WO2004095069A1

WO2004095069A1 - Detector element for combined detection of x-radiation and gamma radiation

Info

Publication number: WO2004095069A1
Application number: PCT/IB2004/050418
Authority: WO
Inventors: Michael Overdick; Herfried Karl Wieczorek
Original assignee: Philips Intellectual Property and Standards GmbH; Koninklijke Philips Electronics NV
Current assignee: Philips Intellectual Property and Standards GmbH; Koninklijke Philips NV
Priority date: 2003-04-24
Filing date: 2004-04-13
Publication date: 2004-11-04
Anticipated expiration: 2005-10-24

Abstract

The invention relates to a detector element of an instrument for combined production of a CT x-ray image and a PET image. The detector element contains a conversion element (1), which produces a charge signal (Q) proportional to the energy of incident x-ray quanta (X) and gamma quanta (y). In an evaluation unit, the charge signal (Q) is converted by an amplifier (30) into a voltage pulse (U), which is subsequently processed separately by a single-pulse analyzer (50) for determining PET information (YPET) and a pulse-sequence analyzer for determining CT information (YCT). By using a common amplifier, the detector element achieves high accuracy even for conversion elements (1) with a weak charge signal.

Description

Detector element for combined detection of x-radiation and gamma radiation

The invention relates to a detector element for combined detection of x- radiation and gamma radiation, which can be used in a device capable of producing x-ray images and PET images. It also relates to a method for producing an x-ray image and a PET image in parallel.

, 5

US 6449 331 Bl discloses a combined CT (Computed Tomography) and PET scanner (PET: Positron Emission Tomography), which has detector elements for sequentially performed detection of x-ray quanta and gamma quanta. The detection of x-ray quanta is

10 necessary for generating x-ray images and the detection of gamma quanta is necessary for detecting the gamma quanta, moving in opposite directions, which result from the annihilation of a positron with an electron, the positron being created by the radioactive decay of a radionuclide introduced into the patient's body. The combination of the detection of x-ray and gamma quanta by the same detector elements has the advantage that the

15 associated apparatus is commensurately smaller and less expensive. The combination also ensures that both types of radiation are detected at the same position. The detector elements described in US 6 449 331 Bl contain a conversion element, which absorbs the quanta to be detected and provides a corresponding charge signal at its output. This charge signal is delivered to a circuit for analyzing individual pulses in which, on the one hand, the energy of

20 an absorbed (gamma) photon can be ascertained using a charge amplifier and, on the other hand, the arrival time of the photon can be ascertained using a fast amplifier. Said information is used in order to carry out a PET coincidence check. The charge signal from the conversion element is sent to an integration circuit, in order to determine an average radiation dose or number of detected quanta. A disadvantage with such detector elements is

25 that the detection of x-ray quanta, on the one hand, and gamma quanta, on the other hand, can only be carried out successively, but not simultaneously. Simultaneous operation of the described processing branches (integration circuit and circuit for analyzing the individual pulses) is not possible because the charge signal from the conversion element would be divided between the two processing branches in a scarcely controllable way. Against this background, it was an object of the present invention to provide means for accurate detection of x-radiation and gamma radiation, with which an x-ray image and a PET image can be produced in parallel.

This object is achieved by a detector element with the features of claim 1, a device with the features of claim 9, and a method with the features of claim 10. Preferred embodiments are contained in the dependent claims.

The detector element according to the invention is used for combined detection of x-radiation and gamma radiation. It can be used, in particular, in a device for (simultaneously or sequentially) producing both x-ray images and PET images, and contains the following components: a) a conversion element for converting x-ray quanta and gamma quanta into an electrical signal dependent on the quantum energy, which is preferably a charge signal; b) an amplifier, coupled to the conversion element, for converting and amplifying the aforementioned electrical signal (for example a charge signal) into an amplifier pulse, the amplitude and time integral of the amplifier pulse having a known relation with the electrical signal. The amplifier pulse may, in particular, be an electrical voltage pulse. Owing to its relation with the electrical signal from the conversion element, the value of the electrical signal can be inferred from both the amplitude and the time integral of the amplifier pulse. c) a single-pulse analyzer, coupled to the output of the amplifier, which is designed for analyzing individual amplifier pulses with respect to their value and the time of their occurrence ("event time"). The single-pulse analyzer therefore provides, in particular, the basis for the coincidence detection of gamma quanta which originate from an annihilation process of a positron and an electron. d) a pulse-sequence analyzer, also coupled to the output of the amplifier, which is designed for determining characteristic features of the sequence of amplifier pulses. These characteristic features may, for example, be a running average of the time integral and/or a running average of the rate of the amplifier pulses, or corresponding values starting from a fixed reference time.

