ES2717179T3 - Device to detect high energy photons - Google Patents

Description

DESCRIPTION

Device to detect high energy photons

The present invention relates to a module for detecting high energy photons. More specifically, the invention relates to a module comprising pixelized solid state detectors at room temperature that detect gamma ray photons, generated by radioactive material within the body of a patient. The invention also relates to systems for detecting high energy photons and to detectors comprising such systems.

The invention finds particular use in medical diagnostic imaging apparatuses.

BACKGROUND

Positron emission tomography (PET) is a diagnostic technique that produces images that show the metabolism and function of tissues and organs (for example, the central nervous system).

Like other diagnostic techniques in Nuclear Medicine, PET is based on detecting and analyzing the distribution within the body of radioisotopes that have been previously administered to a patient. Radioisotopes can be taken orally, can be inhaled as gas or can be given through an injection.

Several positron-emitting radioisotopes are known for medical use. The most commonly used is fluorine-18, which is capable of binding to a glucose marker to obtain 18-fluoro-deoxy-glucose (18F-FDG). In this way, glucose is obtained which is detectable by the emission of a radioactive signal.

After the administration of the radioisotopes, the radioisotopes extend throughout the area of the body to be examined and tend to be picked up, for example, by cancer cells. When the radioisotope decomposes, it emits a positron that, after a few millimeters, is annihilated with an electron. This produces a pair of gamma ray photons that move in the opposite direction, each photon presenting an energy of 511 keV. This pair of gamma ray photons can be detected using the so-called PET scanner. Using the detection position of both gamma-ray photons, the Response Line (LOR) (which is the line connecting the two detection positions of the gamma photons) can be reconstructed. This procedure is illustrated schematically in Figure 14.

Figure 14 shows a conventional PET scanner 1, in which a platform 3 is disposed. On this platform a body 2 of a human or an animal has been schematically indicated. Around the circumference of the PET scanner, a plurality of detectors 4 are arranged. The detector 4a and the detector 4b respectively detect the gamma ray photons moving in the opposite direction. Using this detection, the LOR can be reconstructed.

After collecting several of these events, the points where multiple LORs intersect can be determined. These points indicate a concentration of the radioisotope and, therefore, the possible presence of cancer cells. The PET scanner is connected to a computer, which is responsible for measuring the amount of radioisotopes absorbed by the body and determining the LOR. In this way, it is possible to obtain images that provide details of the structure and function of internal organs and other parts of the body.

In a typical PET protocol, the patient is injected with between 300 and 500 MBq of 18F-FDG (Fluorodeoxyglucose). After allowing it to pass between one and one and a half hours, the patient is placed on the scanner for exploration. A typical PET scan in a conventional PET scanner requires about 30 minutes of scanning time.

PET plays an important role in the diagnosis of tumors. Its accuracy exceeds that of conventional diagnostic imaging systems, as can be seen in the following table (from Journal of Nuclear Medicine Supplement, volume 42, number 5, May 2001 and UCLA):

Diagnostic precision

Type of cancer Representation Conventional PET

Chest 67% 89%

Colorectal 80% 94%

Gastroesophageal 68% 83%

Head and neck 65% 87%

(keep going)

Type of cancer Representation Conventional PET

Liver 81% 93%

Lung 68% 82%

Lymphoma 64% 88%

Melanoma 80% 91%

Pancreatic 65% 81%

Testicular 68% 92%

Uterine / cervical 43% 87%

A PET scanner comprises a plurality of detectors. Today, the best current detector for PET is based on lSO crystals (lutetium oxorthosilicate) with a typical size of 4 mm x 4 mm x 10 mm. The crystals emit flashes of light when they are reached by gamma photons. These flashes of light can be detected using a photomultiplier tube (PMT) that is connected to the crystal. This has also been indicated schematically in Figure 14. The detector 4c comprises a segmented glass 5 and a plurality of PMT's 6. For the person skilled in the art, it will be clear that the plurality of PMT's 6 can also be replaced by a single sensitive photomultiplier. the position (PSPMT).

