EP1890604A1

EP1890604A1 - X-ray arrangement for the image representation of an examination object and use of the x-ray arrangement

Info

Publication number: EP1890604A1
Application number: EP06724603A
Authority: EP
Inventors: Rüdiger LAWACZECK
Original assignee: Bayer Schering Pharma AG
Current assignee: Bayer Intellectual Property GmbH
Priority date: 2005-06-06
Filing date: 2006-04-20
Publication date: 2008-02-27
Also published as: RU2007148223A; MX2007015406A; CA2610964A1; DE102005026940A1; KR20080020616A; JP2008541963A; RU2449729C2; BRPI0611181A2; CN101203179A; CN101203179B; CN102512188A; IL187846A0; AU2006254747A1; WO2006131175A1

Abstract

For the high-contrast representation of small lesions or other target areas in tissue in the human body containing at least one contrasting chemical element, an X-ray arrangement is described, comprising at least one X-ray radiation source which emits essentially polychromatic X-ray radiation, a first detector or a plurality of first detectors, which can be used to determine values of a first intensity of X-ray radiation transmitted through the examination object, a second detector or a plurality of second detectors, which can be used to determine values of a second intensity of X-ray radiation emitted by the examination object, at least one correlation unit, which can be used to correlate the first intensity values of the transmitted X-ray radiation with the second intensity values of the emitted X-ray radiation pixel by pixel, and at least one output unit for the representation of the examination object from pixel signals obtainable by correlation of the first intensity values with the second intensity values. The transmission and emission images are preferably recorded simultaneously. The method can also be combined with other radiological images, e.g. positron emission tomography (PET) or single photon emission computer tomography (SPECT).

Description

X-ray arrangement for image presentation of an examination subject and use of the X-ray arrangement

Description:

The present invention relates to an X-ray arrangement for image display of an examination object containing at least one X-ray contrasting chemical element by means of X-radiation, a use of the X-ray arrangement and an X-ray contrast imaging method on an examination subject, for example a mammal, in particular a human.

X-ray medical diagnostics is a technically advanced field for the diagnosis of diseases, for example for the early detection, radiographic detection, characterization and localization of tumors, vascular diseases and other pathological changes of the human body. The technology is very powerful and has a high availability.

To generate the X-radiation, X-ray tubes are available, for example, with W, Mo or Rh rotary anodes and Al, Cu, Ti, Mo and Rh filters. With suitable filtering, a portion of the bremsstrahlung is filtered out so that in favorable cases substantially the characteristic radiation exits the x-ray tube.

The detectors used are either conventional X-ray films, imaging plates or digital flatbed detectors. Computer tomographs use one or more lines of detector. Also several detectors can be connected in parallel. Semiconductor detectors consisting of cadmium telluride (CT), cadmium zinc telluride (CZT), amorphous selenium, or amorphous or crystalline silicon (MJ Yaffe, JA Rowlands, "X-Ray Detectors for Digital Radiography ", Med. Biol., 42 (1) (1997) 1-39). An example of the construction of such detectors is given in US Pat. No. 5,434,417. In order to also enable an energy sensitivity of the detector, this is formed of several layers. X-ray radiation with different energy penetrates into this detector at different depths and generates in the respective layer by photoelectric effect an electrical signal which can be identifiably read out as a current pulse after the layer and thus according to the energy of the X-ray photons.

Computed tomography (CT) has long been used as a routine procedure in everyday clinical practice. With the CT, sectional images are obtained by the body, with which a better spatial resolution is achieved than with the conventional projection radiography. Although the density resolution of CT is significantly higher than the density resolution of conventional X-ray technology, contrast agents are still needed to reliably detect many pathological changes. These improve the quality of the morphological information. The contrast agent on the one hand shows functional processes in the body (excretion, perfusion, permeability) and on the other emphasizes the morphology by creating contrasts (different contrast agent concentrations in different tissues).

In many cases, the conventional X-ray technique could not be used because the contrast of the tissue to be examined was not sufficient. For this purpose, X-ray contrast agents have been developed which produce a high X-ray density in the tissue in which they accumulate. Typically, iodine, bromine, elements of atomic numbers 34, 42, 44-52, 54-60, 62-79, 82 and 83 are used as contrasting elements and the chelate compounds of elements with atomic numbers 56-60, 62-79, 82 and 83 proposed. As iodine compounds, for example, meglumine-Na or lysine diatrizoate, lothalamate, loxithalamate, lopromide, lohexol, lomeprol, lopamidol, ioversoi, iobitridoi, iopentoi, iotrolan, iodixanol and loxilane (INN) are used (EP 0 885 616 A1).

In some cases, despite the administration of X-ray contrast agents, sufficient tissue contrast could not be achieved. To further enhance contrast, digital subtraction angiography (DSA) was introduced, subtracting pre- and post-contrast (logarithmic) images from each other. A subtraction method for use in mammography is disclosed in EP 0 885 616 A1: For projection mammography, there is proposed, first of all, an

Contrast mammogram, then the patient a conventional urographic X-ray contrast agent rapidly i.v. inject and record a post-contrast mammogram for about 30 seconds to 1 minute after the end of the injection. The obtained data of the two images are then correlated with each other, preferably subtracted from each other.

New developments in the field of CT concern, for example, the application of synchrotron radiation in CT on the excitation side (FA Dilmanian, "Computed Tomography with Monochromatic X-Rays", Am. J. Physiol. Imaging, 314 (1992) 175-193). Good X-ray images are, for example, using the "K-edge

Subtraction CT "(FA Dilmanian, supra, page 179), exploiting the strong increase in the absorption coefficient at the binding energy of the K electrons of an atom The element iodine has a K edge at an energy of 33.17 keV Unfortunately This method works only with the help of the synchrotron radiation available at large storage rings, such as DESY, because only this radiation has the monochromatic and intensity favorable for the process Conventional X-ray tubes do not provide monochromatic radiation but a continuous spectrum not well suited.