In such a detector element, the electrical signals (referred to below as "charge signals" by way of example and without restriction of generality) delivered by the conversion element are first converted or amplified in a single amplifier to form a pulse. The amplifier pulses are then analyzed individually in a first branch and collectively in a second branch. The central amplification has the advantage that all of the charge signal is recorded, and the subsequent analysis is only carried out with the resulting amplifier pulse. High accuracy of the detector element is achieved in this way, which requires comparatively few components overall for the central amplifier. The one-off amplification of the charge signal downstream of the conversion element has the further advantage that it is possible to use conversion elements that provide only a comparatively weak charge signal. Examples include a photodiode or a conversion element, which contains a directly converting material or consists thereof. A directly converting element is characterized in that it converts absorbed x-ray or gamma quanta directly into a charge signal. With indirectly converting materials, conversely, the absorbed quanta are first converted into visible light, which is then converted into a charge signal by a downstream photodetector.

The dependency between the charge signal produced by the conversion element and the absorbed x-ray or gamma quanta is preferably such that the energy deposited by the quanta can be inferred from the value of the charge signal. Preferably, there is a simple proportionality between the charge signal and the deposited quantum energy.

There is preferably also an essentially proportional relation between the amplitude or time integral of the amplifier pulse, on the one hand, and the value of the associated charge signal, on the other hand. In this case, the value of the charge signal can be inferred in a particularly straightforward way from the amplitude or time integral. The relation is furthermore linear in this case, so that the sum of the amplitudes/time integrals of the amplifier pulses of two different charge signals is equal to the amplitude/time integral of the amplifier pulse of the superimposed charge signal.

As was mentioned above, the single-pulse analyzer is used to detect gamma quanta for a PET image. Two items of information about an absorbed quantum are required for this, namely its quantum energy and the time of its arrival. The latter is needed to check for coincidence with a second gamma quantum created by the same annihilation process. According to a first embodiment, the single-pulse analyzer contains two parallel branches, namely: a) an amplitude-determination branch, the input of which receives the amplifier pulse, and which is designed to provide a signal at its output which is proportional to the amplitude or time integral of the amplifier pulse. b) a timestamp branch, the input of which receives the amplifier pulse, and which is designed to indicate the event time of the amplifier pulse at its output. The event time is in this case preferably defined by a unique characteristic position in the amplifier pulse, for example by particular points on its leading edge such as the 50% amplitude value.

According to a second embodiment of the single-pulse analyzer, it has a discriminator as the input stage, which receives the amplifier pulses at its input and which is designed to provide further processing stages only with amplifier pulses above a predetermined amplitude threshold value. Since the gamma quanta on which a PET image is based have a higher energy (about 511 keV) than typical x-ray quanta (about 30 - 140 keV), the discriminator can be used to distinguish "normal" x-radiation from gamma quanta which potentially originate from an annihilation process. The subsequent processing can then be restricted to the relevant gamma quanta.

The pulse-sequence analyzer is used to determine time-averaged properties of the sequence of amplifier pulses. Such features are of prime importance for the production of an x-ray image, which is typically based on the detection of a comparatively large number of x-ray quanta. According to a preferred configuration, the pulse-sequence analyzer has a counter circuit, the output of which indicates the number of amplifier pulses detected since a reference time, and/or the current rate of the amplifier pulses. In addition or as an alternative, the pulse-sequence analyzer may have a dose circuit, which receives the amplifier pulse at its input and which is designed to indicate the integral of the amplifier pulses since a predetermined reference time at its output. The reference time may in this case vary with time, so that the integral effectively indicates a running average of the time integral over the amplifier pulses. From the integral, it is possible to infer the value of the charge signals at the output of the conversion element, and therefore the energy deposited by the x-ray and gamma quanta.

The invention also relates to an instrument for producing both x-ray images and PET images, which contains the following components: a) an x-ray source for emitting x-radiation. b) a detector for detecting x-radiation and gamma radiation, which consists of detector elements of the type explained above. This means that each of the detector elements has a conversion element, an amplifier connected thereto, as well as a single-pulse analyzer and a pulse-sequence analyzer arranged in parallel downstream of the amplifier, c) a data-processing unit, which is coupled to the detector and is designed to reconstruct an x-ray image and a PET image from signals provided by the detector elements.

