends on the geometry of the ball in the radiation field.
Example 2:
In the individual measurements of this test, in each case a 5 cm thick PMMA
disk was positioned as a filter between the detector and the phantom to
simulate the
self-absorption of the x-ray fluorescence radiation through the surrounding
tissue.
In Fig. 5a, the scatter spectra of the water in the phantom are depicted for
the
various positions.
CA 02610964 2007-12-06
26
In a second measurement, the ball was filled with a solution of 50 mrnol/l of
iodine in water (Ultravist ) and measured at 80 kV, 50 mA for each 80 s per
position
(parameters: detector: XR-100.CZT (aperture 0.1 mm)).
The emission spectra obtained in the various positions are reproduced in Fig.
5b.
The intensity of the fluorescence radiation decreased because of the inserted
PMMA
disk. It was verified that the intensity was all the lower the thicker the
disk was.
However, even in the largest layer of the ball ( in the center), the K lines
were still
measurable.
In a third measurement, the ball was filled with a solution of 50 mmol/1 of
gadolinium in water (Gadovist ) and measured at 80 kV, 50 mA for each 80 s per
position (parameters: detector: XR-100.CZT (aperture 0.1 mm)).
The emission spectra obtained in the various positions are reproduced in Fig.
5c.
Also here, the fluorescence radiation decreased because of the inserted PMMA
disk.
Since the Ka and KB lines of gadolinium are approximately 43.0 or 48.7 keV, a
considerably more intensive fluorescence radiation could be detected with the
presence
of the 5 cm-thick PMMA disk than in the case of the iodine emission as before.
Therefore, even in this case, the K lines can still be measured in the largest
layer of the
ball (in the center).
Example 3:
In another test, the intensity values of the fluorescence were determined and
recorded based on the positioning of the ball relative to the x-ray beam.
In a first measurement, the ball was filled with a solution of 50 mmol/1 of
iodine
in water (Ultravist ) and measured at 80 kV, 50 mA for each 80 s per position.
CA 02610964 2007-12-06
27
In Fig. 6a, the intensity of the fluorescence radiation is plotted based on
the
position/shift of the phantom in selected energy bands corresponding to the Ka
line of
iodine at 28.6 keV and the Kp line of iodine at 32.3 keV. The profile of the
emission
intensity produced by the shape of the ball can be detected from this figure.
In a second measurement, the ball was filled with a solution of 50 mmol/1 of
gadolinium in water (Gadovist ) and measured at 80 kV, 50 mA for each 80 s per
position.
In Fig. 6b, the intensity of the fluorescence radiation is plotted based on
the
position/shift of the phantom in selected energy bands corresponding to the
Ka, line of
gadolinium at 43.0 keV and the Ko line of gadolinium at 48.7 keV. The profile
of the
emission intensity produced by the shape of the ball can also be detected from
this
figure.
In a third measurement, the ball was filled with a solution of 25 mmol/1 of
iodine
(Ultravist ) and 25 mmol/1 of gadolinium in water (Gadovist ) and measured at
80 kV,
50 mA for each 80 s per position.
In Fig. 6c, the intensity of the fluorescence radiation is plotted based on
the
position/shift of the phantom in selected energy bands corresponding to the Ka
line of
iodine at 28.7 keV, the KR line of iodine at 32.3 keV, the Ka line of
gadolinium at 43.0
keV, and the Kp line of gadolinium at 48.7 keV. As can be seen in Fig. 6c, the
ball
profile is only insufficiently reproduced in the direct plotting of the signal
intensity as a
function of the position. This can be attributed to the absorption on the
excitation side
and the self-absorption on the emission side, which distort the image. Lower
contrast
medium concentrations and corrections of the absorption of the primary beam
and self-
CA 02610964 2007-12-06
28
absorption of the x-ray fluorescence resulted in a graphic visualization of
the ball in one
dimension.
Example 4:
Fig. 7 shows the CT cross-sectional images that are recorded in the preceding
examples of the x-ray fluorescence. From the top left to the bottom right, the
ball that is
filed with gadolinium, the ball with the mixture that consists of gadolinium
and iodine,
the ball with iodine, the ball with pure water, and the ball filled with air
can be seen.
The air-filled ball has clearly the smallest x-ray attenuation, followed by
the water-filled
ball. When using the ball with 50 mmol/I of opacifying element, the x-ray
attenuation is
more pronounced than that with water; a quantitative evaluation is possible
via the
determination of the Hounsfield units (HU), but only the addition of the x-ray
fluorescence images allows an assessment on the element-specific filling of
the ball.