An alternative possibility is described in DE 101 18 792 A1: Here, a method is proposed for recording projection mammograms X-ray tube with two X-ray anodes of different materials are used. To record the mammograms, the patient is first given an X-ray contrast agent. Then, a first projection image is picked up using the first of the two X-ray anodes, and then a second projection image using a second X-ray anode. By overlaying 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 agent: the emission energy of the first X-ray anode is slightly below the absorption energy of the contrasting element in the X-ray contrast medium and the emission energy of the second X-ray anode is slightly above the absorption energy of the contrasting element. A disadvantage of this method is that conventional x-ray tubes with only one x-ray anode must be replaced with bi-anode tubes.

In addition to transmission radiography, an emission radiography is also described:

Thus, WO 2004/041060 A2 describes a device for the non-invasive in vivo determination of a chemical element in the prostate of a human, comprising a probe, an irradiation system with which the chemical element can be excited to emit radiation, a radiation detector within the probe with which the emitted radiation can be imaged and a signal recording, processing and display system with which the amount of the chemical element in the prostate can be reproduced at different locations corresponding to the image of the emitted radiation. The emitted radiation essentially consists of fluorescence radiation. In the case of examination of the prostate, the distribution of Zn in the tissue is preferably determined.

Furthermore, in DE 36 08 965 A1 a method for determining the proportion of various chemical elements in a layer of Research area described by gamma or X-rays. The Compton and Rayleigh scattered radiation are recorded separately. The course of the differential scattering coefficients determined from the measured values is influenced by the proportions of the various chemical elements contained in the individual picture elements. From this, therefore, the proportion of these chemical elements can be determined. For this purpose, the examination area is penetrated by a primary beam from a multiplicity of directions, and the radiation emerging from the examination area at different angles is detected by a detector arrangement in various positions outside the examination area, after which the differential

Scattering coefficient for different momentum transfers from the measured values obtained for each pixel of the layer is determined.

Further, Quanwen Yu et al., "Preliminary Experiment of Fluorescent X-Ray Computed Tomography to Detective Dual Agents for Biological Study": J. Synchrotron Rad. (2001), 8 1030-1034, has proposed the X-ray fluorescence method for determination This method can be used to obtain images that can simultaneously detect multiagents using the fluorescence Kα line in a single study, such as brain blood flow and density In the presented study, images obtained by this method were compared with images obtained by X-ray transmission tomography.

The X-ray fluorescence method or X-ray diffraction method described in the publications mentioned above, however, has the disadvantage that a representation of small details in an examination subject is not readily possible because of difficulties in imaging. Rather, only roughly resolved representations are obtained, so that smaller details can hardly be depicted. The present invention is therefore based on the problem to avoid the aforementioned disadvantages and in particular to find arrangements and methods with which recordings can be made with different X-ray contrasting chemical elements. Furthermore, the X-ray images should also be receivable in a simple, convenient manner, without incurring high costs. The technology should be available on a broad basis. Even smaller lesions in the examination subject should be able to be visualized with high spatial resolution with the lowest possible radiation dose. Motion artifacts should be avoided.

This problem is solved by the X-ray arrangement for imaging an X-ray examination object containing at least one X-ray-contrasting chemical element according to claim 1, the use of this X-ray arrangement according to claim 11 and the X-ray contrast imaging method according to claim 25. Preferred

Embodiments of the invention are specified in the subclaims.

Insofar as the terms "emission" and "emit" are used below in the description of the invention and in the claims, they are intended to include, for example, X-ray fluorescence, i. the emission of radiation after a

Excitation of the irradiated matter by means of electromagnetic radiation, and on the other preferably Rayleigh scattering understood. In the latter case, the radiation is emitted again without momentum transfer from the irradiated matter, although the irradiation does not cause any excitation of sheath electrons in atoms of this material in excited states, as in fluorescence.

With the X-ray arrangement, X-ray radiation transmitted through and emitted by the examination subject is used for imaging. For this purpose, the X-ray arrangement according to the invention comprises:

a. at least one substantially polychromatic X-ray emitting X-ray source, b. a first detector or a first detector unit (consisting of a plurality of detectors connected in parallel and / or arranged) with which the values of a first intensity of the X-ray radiation transmitted through the examination object can be determined, c. a second detector or a second detector unit, with which or with the values of a second intensity of the X-radiation emitted by the examination object 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 the second intensity values of the emitted X-ray radiation pixel by pixel, and e. at least one output unit for displaying the examination subject from pixel signals obtainable by correlation of the first intensity values with the second intensity values.

The transmitted X-radiation and the emitted X-radiation can be detected either simultaneously or sequentially.

This X-ray arrangement can advantageously be used for imaging an X-ray examination object containing preferably at least one X-ray-contrasting chemical element. The X-ray contrasting chemical element is preferably introduced into the examination subject by means of an X-ray contrast agent and, for example, for this purpose administered to the examination subject, for example a human or an animal.

The low-atom-count contrasting chemical elements naturally occurring in an object under examination have only a small one

Yield of X-ray fluorescence, so that the imaging using these elements does not seem practicable. In addition, the energy of the X-ray fluorescence photons in this case is low, so that their Range in body tissue is low. In particular from the element iodine (Z = 53) with the emission lines 28.6 and 32.3 keV, fluorescence lines are available which leave the examination subject to a sufficient extent and can thus be registered by a detector arranged outside the object. In cases of lower atomic number of the chemical element, an arrangement of the second detector is to be selected in which this is arranged as close as possible to the region to be examined (ROI: region of interest).