The invention furthermore relates to a method for producing an x-ray image and a PET image in parallel, which contains the following steps: a) the x-ray quanta and the gamma quanta are converted into an electrical signal, preferably a charge signal. b) the electrical signal (charge signal) is converted and amplified to form an amplifier pulse, the amplitude and time integral of the amplifier pulse being related to the electrical signal so that the value of the electrical signal can be inferred from their values. c) the values and the times of individual amplifier pulses are determined, so that it is possible to detect coincident gamma quanta from an annihilation process. d) for the chronological sequence of amplifier pulses, their rate and/or their time integral are determined so as to provide the information necessary for producing an x-ray image.

It is important for the method that the charge signal is converted only once, centrally, into an amplifier pulse which is then available as a value accurately and unequivocally related to the charge signal for the subsequent analysis of the individual pulses and the pulse sequence. These and other aspects of the invention are apparent from and will be elucidated with reference to the embodiments described hereinafter.

In the drawings, in which the same reference numbers stand for the same components:

Fig. 1 shows the general structure of a detector element according to the invention for x-radiation and gamma radiation;

Fig. 2 shows a first possible embodiment of the amplifier; Fig. 3 shows a second possible embodiment of the amplifier; Fig. 4 shows the detailed structure of a special embodiment of the detector element in Fig. 1 ;

Fig. 5 shows a variant of the PET channel in a detector element according to Fig. 1;

Figs 6 - 9 show various possible arrangements of a detector and an x-radiation source in a device for combined production of CT and PET images.

CT images and PET images are commonly recorded in medical diagnosis, it being very helpful to have both types of image of the same body part available in parallel for various diagnostic tasks. While CT images provide morphological information about the body part, information about functional aspects can be derived from PET images. The two types of image should ideally be produced at exactly the same position in the patient, in order to ensure diagnostic interpretability. In this context, machines in which CT and PET systems are combined in a

"gantry", in order to be able to take CT and PET images of the same patient in close succession, are known from the prior art. Disadvantages with this, however, are the great outlay involved in providing two independent systems and the long length of the imaging tunnel which is necessary, and which the patient often finds uncomfortable. For this reason, combined devices have been proposed in which the same detector elements undertake the detection of x-radiation for the CT system and gamma radiation for the PET system. The different requirements of the respective systems, however, are problematic for such combinations. For instance, CT systems are essentially characterized by the following features: - a geometry in which an x-radiation source is provided, with a detector extending over about 50° facing it, the two components rotating about the patient axis; an x-ray quantum energy of typically 30 - 140 keV; very high quantum fluxes of typically 3 x 10⁹ photons/s mm²; large dynamic range with very good linearity. Conversely, PET systems are characterized by: a geometry with an immobile 360° detector;

a gamma quantum energy of 511 keV; very low quantum fluxes of typically less than 10³ photons/s mm²; - single-pulse detection required with very good energy resolution (8 - 20%

FWHM) and good time resolution (5 ns); coincidence detection of gamma quanta.

In order to be able to satisfy these different requirements, the invention proposes a detector element for the combined detection of x-radiation for x-ray images and gamma radiation for PET images, which has the structure outlined in figure 1. The first stage of this detector element is a conversion element 1, which converts incident x-ray quanta X or gamma quanta γ into electrical signals. The electrical signals are preferably charge signals Q, the charge being proportional to the energy deposited by an absorbed quantum. The conversion element 1 is preferably configured so that it has an absorption probability of at least 95% in the energy range of the x-ray quanta (30 - 140 keV) and an absorption probability of at least 80% in the range of the PET gamma quanta (511 keV). This absorption property is achieved, for example, by a 2.0 cm thick GOS scintillator (GOS: gadolinium oxysulfide Gd₂θ₂S). In the known way, the conversion element 1 may be fonned in two parts by a scintillator/photodetector combination, or in one part by a directly converting sensor.