The light output of LSO crystals for 511keV gamma rays is approximately 4000 phe. The total width at half the maximum (FWHM) that can be obtained with LSO crystals at 511 kV is around 10%. This limited resolution of energy will reduce the ability to eliminate scattered events, which are a type of noise in the reconstructed image. The typical length of the LSO crystal (in the radial direction, when used in a PET scanner) is approximately 10 mm, which implies that the intrinsic uncertainty of the detector of the point of incidence in the radial direction is approximately 3 mm, which will cause an error in the projection of the response line. This can easily be understood by looking at Figure 14. The PMT that records an event essentially provides a two-dimensional coordinate. The radial position where the gamma photon hits the crystal is lost. This loss of information results in a parallax effect, which can lead to an error in the LOR projection. This error naturally deteriorates the quality of the reconstructed image.

Another drawback is the shape of the crystals of the detector, which is a rectangular parallelepiped and when it has a cylindrical shape from said components (for example, for a PET scanner), it is inevitable that it presents cracks at the contact points of the crystals.

Another example of a medical imaging device is a gamma camera. Gamma cameras are also widely used in nuclear medicine. It consists of a single detector plane, formed by a plurality of bright crystals, and a collimator in front of it. Only photons in a small angular range through the holes of the collimator strike the detector; the others are absorbed by the collimator. Therefore, the plane of the detector records a two-dimensional projection of the source distribution.

Single photon emission computed tomography (SPECT) is a nuclear medicine tomographic imaging technique that uses a gamma camera. SPECT representation is done using the gamma camera to acquire multiple two-dimensional images at different angles. Then a computer is used to apply a tomographic reconstruction algorithm to the multiple projections, which produces a three-dimensional image.

Compton cameras are another example of medical diagnostic imaging devices. Compton cameras are used to reconstruct distributions of radioisotopes that emit gamma rays. The range of applications is wide: in addition to the diagnosis in nuclear medicine, they can also be used to control the dismantling of nuclear plants and also for national security (for example, to scan transport containers to detect radioactive material).

A Compton camera has two detection planes. The photons emitted from the source are scattered in the foreground (Compton dispersion) and absorbed in the second plane (photoelectric effect). In both planes the position of the interaction and the deposited energy are measured. The foreground is usually made in a semiconductor material and the second plane in bright crystals. The detectors work in coincidence, so that only photons are recorded that interact with both detectors and deposit a total energy within a certain window. Using the detection position and the energy of the photon, the point of origin of the photon can be calculated, using the so-called Compton formula.

In summary, several techniques and devices (PET, SPECT, gamma camera, Compton camera) that conventionally use detectors based on bright crystals are known in nuclear medical representations.

Detectors based on bright crystals have several drawbacks: in PET scanners, there are cracks inherently present at the contact points of the adjacent parallelepiped crystals. Such as described above for PET scanners, a parallax effect can occur. The crystals used in PET scanners, gamma cameras or Compton cameras have a size of 4 mm x 4 mm x 10 mm, and this determines the intrinsic error (and intrinsic spatial resolution) they present. It is not possible to see anything smaller than 4-5 mm in the case of scanning ^p E ^t . For SPECT, this is of the order of 15-20 mm. The spatial resolution will even be considerably worse, since the detectors do not work with an efficiency of 100%.

Another problem with bright crystals is that the quantum efficiency of detection (DQE) of the bright crystals is quite poor. To improve this DQE, one would have to increase the length of the crystal to increase the probability of capturing the gamma photons. However, adding more material would not necessarily improve the quality of the signal, since a part of the light produced in the crystals in a position that is relatively far from the photomultiplier would be attenuated before reaching the PMT. A gamma photon captured by the crystal near the PMT will produce more photons than a gamma photon captured further away from the PMT. Therefore, the resolution of energy worsens at the expense of detecting more gamma photons. Therefore, the standard length of the bright crystals (10 mm) used in gamma ray detectors is a compromise to obtain a reasonable DQE and, at the same time, have an acceptable energy resolution.

To overcome the drawbacks of the above-mentioned bright crystals, it has been suggested to use pixelized solid state detectors at room temperature in PET detectors. With the pixelated solid state detectors a high spatial resolution can be achieved due to the fact that the detector can be segmented to submillimeter pixels (or voxels).