The X-ray arrangement is used for carrying out the X-ray contrast method according to the invention. The method has the following procedural steps:

a. Preferably administering at least one chemical element giving an X-ray contrast, b. Radiating the examination object with substantially polychromatic X-ray radiation, c. Determining values of a first intensity of the X-radiation transmitted by the examination subject, d. Determining values of a second intensity of the X-radiation emitted by the examination subject, e. Correlating the first intensity values of the transmitted X-ray radiation pixel by pixel with the second intensity values of the emitted X-radiation, f. Representation of the examination object from obtained by correlation of the first intensity values with the second intensity values

Pixel signals.

In contrast to the known methods in which either only X-ray transmission tomography (TXCT) is performed or only X-ray fluorescence (FXCT) is detected, transmission and emission are measured simultaneously or sequentially, and these two techniques are combined in accordance with the present invention each image obtained by a suitable correlation method be superimposed. This procedure uses the respective advantages of both techniques:

Although the X-ray transmission tomography offers the advantage of a high achievable temporal and spatial resolution, so that in principle even the smallest lesions or other details can be resolved in an examined human body. However, the contrast obtained is often not sufficient to visualize these details as well. This applies in particular to examinations of lesions in the soft tissue. In addition, examinations of certain body regions with the TXCT method are also affected by the skeleton.

On the other hand, the X-ray fluorescence tomography offers the advantage of an extraordinarily high-contrast representation, since only certain chemical elements emit electromagnetic radiation with suitable excitation of these elements, so that these elements located in the examination region (ROI) are suitable as extremely sensitive measuring probes. However, the FXCT method suffers from the disadvantage of a low spatial resolution, so that smaller lesions can no longer be displayed.

Only by correlating the intensity values of the transmitted X-ray radiation pixel by pixel with the intensity values of the emitted X-radiation and displaying the examination subject from pixel signals which have been obtained by this correlation can a high-contrast and detailed image of the examination region (ROI) be generated. Although the image part resulting in the contrast has a low resolution. By correlating the respective values with each other, however, this deficiency is largely eliminated, since the required detailed information comes from the intensity values of the radiation measured by TXCT.

The invention can be used in particular for human examination. The invention is for the production of radiographs for the representation of Masses, vessels and perfusions suitable, for example, the representation of the esophageal gastrointestinal passage, bronchography, cholegraphy, angiography and cardiograography, cerebral angiography and perfusion measurements, mammography and lymphography. The focus of the application technology of the invention lies in the

Computed tomography (MS-CT; μCT) and its fusion modalities (PET-CT (positron emission tomography), SPECT (Single Photon Emission Computer Tomography), sonography and other methods of optical imaging). In principle, the invention can also be used for the investigation of non-living materials, for example in the field of material testing.

To carry out an investigation, the transmitted radiation is recorded by means of the first detector, which is located in the beam path of the x-ray tube which is attenuated by the examination subject. The emitted radiation is measured by means of the second detector, which is arranged outside this beam path, preferably at an angle of approximately 90 ° to the beam path. In principle, however, this second detector can also be arranged in any other angular position relative to the X-ray beam, for example 45 ° or 135 ° to the beam emanating from the X-ray radiation source, without however being detected by the beam passing through the examination subject. When the x-ray tube is in the 12 o'clock position, ordinary CT scanners are equipped with a series of detectors at the opposite 6 o'clock position. Preferably, the second detector can be arranged in the 3 o'clock position and / or the 9 o'clock position. Both X-ray fluorescence and X-ray scattering (Rayleigh scattering, Compton scattering) can be recorded by means of this second detector.

For selective image acquisition with the second detector by utilizing the emitted X-radiation, the emitted X-radiation can be measured in terms of its energy resolved. It is particularly advantageous, in the presence of a given emitting chemical element in Object to be examined to discriminate the X-ray received from the second detector, which originates from this contrasting element of other emitted X-rays, for example, scattered radiation (Compton, Rayleigh radiation) and derived from other chemical elements fluorescent radiation. This makes it possible to make certain areas (ROI) very selectively visible by utilizing, for example, the accumulation of contrasting chemical elements in certain organs of a human body, so that a particularly large contrast of the visualized tissue to surrounding tissue arises. The structure caused by the skeleton in an image representation also occurs in such a case compared to the representation of the tissue, so that the framework practically does not disturb the image representation.

For the detection and characterization of the emission radiation, an energy-dispersive detector is preferably used. However, it is also possible to use simpler detectors and to guarantee the characterization of the emission by X-ray optical modules (filter combination, monochromators).

This principle can also be applied in the same way to the measurement of

Intensity values of the transmitted X-radiation can be applied with the first detector. Also in this case, a selective representation of the areas in the examination subject (ROI) is achieved, in which accumulate the contrasting chemical elements.

Therefore, with the invention, soft tissue can be represented in high contrast, for example in humans. By matching the energy or energy interval of the transmitted and emitted X-radiation received by the detectors to the nature of the contrasting chemical element, an efficient

Contrast increase over conventional methods can be achieved. To generate the X-radiation, it is possible to use a normal, commercially available X-ray tube with a continuous spectrum, for example a tube with a Mo, W or Rh anode. Depending on the nature of the contrasting chemical element contained in the examination subject, a voltage is applied which enables an emission of the continuous radiation in the range up to, for example, more than 100 keV.

In principle, the X-ray radiation source can be operated without filtering the emitted radiation, so that polychromatic radiation impinges on the examination subject in the entire spectral range. In order to reduce the radiation exposure of the examination object, 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 required or advantageous for the detection. For this example, an AI or a Cu filter is used, the energy in the range <20 keV (soft radiation) filters out. The continuous spectrum is thus to be understood as meaning an X-ray emission in a range of ≥ 0 keV, preferably ≥ 15 keV, particularly preferably ≥ 17 keV and most preferably ≥ 20 keV, up to, for example, 100 keV, with no spectral range within these limits being highlighted over others or excluded. The upper limit of the

Emission spectrum is determined by the voltage applied to the x-ray anode. The low-energy region of the radiation is preferably filtered out in order to eliminate dose-relevant radiation for the human body.