According to figure 1, the charge signal Q produced by the conversion element 1 is delivered to an evaluation unit 2 electrically connected to the conversion element 1. The evaluation unit 2 is generally characterized in that it processes the electrical signals from the conversion element 1 and provides the information necessary for the CT operation and PET operation at its output (or its outputs). For CT operation, this involves a signal Y_CT (analog, digital or mixed) which corresponds to the radiation dose detected by the conversion element during a time interval. This may involve the number of quanta detected and/or the energy deposited in the conversion element 1 by the quanta. For PET operation, the evaluation unit 2 delivers information about the energy Y_PET deposited for each detected gamma quantum, as well as information either implicitly or explicitly about the time of the detection, which is necessary for the coincidence detennination.

According to the structure shown in figure 1 for the evaluation unit 2, the charge signal Q from the conversion element 1 is first sent to an (input) amplifier 3 which, in a way which will be explained in more detail, produces an amplifier pulse U therefrom and provides it via its output to a coupling element 4. The coupling element delivers the amplifier pulse, on the one hand, to a "PET channel" with a single-pulse analyzer 50, the output of which provides the infonnation Y_PET necessary for the PET images and, on the other hand, to a "CT channel" with a pulse-sequence analyzer 60, which provides the information Y_CT necessary for the CT images at its output. In the simplest case, the coupling element 4 may be permanent electrical coupling between the amplifier 30 and the single-pulse analyzer 50 or the pulse-sequence analyzer 60, respectively. As an alternative, the element 4 may also be a selector switch by which the amplifier signal U can be appropriately directed to the PET channel or the CT channel, according to the operating mode (CT or PET). The amplifier 30 is characterized by the following properties:

(El) a well-defined voltage pulse U is produced at the output of the amplifier for each (fast) charge signal Q at the input of the amplifier 30.

(E2) the amplitude of the voltage pulse U is proportional to the charge signal Q. (E3) the time integral of the voltage pulse U is likewise proportional to the charge signal Q.

There are a number of viable options for producing an amplifier 30 with such properties. In this context, figure 2 shows a first variant with a so-called CR-RC "pulse shaper". In the example which is represented, this consists of a series circuit containing: a first operational amplifier 31 , the input and output of which are coupled through a capacitor 32; a second capacitor 33 downstream of the output of the operational amplifier; a second operational amplifier 35, the connection between its input and the second capacitor 33 being routed via a grounded resistor 34; a resistor 36, the output of which is grounded through a capacitor 37.

The output signal of such an amplifier has the form:

U (t) = A_! • Q • (t Iτ) ■ exp (-t /τ), where A] is the amplitude constant and τ is the time constant of the amplifier/pulse shaper. Other variants of a CR-RC pulse shaper are known from the literature (W.R. Leo, Techniques for Nuclear and Particle Physics Experiments, 2nd ed., Springer, Berlin, 1994, p. 281).

Figure 3 shows another possible embodiment of the amplifier 30. It is formed by an operational amplifier 31, the input and output of which are coupled in parallel, on the one hand through a capacitor 32 and, on the other hand, through a resistor 38. For a (fast) charge signal Q at the input, this circuit produces a profile of the output voltage U according to:

U(t) = A₂ - Q - exp (-t/τ).

Figure 4 shows possible embodiments of the single-pulse analyzer 50 and the pulse-sequence analyzer 60 of figure 1 in more detail. The coupling element 4 of figure 1 is a simple electrical connection in this case.

In the example which is represented, the pulse-sequence analyzer 60 (CT channel) is constructed in tandem, the information from one of these two channels being in principle sufficient to produce an x-ray image.

The pulse-sequence analyzer 60 has, on the one hand, an integrating channel consisting of an integrator 63 and a sample-and-hold circuit 64. Owing to property (E3) of the amplifier 30, its output signals U can be readily integrated in order to obtain the total charge in a particular time window (cf. also DE 199 45 757 Al). A signal E_tot which indicates the total deposited energy is therefore provided at the output of the sample-and-hold circuit 64. The CT channel 60 also has a channel that counts the voltage pulses U. It consists of a CT event discriminator 61 which, owing to property (E2) of the amplifier 30, may be a simple voltage discriminator (W.R. Leo, op. cit, p. 286) which produces a digital output pulse for events above a voltage threshold. This output pulse can then be counted by a counter 62 downstream. The number n of events since a predetermined time can therefore be obtained from the output of the counter. As an alternative, this number may also noπnalized with respect to time, in order to provide a time-averaged rate.