One of the problems of using solid-state detectors in a PET scanner is that thick detectors are needed to achieve high gamma-ray absorption. In particular, a thickness of 4 cm is needed (if CdTe is used) to capture 90% of the photons with an energy of 511 kV. In the literature, it has been suggested to use very large sensors (for example, 10 mm x 10 mm x 10 mm) of Cd (Zn) Te in which a pixel reading chip is connected to the back of the crystal. This solution seems easy on paper, but in reality it is not. In the first place, the cost of CdTe detectors, which are so large and of good quality, is very high. Secondly, the time collection of the signal will be very long and, therefore, it will not be possible to use it in PET as an activator. Third, with such a thick detector (10 mm), energy spectroscopy deteriorates significantly due to the capture and lifetime of the electron hole.

US 2007/0057191 describes a radiological imaging system comprising a first and a second imaging apparatus. Said first image forming apparatus comprises a series of semiconductor radiation detectors (not pixilated). The set of detectors is connected through cables to ASICs.

DESCRIPTION OF THE INVENTION

An object of the present invention is a module for detecting high energy photons, which at least partially overcomes the drawbacks of the prior art detectors.

Another objective of the present invention is a module for detecting high energy photons with high spatial resolution and high detection efficiency. The high spatial resolution provides the exact position (x, y, z) of the incidence point of the gamma photon; High detection efficiency helps reduce radiation doses.

Another objective of the present invention is a module for detecting high energy photons with a better energy resolution to eliminate a large fraction of the scattered events and increase the signal to noise ratio (SNR). This results in a high contrast image.

To achieve the above objectives, the invention features a module according to claim 1: a module for detecting high energy photons comprising a plurality of devices for detecting high energy photons, each of which comprises one or more high-energy detectors. pixelated solid state to detect high energy photons; means for providing a high voltage for biasing said solid-state detectors; at least one reading element connected to said pixelated solid-state detectors; an input / output element connected to said at least one reading element for data input and output; and a base layer for mounting the pixelated solid-state detectors, the reading elements and the input / output element. Each pixelated solid state detector is mounted on a reading element. The devices are arranged one above the other and connected to each other to support each other and so that the combined pixelated solid state detectors of the devices can provide the three-dimensional coordinates of high energy photon incident points.

Instead of using a brilliant crystal (with its drawbacks mentioned), the invention features a pixelated solid state detector for detecting high energy photons. A typical pixel size can be 1 mm x 1 mm, but it is possible to perform pixelized solid state detectors with pixel sizes of, for example, 10 | jm x 10 | jm. The invention thus offers a tremendous potential for spatial precision. This precision can never be achieved by using bright crystals. Therefore, a device can be adapted to the specific needs in different applications. It is even possible to vary the pixel size within a single solid state detector. Also, there are no dead areas between the pixels of the detectors. Even two pixels together can still detect an event that occurs between two pixels.

In addition, due to its modular and simple structure, a plurality of devices can be easily combined to obtain the appropriate size in three dimensions.

Each device essentially forms a two-dimensional detector. Placing at least one device on top of another, a three-dimensional detector is created. Because the detector is three-dimensional, a parallax effect (as explained in PET scanners) can be avoided. Due to the modular structure of the device, as many devices as desired can be combined to build a detector of greater thickness and, therefore, higher DQE, but without compromising energy resolution or spatial resolution.

The typical thickness of the solid-state detectors used in the devices can be 2 mm (although it is clear that other thicknesses are possible). It is possible to combine a plurality of these devices to an appropriate thickness without encountering problems related, for example, with time collection of conventional thick solid state detectors.

In addition, the invention features a detector that can be used in strong magnetic fields. Therefore, a detector based on the invention can be combined, for example, with an MRI scanner in a single scanner, which allows performing a PET scan and an MRI scan simultaneously.

In some embodiments, the base layers of the devices have an isosceles trapezoidal shape. By giving the base layers of the devices an isosceles trapezoidal shape and adapting the size and shape of the reading element or elements and the pixelated solid state detectors mounted in this trapezoidal shape, a device is obtained which is specially adapted to fit in a circular ring. This is particularly useful for applications in PET scanners.