Normally, the examination object is examined with polychromatic X-ray radiation with a suitable detector. Optionally, an energy dispersive detector can be used to determine the energy of the incident photons.

As energy-dispersive detectors and detector units there are basically two versions available: a. Energy-dispersive detectors of the type of Cd (Zn) Te detectors, as described in the introduction to the description. With such a series of detectors, X-ray spectra of the emitted X-ray radiation can be measured pixel by pixel. b. Simple X-ray detectors are used. In front of the detector, a discriminator is arranged, which consists in the simplest case of a suitable filter combination. For energy selection but also monochromators can be used, which are set for example on the X-ray fluorescence of the administered contrast agent. c. It is technically quite possible to adapt the detector directly to the contrast agent. Thus, Gd (Zn) Te or Dy (Zn) Te detectors can be used.

In all cases, the detector is positioned as much as possible so that a minimum of Compton scattering is measured.

In order to determine the values of the intensity as well as the energy of the X-ray radiation emitted by the examination subject, the detected photons are divided into at least two different energy ranges, which contain, for example, the K ₀ and the Kβ emission lines. If necessary, a Compton correction can be carried out to increase the element specificity. As the examples given below show, this is not always necessary.

Apart from a native X-ray contrast, so can the

Object to be examined, for example a human, to carry out the method according to the invention, an X-ray contrast agent can be administered. The X-ray contrast agent can be administered, for example, enterally or parenterally, in particular by IV, im. or subcutaneous injection or infusion. Subsequently, the X-ray image is created. Suitable contrast agents which have high attenuation coefficients in the selected spectral range are suitable. Contrast agent whose absorbing element is the K edge of the absorption spectrum in the selected Spectral range is just as well suited. Such X-ray contrast agents contain contrasting chemical elements with an atomic number of 35 or greater than 35 - this is, for example, bromine-containing contrast agents -, with an atomic number of 47 or greater than 47 - this is iodine-containing

Contrast agents - having an atomic number of 57 or greater than 57 - these are lanthanide-containing contrast agents, in particular gadolinium-containing contrast agents - or having an atomic number of 83 - these are bismuth-containing contrast agents -. Therefore, X-ray contrast agents containing contrasting chemical elements having an atomic number of 35 (bromine) to 83 (bismuth) are suitable. Particularly suitable are contrast agents with contrasting chemical elements with an atomic number of 53 (iodine) - 83 (bismuth). Also suitable are X-ray contrast agents with contrasting chemical elements with an atomic number of 57 or greater than 57 (lanthanides) - 83 (bismuth) and more preferably means with contrasting chemical elements with an atomic number of 57-70 (lanthanides: La, Ce, Pr, Nd , Pm, Sm, Eu ₁ Gd, Tb, Dy, Ho, Er, Tm, Yb).

Suitable iodine-containing X-ray contrast agents are, for example

Compounds containing triiodo-aromatics, such as amidotrizoate, lohexol, lopamidol, lopansäure, lopodinsäure, lopromid, lopronsäure, lopydon, lotalaminsäure, lopentol, iversol, loxaglat, lotrolan, iodixanol, lotroxinsäure, loxaglinsäure and loxitalaminic acid and losimenol (INN). Brand name for iodine-containing x-ray contrast agents are Urografin ^® (Schering), Gastrografin ^® (Schering), Biliscopin ^® (Schering), Ultravist ^® (Schering) and Isovist ^® (Schering).

Are also suitable as X-ray contrast agent 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 (Gadobenate, MultiHance), Gd-DTPA-BMA (gadodiamide, Omniscan ^® (Amersham Health)), Dy-DTPA-BMA, Gd-DTPA-polyiyine, Gd-DTPA cascade polymers, among others, where DTPA = diethylenetriaminepentaacetic acid, DOTA = 1, 4,7,10-tetraazacyclododecane, HP-DO3A = 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-13-oic, benic acid), DTPA-BMA = Diethylenetriaminepentaacetate-bis (methylamide), DTPA-polylysine = diethylenetriamine pentaacetate polylysine, DTPA cascade polymers.

The X-ray contrast agents can be administered enterally and parenterally. In parenteral administration, the intravenous (i.v.) application is preferably selected. Preferred dosages in the iodine-containing nonionic contrast media are doses up to 0.75 g I / kg body weight. This corresponds to about 6 mmol I / kg body weight. The dose may further be preferred at 1.5 g I / kg body weight (corresponding to about 12 mmol I / kg

Body weight) and in exceptional cases up to 2 (corresponding to about 16 I) or 5 g I / kg body weight (corresponding to about 39 mmol I / kg body weight) can be increased. For the lanthanide complexes, the preferred dose is 0.1 mmol / kg body weight. Also suitable and furthermore preferred are doses of up to 0.3 mmol / kg body weight or up to 1 mmol / kg body weight.

The emission lines of gadolinium are at 43.0 and 48.7 keV, i. far above the emission lines of iodine, which are at 28.6 and 32.3 keV. The metal complexes may, for example, also contain all other lanthanides, such as lanthanum, dysprosium or ytterbium, instead of the gadolinium atoms.