In the embodiment according to figure 4, the PET channel 50 firstly consists of a PET event discriminator 51, which only processes events above a predetermined energy. For example, it may be set to an energy above about 150 keV or above about 450 keV. The energy infoπnation is recorded on the output side by a sample-and-hold element 52, and is provided as a signal of the single-pulse energy Ep at the output of the evaluation unit 2. In parallel with this, the time Tp at which the voltage pulse U occurred is recorded and signaled out by a timestamp unit 53, connected to the discriminator 51 on the output side. For each potential PET event, its energy and information about the time of this event are therefore available at the output of the PET channel.

Figure 5 shows an alternative embodiment of the single-pulse analyzer or PET channel 50. As explained above, the PET channel 50 must deliver the energy (corresponding to the quantity of charge Q) of an event, as well as a timestamp that is as accurate as possible. Owing to property (E2), the energy can be obtained from the amplitude information of the amplifier pulse U. As shown in figure 5, the series circuit containing a pulse stretcher 54 (cf. W.R. Leo, op. cit., p. 284) and a sample-and-hold element 52 may be used for this. Optionally (in particular together with the embodiment of the amplifier 30 according to figure 3), a differentiating element may also be interconnected at the input of the pulse stretcher 54, so that the correct amplitude information is actually measured when there are closely consecutive signals. A signal representing the energy Ep of an individual event is therefore available at the output of the sample-and-hold element 52.

A timestamp Tp is produced according to figure 5 using a leading-edge discriminator 55 with a downstream timestamp generator 56. This and other methods for producing time information, for example a constant-fraction discriminator (CFD), are described in the literature (W.R. Leo, op. cit., pp. 326 ffi). A simplified time-recording method consists in first recording time information with the leading-edge discriminator 55, and then correcting it with the aid of the amplitude information (so-called "Amplitude based Time Walk Correction", cf. G.H. Sanders et al., A high performance timing discriminator, Nuclear Instruments and Methods, 180 (1981), pp. 603 - 614). Such a correction is advantageously carried out in downstream logic (not shown) or in software.

The PET and CT channels of the detector elements are preferably contained in integrated circuits (for example CMOS) inside a pixel area. A bump-bond technology, for example, may be used for connecting the conversion elements 1 to the evaluation units 2.

The detector element described above with reference to figures 1 to 5 is intended to be used in a device which allows CT x-ray images and PET images to be produced in parallel. In this context, figures 6 to 9 represent schematic longitudinal sections of various alternative arrangements of a detector 100, which is formed by the detector elements, for example arranged in rows and columns, and an x-radiation source (x-ray tube) 110.

In the embodiment represented in figure 6, the detector 100 is fonned as a static cylinder, or tunnel. The x-ray tube 110 is arranged axially alongside the detector 100, and irradiates it at least partially. During imaging, the x-ray tube 100 rotates about the patient (not shown) who is in the middle of the detector tunnel 100.

In the embodiment represented in figure 7, the x-ray tube 110 is arranged inside the detector 100. It again rotates about the patient during imaging. As an alternative, according to Fig. 7, it is also possible for only the target of an electron beam to rotate.

In the embodiment of figure 8, the x-ray tube 110 is rigidly connected to the cylindrical detector 100, and it can irradiate the opposite inner surface of the detector through a gap in the detector wall. The x-ray tube 110 and the detector rotate together about the patient during imaging.

Figure 9 shows another variant, in which the inner part of the device comprises the rotating x-ray tube 110 and a co-rotating detector 100' opposite it, the outer parts of the device being formed by static cylindrical detector sections 100.

Besides the elements shown in figures 6 to 9, it is also possible to use a radioactive preparation as the radiation source for so-called "attenuation correction". This is already done in modern PET systems, with the radioactive source likewise rotating about the patient. Distinction can be made between three different operating modes for the combined CT/PET machine:

CT operation; attenuation recording with a radioactive source;

PET operation. These three operating modes are well known from established CT and PET machines. In the device according to the invention, the CT channel 60 of the detector elements is used for the CT operation, and the PET channel 50 is used for the PET operation. Either the CT channel or the PET channel, or both simultaneously, may be used for the attenuation recording with a radioactive source. Optionally, the attenuation recording with a radioactive source may be omitted entirely, if the attenuation for the PET operation can also be determined reliably from the CT data.