Optionally, both the pixelated solid state detectors and the base have an isosceles trapezoidal shape. The detectors may also have a different shape, provided they are adapted to the shape of the base layer. It would be possible, for example, a trapezoidal shape using two separate detectors that have quadrilateral shapes. However, a particularly interesting option is to give the solid-state detectors a corresponding isosceles trapezoidal shape, so that they can fit exactly in the base layers. However, the reading elements usually retain their rectangular shape. In some embodiments according to the invention, the device further comprises an intermediate layer mounted between the solid state detector and the reading element to accommodate any displacement between the pixel pads of the detectors and the corresponding pixel channels in the reading elements.

In the case that the pixelated solid state detector has an isosceles trapezoidal shape, the pixel size of the detector varies along the height of the trapezoid to accommodate the change in width (with a constant number of pixels throughout the height of the trapeze).

In some embodiments, the base layers of the devices have a rectangular shape. This shape is especially useful for applications such as gamma cameras and Compton cameras.

Preferably, in the module according to the invention, the input / output elements of the plurality of devices are connected to a single interface. A particularly advantageous way to build a module is by connecting the signal input / output elements of the plurality of devices to a single interface element, which may be, for example, a printed circuit board (PCB). The data collected by the plurality of devices can, therefore, be collected and transmitted using the unique interface.

Optionally, in the module according to the invention, at least part of the solid-state detectors of the devices are arranged in a staggered manner with each other. By arranging the solid state detectors in a stepped manner, a more rigid structure is obtained, since the detectors of one device provide support to the other devices.

Optionally, the module comprises devices in which the pixelated solid state detector, the reading element and the input / output signal element are housed in an upper face of the base layer and in the that the lower face of the base layer comprises means for providing a high voltage for biasing the pixelated solid state detectors in an adjacent device. In the art, various ways of providing high voltage to polarize a solid state detector are known, for example, through suitable wiring. Taking advantage of the fact that a module is formed from a plurality of basic devices, a device can provide the means for biasing the pixelated solid state detector of an adjacent device by arranging them on the underside of the device. Only one device on the "edge" of a module (probably the first) may need other means, for example, an additional kapton layer.

Optionally, the module comprises devices in which the solid state detectors are mounted on the reading elements. Each pixel pad of a solid-state detector may be connected to a channel of input pixels in the reading element (no wiring is needed). In other embodiments, the reading elements and the solid state detector may be arranged side by side (instead of one on top of the other).

In some embodiments, the devices further comprise an intermediate layer mounted between the solid state detectors and the reading elements to accommodate any displacement between the pixel pads of the solid state detectors and the pixel channels in the elements of the solid state detectors. reading. In embodiments where the size of the solid state detector varies along a given direction (eg, in devices where the base layer has an isosceles trapezoidal shape, which are specially adapted to be used to form a circular ring, the shape of the solid state detector varies in a radial direction), the pixel size also varies along this direction. For this reason, an intermediate layer can serve as an interface for adjusting any displacement between the pads in the detector and input channels in the reading element. The intermediate layer can be a kapton layer.

In another aspect, the invention features a system for detecting high energy photons according to claim 10, that is, comprising a plurality of modules, wherein the plurality of module interfaces are connected through suitable connectors to a single interface bus. The single interface bus can collect the data of all the events that occur in the solid state detectors through the data collected by the read elements. It can also provide the necessary operating voltages. Optionally, a system comprises only modules formed of devices having a trapezoidal-shaped base layer (and other elements with the appropriate shape). In this way, a ring-shaped detector can be constructed for a PET scanner.

Optionally, a system comprises only modules formed of devices having a base layer of rectangular shape (and other work elements with the appropriate shape). In this way, a gamma camera or a Compton camera can be constructed with a detector in the shape of a rectangular prism.