Digital detectors have been offered by several manufacturers for some time (for example: The BBI Newsletter, February 1999, page 34; HG Chotas, JT Dobbins, CE, Ravin, "Principles of Digital Radiography with Large-Area, Electronical Iy Readable Detectors : A Review of the Basics ", Radiol., 210 (1999) 595-599) .They often consist of amorphous silicon or other semiconductor materials The following detectors are suitable: detectors with phosphor plates (for example from Fuji Chemical

Industries, Konica), with amorphous silicon (for example from GE Medical, Philips Medical, Siemens Medical), with selenium (for example from Philips Medical, Toshiba), with gadolinium hyposulfite (from Kodak, for example), with cadmium telluride (CT) or cadmium-zinc. Telluride (CZT) semiconductors, with yttrium oxyorthosilicate, with lutetium oxyorthosilicate, with sodium iodide or bismuth germanate. Particularly good results are achieved with the so-called C (Z) T detectors, i. Detectors consisting of a cadmium (zinc) telluride (C (Z) T) semiconductor.

The construction of an energy dispersive detector formed from a semiconductor is described in detail in US Pat. No. 5,434,417. In this case, segmented semiconductor strips are provided, which are irradiated from the front side with the X-rays. The radiation penetrates into it

Semiconductor material until it interacts with the semiconductor material. The penetration depth depends on the energy of the X-ray photons. With larger energy of the X-ray photons, the radiation penetrates deeper until it interacts with the detector material and generates a current pulse by a photoelectric effect, as at lower energy

X-ray photons. The current pulses can be derived in the individual segments of the detector by means of attached electrical contacts. The current pulses are processed with a preamplifier.

On the one hand, the detector can be designed in the form of a flatbed detector. In this embodiment, all pixels are detected simultaneously and forwarded to the correlation unit for evaluation. The detector in this case consists of a planar arrangement of individual detector sensors, preferably in a row and columns of such sensors having matrix.

Furthermore, it is also possible to provide a detector unit which serves to determine the emitted X-ray radiation and, if appropriate, to receive a Emission image and this is designed with an X-ray optical module for energy selection.

Instead of the flatbed detector, it is also possible to use line detectors or a matrix of a plurality of detectors suitable for recording a single pixel. In the case of the latter detectors, the X-ray radiation from the examination object is simultaneously conducted via X-ray light guides. A plurality of such optical fibers is combined to form an area detector.

Furthermore, the detector can be designed to receive a single pixel and be movable to accommodate all pixels. In this embodiment, the detector can only detect energy-dependent intensities in a single pixel during the measurement. The intensities of the individual pixels are detected in succession, for example line by line, and forwarded to the correlation unit for further processing.

In addition, the detector can also have an array of detector sensors designed to receive one pixel each and can be moved to accommodate all the pixels. As an array of detector sensors according to the present invention, both a row of detector sensors and another arrangement, for example, a matrix-like arrangement, of detector sensors understood. In this embodiment, the detector detects the intensity values in the individual pixels line by line or optionally also in blocks. To record all intensity values, the detector is preferably moved perpendicular to a main axis of the array during the measurement. The intensity values determined during the measurement are transmitted to the correlation unit.

To image, for example, the distribution of contrasting chemical elements in the examination subject, it is advantageous to detect the radiation intensities emitted by the respective spatial elements, each with the same weighting. Furthermore, it is for this purpose too advantageous to apply the respective Raumeiemente with jeweiis same radiation intensity from the X-ray source. In practice, however, these premises only prove to be approximate since, on the one hand, the irradiated X-radiation is attenuated to varying degrees by absorption depending on the path traveled by the radiation in the examination subject and, on the other hand, the radiation emitted by the spatial elements in the examination subject is still weakened in the object under investigation path to the detector to varying degrees attenuated by the self-absorption.

This problem occurs in all emission spectroscopic methods. In order to solve the problem, the second intensity values are first corrected taking into account the absorption of incident X-radiation and / or the inherent absorption of the emitted X-radiation in the examination subject, and the first and second intensity values are correlated with each other only after this correction, pixel by pixel. Such a correction can be carried out by means of numerical methods, taking into account the geometry of the examination subject and an at least approximate location-dependent radiopacity. To determine the location-dependent radiopacity, the images generated from the first intensity values can be used. To determine the location-dependent absorption and self-absorption, the location-dependent radiopacity obtained from this measurement can be used as a first approximation because the absorption coefficients for the irradiated X-radiation are similar to those of the emitted radiation.

Because of the inherent absorption of the emitted radiation, it may also be advantageous to move the position and angular position of the second detector relative to the examination area (ROI) during the measurement, such as on a circular segment path, to structural inhomogeneities in the examination subject, which vary depending on the viewing angle and location have an absorbent effect, compensate. The images would be obtained in this case after correction of the self-absorption by averaging. The signal originating from the preamplifier is then passed into the at least one correlation unit with which the intensity of the transmitted X-ray radiation from a pixel of the examination object is correlated with the image of the emitted X-radiation (X-ray scattering and X-ray fluorescence) from the same pixel. The correlation unit may be a correspondingly programmed data processing system.

In order to correlate the intensity values of the photons of the two modalities (transmission image and emission image), these are pixel-wise correlated, preferably subtracted from one another or divided by one another.

For this purpose, in one case a comparator and in the other case a divisional element can be used for pixel-by-pixel correlation.

Of course, other mathematical operations for correlating the intensity values of the transmitted and emitted

X-rays are performed by a pixel.

For processing the measured intensity values of a pixel, the following devices are preferably provided which can be realized in a data processing system, namely:

d1. a first memory unit, with which the first intensity values of the transmitted X-ray radiation can be stored pixel by pixel, d2. a second memory unit, with which the second intensity values of the emitted X-ray radiation can be stored pixel by pixel

(eg with the elements I ₁ Gd, Yb), d3. an arithmetic unit which provides for a suitable correlation of the two generated image data sets and thus generates or calculates 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 the two biid points in transmission and emission, with the emission image being matched to the contrast agent used via the characteristic emission lines. A mixture of x-ray contrast agents (eg Ultravist ^® and Gadovist ^®) or both iodine and a lanthanide (such as Gd or Dy) containing substances is used, each characteristic emission lines may be used for the emission imaging, wherein the measured data sets then correlated imagewise together and, alternatively, the respective intensity values are correlated pixel by pixel, and the obtained data are subsequently used for image presentation. For this purpose, the data obtained are transmitted pixel by pixel 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 serve the following figures and examples. In order to provide a direct impression of the operation of the invention, it was refrained in all cases from correcting the measured X-ray spectra according to the absorption of the excitation beam and the self-absorption. They show in detail:

1 shows an illustration of a test arrangement in a computer tomograph;

2 shows a schematic representation of the arrangement for image acquisition or the experimental setup;

Fig. 3: a schematic representation of the experimental arrangement for

Generation of the first phantom measurements; Fig. 4: emission spectra of the phantom of Fig. 3, filled with water (Fig.