It is furthermore possible for the CT operation and the PET operation to take place in parallel, in which case the PET channel 50 and the CT channel 60 are used simultaneously. Owing to the different quantum energies, the PET channel is not affected by the x-ray quanta of the CT imaging, while, on the other hand, the infrequent gamma rays from PET events also cannot compromise the imaging in the CT channel. When the single- pulse signals are being used in CT operation, the beam hardening in the patient is simultaneously measured by the CT imaging. The detector elements may also be switched over between the single-pulse analyzer 50 and the pulse-sequence analyzer 60 automatically inside the evaluation units 2. This may be controlled by information about the signal level, for example from immediately preceding measurements or from the count rate in the single-pulse analyzer 50.

In order to permit CT operation of the combined device, the area of the detector elements (pixel size) should be approximately lxl mm². While this is a typical pixel size for computer tomography, very much larger pixels are normally used for PET. Owing to Compton scattering during PET operation of the combined device with pixels that are so small, the energy (511 keV) of a PET gamma quantum may not remain in one pixel but instead be distributed over two pixels lying not far apart. By suitable configuration of the PET channel 50 (for example as in figure 4, with an energy threshold of about 150 keV in the PET event discriminator 51), in such a case the two pixels lying not far apart would each register an event with the same timestamp. This could be ascertained using an external coincidence unit, and used to efficiently take such cases into account as well. When designing a detector (figures 6 to 9), it is furtliermore possible to select smaller pixel sizes in the combined CT PET regions than in the pure PET regions. Another option is for a plurality of pixels to be simultaneously connected to a readout channel in PET operation (binning), so as to obtain effectively larger pixels.

Claims

CLAIMS:

1. A detector element for detecting x-ray quanta (X) and gamma quanta (γ), containing: a) a conversion element (1) for converting x-ray quanta (X) and gamma quanta

(y) into an electrical signal (Q), which is preferably a charge signal; b) an amplifier (30), coupled to the conversion element (1), for converting the electrical signal (Q) into an amplifier pulse (U), the amplitude and time integral of which are related to the electrical signal (Q); c) a single-pulse analyzer (50), coupled to the output of the amplifier (30), for analyzing individual amplifier pulses (U) according to value and event time; d) a pulse-sequence analyzer (60), coupled to the output of the amplifier (30), for deteπriining characteristic features of the sequence of amplifier pulses.

2. A detector element as claimed in claim 1, characterized in that the conversion element (1) is a photodiode or contains a directly converting material.

3. A detector element as claimed in claim 1, characterized in that the electrical signal (Q) is essentially proportional to the energy which the x-ray quanta (X) or gamma quanta (γ) have deposited in the conversion element (1).

4. A detector element as claimed in claim 1, characterized in that the amplitude and the time integral of an amplifier pulse (U) are essentially proportional to the value of the associated electrical signal (Q).

5. A detector element as claimed in claim 1, characterized in that the amplifier pulse (U) is a voltage pulse.

6. A detector element as claimed in claim 1, characterized in that the single-pulse analyzer (50) contains the following branches: a) an amplitude-determination branch (52, 54), the output of which provides a signal (E_p) proportional to the amplitude or time integral of the amplifier pulse (U); b) a timestamp branch (55, 56), the output of which indicates the event time (T_p) of the amplifier pulse (U).

7. A detector element as claimed in claim 1, characterized in that the single-pulse analyzer (50) has a discriminator (51) as the input stage, which only delivers amplifier pulses (U) above an amplitude threshold value to further processing stages (52, 53).

8. A detector element as claimed in claim 1, characterized in that the pulse- sequence analyzer (60) has a counter circuit (61, 62), the output of which indicates the number (n) and/or the rate of the amplifier pulses (U), and/or a dose circuit (62, 63), the output of which indicates the integral (E_ot) of the amplifier pulses (U).

9. A device for producing x-ray images and PET images, containing: a) an x-ray source (110); b) a detector (100, 100') for detecting x-radiation and gamma radiation, which consists of detector elements as claimed in one of claims 1 to 8; c) a data-processing unit, which is coupled to the detector and is designed to reconstruct an x-ray image and a PET image from signals provided by the detector elements.

10. A method for producing an x-ray image and a PET image in parallel, wherein a) x-ray quanta (X) and gamma quanta (γ) are converted into an electrical signal (Q), preferably a charge signal; b) the electrical signal (Q) is amplified to form an amplifier pulse (U), the amplitude and time integral of which are related to the electrical signal (Q); c) the values and the times of individual amplifier pulses (U) are determined; d) the rate and/or time integral of the chronological sequence of amplifier pulses (U) are determined.