BRIEF DESCRIPTION OF THE DRAWINGS

Next, embodiments of the present invention will be described, by way of non-limiting example only, with reference to the accompanying drawings, in which:

Figure 1a is a schematic representation of a preferred example of the device for detecting high energy photons that can be used in a module according to the invention;

Figure 1b is a schematic representation of another preferred example of the device for detecting high energy photons that can be used in a module according to the invention;

Figures 2 to 5 are schematic representations of the step-by-step construction of the device of Figure 1a;

Figure 6a is a schematic representation of the module for detecting high energy photons according to the invention;

Figure 6b is a schematic representation of the module for detecting high energy photons of Figure 6a with a stepped configuration;

Figure 7 is a schematic representation of the module of Figure 6 mounted on a support to obtain a system for detecting high energy photons according to the invention;

Figure 8 is a schematic representation of the system for detecting high energy photons according to the invention, comprising a plurality of modules of Figure 6a;

Figures 9a and 9b are top views of still another example of the device that can be used in a module according to the present invention, the device having an isosceles trapezium shape;

Figure 10 is a schematic representation of a partial view of a gamma ray detector having a ring shape, comprising a plurality of systems of Figure 8;

Figure 11 is a schematic representation of the full view of the gamma ray detector of Figure 10;

Figure 12 is a schematic representation of a gamma ray detector having a parallelepiped shape, comprising a plurality of systems of Figure 8; Y

Figure 13 is a schematic representation of a Compton camera, comprising the detector of figure 12.

Figure 14 shows a schematic representation of a conventional PET scanner.

DESCRIPTION OF REALIZATIONS

Next, a preferred example of a device for detecting high energy photons that can be used in a module according to the invention will be described. Next, preferred embodiments of a module (comprising a plurality of these devices), a system (comprising a plurality of modules), a PET scanner, a gamma camera and a Compton camera according to the invention will be described.

In Figure 1a, a preferred example of a device that can be used in a module according to the present invention is shown. The device 10 for detecting high energy photons comprises four solid-state modular detectors pixelated at room temperature 11 in a mosaic / stacking scheme; four ASICs (integrated circuits for specific applications), one for each solid state detector 11, which act as reading elements for them; a connector of an input / output element 12 connected to the ASICs for the input and output of data (that is, mainly to obtain the values generated by the ASICs from the parameters captured by the solid state detector); and a layer of kapton 13 acting as the base layer, on which the solid state detector 11, the ASICs and the input / output element 12 are mounted. In addition, connectors 20 are provided (not visible in figure 1a) ) to connect each individual ASIC 30 to the input / output element 12. More specifically, the front plane of the kapton layer 13 is used to mount the ASICs 30, the connectors 20, the solid state detectors 11, and the entry / exit 12; in the backplane, the kapton layer 13 comprises connections to a power supply that can provide high voltage for biasing solid state detectors 11 of a second adjacent device disposed next to the side of the device shown. (This arrangement will be shown later, for example, in Figure 6a.)

In a device, a different number of solid state detectors may also be arranged.

In this example, a solid state detector was used at room temperature. The solid state detectors at room temperature can be, for example, Si, Ge, GaAs, CdTe, CdZnTe or Hgl2. Preferably, a solid state semiconductor with a high Z (high number of protons in the core) and a good mobility between hole and electron is used. However, any pixelated solid state detector can be used.

In a currently preferred example, the pixelated solid state detectors 11 are CdTe detectors approximately 2 cm x 1 cm x 0.2 cm in size; The pixel size is approximately 1mm x 1mm. The size of the detectors can also be selected differently. By using the device with a thickness of 0.2 cm, it is possible to build a thick solid state detector by combining a plurality of these devices. However, this detector will not experience problems such as collecting time on thick conventional solid-state detectors.

In addition, each solid-state detector 11 is mounted on the corresponding ASIC, so that each pixel pad is independently connected to its own channel.

In Fig. 1b, an alternative example of the device that can be used in a module according to the invention is shown schematically. In this case, a single device 10 is formed by arranging two rows of solid state detectors 11 in a single device. The solid-state detectors are mounted on a base layer 13. For each row of solid-state detectors, an input / output element 12 is provided. An advantage of this device design is that it can be made slightly thinner than if it were they combine two devices 10 according to figure 1a.

Next, a step-by-step construction (illustrated in Figures 2-5) of the device 10, which is shown in Figure 1a, will be described.

Figure 2 shows the kapton layer 13 which, at the top, disposes the input / output element 12 and the connectors 20 for connecting the input / output element 12 to each individual ASIC.