4a), Ultravist ^® (Figure 4b), Gadovist ^® (Figure 4c)..; FIG. 5: emission spectra of the phantom of FIG. 3 filled with water (FIG.

.. 5a), Ultravist ^® (Figure 5b), Gadovist ^® (Figure 5c), wherein each a 5 cm thick PMMA disk was placed between the detector and phantom; FIG. 6: Intensity of the emission as a function of the position / displacement of the phantom from FIG. 3 in selected energy bands (corresponding to the K ₀ and K ₂ lines (iodine: FIG. 6 a, gadolinium: FIG. 6 b) Gadolinium: Fig. 6c) Fig. 7: CT sectional images (transmission images) of the phantom filled with Gd, an iodine / Gd mixture, iodine, air and water.

FIG. 1 shows a photographic representation of a test arrangement in a computer tomograph with a rubber ball 1 which is fastened to a stand 2. The rubber ball is located in the center of the computer tomograph. In various experiments, the rubber ball was filled with air, water and different contrast agent solutions. The ball was between the CT tube (above the rubber ball, not shown) and a line detector (below the table visible under the rubber ball, not visible).

At a 90 ° angle to the connecting line between the CT tube, the rubber ball and the detector, a measuring chamber 3 for detecting the X-ray fluorescence was positioned. With this experimental setup, a contrast-filled tissue, tumor or the like was used as

Object of examination simulated, which is examined in the computer tomograph. For this purpose, the object was scanned in layers and thereby measured the scattering spectra.

The experimental setup used for this experiment is shown in detail in FIG. The diagram shown there shows the ball 1, which was located as a phantom in the isocenter of the gantry 4. The CT tube 5 was located at the 12 o'clock position and stood there. The measuring chamber 3, consisting of a detector 6 and a lead tube 7, was directed at the phantom (ball) at an angle of 90 ° to the X-ray cone beam emerging from the CT tube (in the z-direction, see arrow). , For the detection of the X-ray radiation, a CZT detector 6 with a 3 mm × 3 mm × 2 mm cadmium-zinc-telluride crystal and 100/400 μm pinhole diaphragms was used (Amptek Inc., USA). The data recorded by the fluorescence detector were forwarded by the detector via an amplifier 8 to a multi-channel analyzer 9 and then a

Excel ^® (Microsoft) -Graphiktabelle supplied stored on a PC 10th The signal intensities SI = SI (E) were thus available in digital form as a function of the energy E.

In Fig. 3 is a schematic representation of the experimental arrangement for

Generation of the first phantom measurements shown. Part of the measuring chamber for measuring the fluorescence 3 can be seen on the left in the illustration, while the ball 1 is shown in the middle of the illustration. The individual sectional planes running vertically in FIG. 3, from which the fluorescence arrives at the measuring chamber, have been cut through from above

X-ray fan beam generated. The dashed lines mark the respective positions of the CT tube above the image section. The horizontal scale indicates the shift of the fan beam and thus indicates the respectively addressed cutting plane (excited layer) in the ball.

A "zero measurement" was performed at +45 mm and thus outside the excitation beam.

After each acquisition of a spectrum, the entire measurement setup was moved 10 mm further into the gantry (in the z direction) and the new spectrum was recorded. Layer by layer, different spectra were generated depending on the position of the ball in the beam or the ball geometry.

With this measurement setup, it was thus possible to measure the X-ray fluorescence as a function of 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, then Self absorption of Emission minimal and at z = 0 maximum; because of the spherical geometry, an absorption effect in the radiation at higher contrast agent concentrations is noticeable).

Example 1 :

In a first measurement, the ball was filled with water and measured at 80 kV, 50 mA for every 80 s per position of the ball in the beam according to FIG. 3 (parameter: detector: XR-100.CZT (pinhole 0.1 mm), Distance ball - detector: 18.0 cm, distance ball - CT tube: 32.0 cm).

FIG. 4a shows the scattered spectra of the water in the phantom for the different positions.

In a second measurement of the bale parameter was filled with a solution of 50 mmol / l of iodine in water (Ultravist ^®) and at 80 kV, 50 mA for each measured 80 seconds per position (: Detector: XR-100.CZT (pinhole 0 ,1 mm)

The obtained emission spectra at the various positions are shown in Fig. 4b. The K ₀ and Kβ lines of iodine (28.6 and 32.3 keV) are clearly visible. The graph shows a dependence of the measured intensity of the X-ray fluorescence on the geometry of the phantom. The larger the irradiated layer of the phantom, the higher the measured intensity.

In a third measurement the ball parameter was filled with a solution of 50 mmol / l gadolinium in water (Gadovist ^®) and at 80 kV, 50 mA for each measured 80 seconds per position (: Detector: XR-100.CZT (pinhole 0 ,1 mm)).

The obtained emission spectra at the various positions are shown in Fig. 4c. The K ₀ and Kβ lines of gadolinium (43.0 and 48.7 keV) are clearly visible. It showed that the intensity of measured emission radiation, in particular in the field of K-lines on the geometry of the ball in the radiation field is dependent.

Example 2:

In the individual measurements of this experiment, a 5 cm thick PMMA disc was positioned as a filter between the detector and the phantom to simulate the intrinsic absorption of the X-ray fluorescence by surrounding tissue.

FIG. 5a shows the scattered spectra of the water in the phantom for the different positions.

In a second measurement of the bale parameter was filled with a solution of 50 mmol / l of iodine in water (Ultravist ^®) and at 80 kV, 50 mA for each measured 80 seconds per position (: Detector: XR-100.CZT (pinhole 0 ,1 mm)).

The obtained emission spectra at the various positions are shown in Fig. 5b. The intensity of the fluorescence radiation decreased due to the inserted PMMA disk. It was verified that the intensity became the lower the thicker the disk was. However, even at the largest layer of the ball (in the middle), the K-lines remained measurable.

The obtained emission spectra at the various positions are shown in Fig. 5c. Again, the fluorescence radiation decreased by the inserted PMMA disc. Since the K ₀ and Kβ lines of gadolinium are at 43.0 and 48.7 keV, respectively, a significantly more intensive fluorescence radiation was detectable in the presence of the 5 cm thick PMMA disc in the faiie of iodemission as before. Therefore, even in this case, even with the largest layer of the ball (in the middle), the K-lines remained measurable.

Step Example! 3:

In a further experiment, the intensity values of the fluorescence were determined and recorded as a function of the positioning of the ball in relation to the X-ray beam.

In a first measurement the ball is filled with a solution of 50 mmol / l of iodine in water (Ultravist ^®) and at 80 kV, 50 mA, each for 80 seconds per position measured.

In Figure 6a, the intensity of the fluorescence radiation is plotted against the position / displacement of the phantom in selected energy bands corresponding to the Kα line of iodine at 28.6 keV and the Kβ line of iodine at 32.3 keV. The profile of the emission intensity caused by the shape of the ball can be seen from this figure.

In a second measurement the ball is filled with a solution of 50 mmol / l gadolinium in water (Gadovist ^®) and at 80 kV, 50 mA, each for 80 seconds per position measured.

In Fig. 6b, the intensity of the fluorescence radiation is plotted against the position / displacement of the phantom in selected energy bands corresponding to the K _α line of gadolinium at 43.0 keV and the Kβ line of gadolinium at 48.7 keV. The profile of the emission intensity caused by the shape of the ball can also be seen from this figure.

In a third measurement the ball is filled with a solution of 25 mmol / l iodine (Ultravist ^®) and 25 mmol / l gadolinium in water (Gadovist ^®) and at 80 kV, 50 mA for each measured 80 seconds per position.

In Fig. 6c, the intensity of the fluorescence radiation is dependent on the position / displacement of the phantom in selected energy bands corresponding to the Kα line of iodine at 28.7 keV, the Kp line of iodine at 32.3 keV, the Kα line of gadolinium at 43.0 keV and the Kβ line of gadolinium at 48.7 keV. As can be seen from Fig. 6c, the ball profile is reproduced insufficiently in the direct application of the signal intensity as a function of position. This is due to the absorption on the excitation side and the self-absorption on the emission side, which falsify the image. lower

Contrast agent concentrations and corrections of the absorption of the primary beam and intrinsic absorption of the X-ray fluorescence led to a visual representation of the ball in one dimension.

Example 4:

Fig. 7 shows the CT sectional images taken on the foregoing examples of X-ray fluorescence. From top left to bottom right is the gadolinium, the mixture of gadolinium and iodine, the iodine, the pure water and the air-filled ball. The air-filled ball has the least noticeable X-ray attenuation, followed by the ball filled with water. Using the ball with 50 mmol / l contrast element, the X-ray attenuation is more pronounced than with water, a quantitative evaluation is by determining the Hounsfield units (HU) possible, but only the addition of X-ray fluorescence images allows a statement about the element-specific filling of the ball.

Claims

claims:

1. X-ray arrangement for imaging an examination subject containing at least one X-ray contrasting chemical element by means of X-radiation transmitted through the examination subject and emitted by the examination subject

a. at least one substantially polychromatic X-ray emitting X-ray source, b. a first detector or a first detector unit, with or with the values of a first intensity of the by the

Examination object transmitted X-ray radiation can be determined, c. a second detector or a second detector unit, with which or with the values of a second intensity of the X-radiation emitted by the examination object 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 the second intensity values of the emitted X-ray radiation pixel by pixel, and e. at least one output unit for displaying the

Object to be examined from pixel signals obtainable by correlating the first intensity values with the second intensity values.

2. X-ray arrangement according to one of the preceding claims, characterized in that the correlation unit has the following facilities: d1. a first memory unit, with which the first intensity values of the transmitted X-ray radiation can be stored pixel by pixel, d2. a second memory unit with which the second intensity values of the emitted X-ray radiation can be stored pixel by pixel, d3. a computing unit, with which the first intensity values of the transmitted X-ray radiation can be correlated with the second intensity values of the emitted X-ray radiation pixel by pixel.

3. X-ray arrangement according to one of the preceding claims, characterized in that the second intensity values can be determined as a function of the energy of the emitted X-ray radiation in a resolved manner.

4. X-ray arrangement according to one of the preceding claims, characterized in that with the second detector or with the second detector unit of the X-ray radiation emitted by the contrasting chemical element contained in the examination object of other emitted X-radiation via their energy is discriminated.

5. X-ray arrangement according to one of the preceding claims, characterized in that the first intensity values and the second intensity values after a previous correction taking into account the absorption of incident X-radiation and / or the

Self-absorption of the emitted X-ray radiation in the examination subject pixel by pixel are correlated.

6. X-ray arrangement according to one of the preceding claims, characterized in that the first and / or the second detector is a flat-bed detector.