Figure 3 shows the same kapton layer 13 with an ASIC 30 mounted thereon. The ASICs have been made thinner to a thickness of 50 pm and the input / output connectors of the ASICs are made with metalized tracks, so that it is possible to connect the ASIC to the kapton layer by means of "bump-bond" or a conductive adhesive In addition, other forms of connection may be used, for example, anisotropic conductive films, such as those commercially available from 3M®. It is also possible to use a cable connection. Layer 13 shown It is designed to contain four ASICs, but obviously the number of ASICs that connect can be different according to the requirements of the device 10.

Figure 4 shows the same layer 13 with the four ASICs 30 mounted thereon. In this preferred example, after assembling the four ASICs 30, an intermediate layer (kapton layer) is arranged over the ASICs. The object of said intermediate layer is to connect each pixel pad of the corresponding solid state detector 11 to an input pixel channel in the ASIC 30. Since the size of the solid state detectors 11 can vary along one direction determined, as will be described below for some examples (eg, Figure 9a), the intermediate kapton layer serves as an interface for adjusting any deviation between the pads in the solid state detector 11 and the input channel of the ASIC 30 The intermediate layer is not necessarily available in all cases, especially if the size of the solid-state detectors is constant. Figure 5 shows the solid state detectors 11 mounted on the intermediate kapton layer. In other examples, the solid state detectors may be mounted directly on the reading element (e.g., ASIC). The pixel size in this preferred example is selected to be approximately 1 mm x 1 mm. With a pixel pad of this size, it is possible to connect the solid state detectors and the intermediate kapton layer or the ASICs and the intermediate kapton layer with a conductive adhesive and avoid the use of solder balls and, therefore, both Heat the solid state detector.

The device 10 obtained, in the preferred example shown, has a solid state detector with a length of 4 cm, a width of 2 cm and a thickness of 0.2 cm. The effective thickness of the device is approximately 2.2 mm (including the detector, the kapton layer (s) and the reduced thickness ASIC).

The device 10 obtained in this manner, and shown in Figure 1a, comprises a body having two large parallel faces (the front plane and the back plane) and substantially thin lateral edges. The operation of device 10 is as follows.

A gamma ray (high energy photon) impinges on the pixelated detector (11), which is mounted on the device (10), on its thin side (as indicated schematically in Figure 1a). In prior art devices using solid-state detectors, these are usually oriented to capture gamma rays that strike the larger face of the detector (i.e., in one direction of the electric field applied to the solid-state detector). However, the device that can be used in a module according to the invention is oriented so that it can capture gamma rays incident on the thin side of the detector (i.e. in a direction perpendicular to the direction of the electric field applied to the detector). ).

In the device that can be used in a module according to the invention, the gamma rays can also impinge on another side, but, as will be shown below, the main incidence direction will be substantially on the thin side. The probability that the photon is captured in the detector (11) increases with the distance traveled by the particle within the detector material. At the point of incidence of the captured photon, it will experience different processes, and the dominant one is the photoelectric effect. After depositing all the energy in the detector (11), the equivalent amount of electron-holes (e-h) is released in response. A high voltage (HV) is applied to the detector to allow the e-h to derive by inducing a signal at the pixel electrode, which then amplifies and processes the ASIC. The ASIC (30) will indicate the position of the point of incidence and, therefore, the coordinate of the pixel where the incident occurred. In addition, the pulse of the signal will provide information about the energy of the photon. In addition, the moment of incidence can be determined with respect to a global clock from the increasing time of the signal produced.

Next, a module for detecting high energy photons will be described.

As can be seen in Figure 6a, the module 60 comprises a plurality of devices 10 for detecting high energy photons (described above) which are connected through their input / output elements 12 to an interface, which in this case , is a printed circuit board 61 (PCB). The devices are arranged in a row with their large faces adjacent to each other. It does not matter if they are all arranged with the front sides of the base layers facing the same direction, or not. Obviously, due to the modularity of the devices 10, the module 60 can be of any size that the designers prefer. In the case of figure 6a, the module comprises thirty devices. Using the size of the devices that were previously used, the module obtains a thickness of approximately 6.6 cm.