7. X-ray arrangement according to one of claims 1-5, characterized in that the first and / or the second detector is designed to receive a single pixel and movable to accommodate all the pixels.

8. X-ray arrangement according to one of claims 1-5, characterized in that for determining the emitted X-radiation is provided with a trained X-ray optical module for energy selection detector unit.

9. X-ray arrangement according to one of claims 1-5, characterized in that the first and / or the second detector comprises an array of detector housings designed to receive one pixel each and is movable to accommodate all the pixels.

10. X-ray device according to one of the preceding claims in the radiological findings in combination with other radiographic methods of imaging, such as positron emission tomography (PET), single-photon emission computer tomography (SPECT) and

Sonography and methods of optical imaging.

11. Use of the X-ray device according to one of claims 1 - 10 for imaging an examination object containing at least one X-ray contrasting chemical element by means of transmitted through the examination object and emitted by this X-ray.

12. Use of the X-ray device according to claim 11, characterized in that the following method steps are carried out:

a. Irradiation of the examination subject with substantially polychromatic X-radiation, b. Determining values of a first intensity of the X-ray radiation transmitted through the examination object, c. Determining values of a second intensity of the X-radiation emitted by the object under examination, d. Correlating the first intensity values of the transmitted

X-ray pixel by pixel with the second intensity values of the emitted X-radiation, as well as e. Representing the examination subject from pixel signals obtained by correlating the first intensity values with the second intensity values.

13. Use of the X-ray device according to one of claims 11 and 12, characterized in that the second intensity values are measured in a resolved manner as a function of the energy of the emitted X-radiation.

14. Use of the X-ray device according to one of claims 11-13, characterized in that X-ray radiation emitted by the contrasting chemical element contained in the examination object is emitted by other X-radiation emitted via the latter

Energy is discriminated against.

15. Use of the X-ray device according to one of claims 11-14, characterized in that the first intensity values and the second intensity values are available after a previous correction

Considering the absorption of incident X-ray radiation and / or the intrinsic absorption of the emitted X-ray radiation in the examination object pixel by pixel are correlated.

16. Use of the X-ray device according to one of claims 11 - 15, characterized in that a first and a second detector or a first and a second detector unit are provided.

17. Use of the X-ray device according to claim 16, characterized in that the first and / or the second detector is a flatbed detector.

18. Use of the X-ray device according to claim 16, characterized in that the first and / or the second detector is designed to receive a single pixel and is moved to receive all the pixels.

19. Use of the X-ray device according to claim 16, characterized in that the first and / or the second detector has an array of detector housings designed to receive one pixel in each case and is moved to accommodate all the pixels.

20. Use of the X-ray device according to claim 16, characterized in that for determining the emitted X-radiation, a detector unit designed with an X-ray optical module for energy selection is provided.

21. Use of the X-ray device according to any one of claims 11 - 20, characterized in that the contrasting chemical element is selected from a group comprising bromine, iodine, lanthanides and bismuth.

22. Use of the X-ray device according to any one of claims 11 - 21, characterized in that the contrasting chemical element is administered enterally or parenterally.

23. Use of the X-ray device according to one of claims 11 - 22 for element-specific pictorial or quantitative representation of a the examination region containing at least one contrasting chemical element in the examination subject.

24. Use of the X-ray device according to one of claims 11 - 23 in the radiological examination in combination with other radiological methods of imaging such as positron emission tomography (PET), single photon emission computer tomography (SPECT) and sonography and methods the optical imaging.

25. Imaging X-ray contrast method on an examination subject by means of the X-ray radiation transmitted through the examination subject and comprising it, comprising the following method steps:

a. Irradiation of the examination object with substantially polychromatic X-radiation, b. Determining values of a first intensity of the X-ray radiation transmitted through the examination object, c. Determining values of a second intensity of the X-radiation emitted by the object under examination, d. Correlating the first intensity values of the transmitted

26. X-ray contrast method according to claim 25, characterized in that the examination object is first administered at least one X-ray contrasting chemical element before carrying out the method steps a) to e).

27. X-ray contrast method according to one of claims 25 and 26, characterized in that the second intensity values in Depending on the energy of the emitted X-ray radiation can be measured dissolved.

28. X-ray contrast method according to any one of claims 25 - 27, characterized in that the X-ray radiation emitted by the contrasting chemical element contained in the examination subject is discriminated from other emitted X-ray radiation via its energy.

29. X-ray contrast method according to claim 25, characterized in that the first intensity values and the second intensity values are correlated with each other after a previous correction taking into account the absorption of incident X-radiation and / or the intrinsic absorption of the emitted X-radiation in the examination subject pixel by pixel.

30. X-ray contrast method according to any one of claims 25 - 29, characterized in that a first and a second detector or a first and a second detector unit are provided.

31. X-ray contrast method according to claim 30, characterized in that the first and / or the second detector is a flatbed detector.

32. X-ray contrast method according to claim 30, characterized in that the first and / or the second detector is designed to receive a single pixel and is moved to receive all the pixels.

33. X-ray contrast method according to claim 30, characterized in that the first and / or the second detector has an array of detector elements designed to receive one pixel in each case and is moved to accommodate all the pixels.

34. An X-ray method according to claim 30, characterized in that for determining the emitted X-radiation, a trained with a X-ray optical module for energy selection detector unit is provided.

35. X-ray contrast method according to any one of claims 25 - 34, characterized in that the contrasting chemical element is selected from a group comprising bromine, iodine, lanthanides and bismuth.

36. X-ray contrast method according to any one of claims 25 - 35, characterized in that the contrasting chemical element is administered enterally or parenterally.

37. X-ray contrast method according to any one of claims 25 - 36 in the radiological findings in combination with other radiographic methods of imaging such as positron emission tomography (PET), single-photon emission computer tomography (SPECT) and sonography and methods of optical imaging.