The PCB 61 comprises connectors 62, for example, low profile, to access the ASICs 30, that is, the connectors 62 of the PCB are connected to the input / output signal connectors 12, which have access (through from the connectors 20) to the ASICs. In this way, the data processed in the ASICs 30 can be obtained in the connectors 62 and read from them. Optionally, said module comprises devices, in which are housed the pixelated solid-state detector, the reading element and the signal element of input / output, on an upper side of the kapton layer and wherein the lower side of the kapton layer comprises means for providing a high voltage for biasing the pixelized solid state detectors in an adjacent device.

Each pixelated solid state detector can essentially provide two-dimensional information about where the incidence occurred with a gamma ray, as described above. In the module, a plurality of devices are arranged so that a three-dimensional solid-state detector is obtained. Each device is a two-dimensional detector, but by placing several devices on top of each other, a three-dimensional detector is obtained. The dimensions of the module 60, using the dimensions mentioned above with respect to Figure 5, are approximately 5 cm x 2 cm x 6.6 cm and contain 48,000 voxels. Each voxel has dimensions of 1 mm x 1 mm x 2 mm and is connected to its own independent pixel channel in ASIC 30. This is equivalent to having approximately 725 channels / cm3.

The operating principle of the module remains the same. A gamma ray hits a pixelated detector. At the point of incidence of the captured photon, electrons-holes (e-h) are created. Due to the high applied voltage, the e-h drift induces a signal at the pixel electrode, which then amplifies and processes the ASIC. The ASIC will indicate the position of the point of incidence and, therefore, the coordinate of the pixel where the incident occurred. In addition, the interface or PCB (61) has data on which ASIC the event was registered. Therefore, the PCB has data on the voxel coordinate where the incident occurred.

In some embodiments, as shown in Figure 6b, to obtain a more robust module, the detectors 11 of the devices 10 comprised in the module according to the invention have different sizes and are arranged in a stepped configuration. In this way, the gaps between the pixelated detectors 11 of a device are compensated with the pixelated detectors of the adjacent devices. In summary, this configuration provides more mechanical strength for the module 60 to stand alone. Obviously, this has no further consequences for the functionality of the module.

It should be noted that the device shown in Figure 1b is especially suitable for use in such stepped configuration. A stepped configuration can be provided to a single device between the solid state detectors on one side of the device and the solid state detectors on the other side.

Next, a system for detecting high energy photons will be described.

Figure 7 illustrates an interface bus 70 to which a module 60 is connected. The interface bus in this embodiment is a PCB. The PCB 70 comprises a plurality of connectors 71, and the connectors 62 of the interface 61 are connected to the connectors 71 of the PCB 70. In this way, a system for detecting high energy photons is obtained. Said system may comprise one or more modules depending on the requirements (ie, depending on the size to be scanned). The interface bus 70 can supply the high voltages for the solid state detectors, and provide power to the ASICs, as well as read from them and write on them.

Figure 8 shows the system 80 comprising a plurality of modules 60. Said modules 60 are arranged in a row, so that a large face of an end device of a module is adjacent to a large face of an end device of the next module. (the modules are aligned in the same direction of the applied electric field within the pixelated solid state detectors).

It will be clear to the expert that many other possible module arrangements can also be used to form a system.

From the system for detecting high energy photons 80 described it is possible to obtain different medical diagnostic imaging devices.

Figure 9a illustrates a different example of a device that can be used in a module according to the invention. The device 10 comprises a body having two large parallel faces and substantially thin lateral edges, the large faces having an isosceles trapezoidal shape. This form has been obtained by giving the base layer of the device the shape of an isosceles trapezoid (and adapting the other elements of the device appropriately). The solid-state detectors 11 used in this example are in the form of an isosceles trapezoid. The ASICs 30, however, are rectangular. In these examples an intermediate kapton layer that has been mentioned above may be necessary to accommodate any displacement between the pixels in the detector 11 and the corresponding pixel channels in the ASIC 30. The ASICs 30 are indicated partially in dotted line, since that are arranged below the detectors and, therefore, are not entirely visible.

The chosen form of the device allows the construction of a detector especially suitable for a PET scanner, since they can be installed in a ring without significant spaces between the modules. In PET scanners, two photons must be registered in the same time window (coincidence) in order to form the Response Line (LOR). In the devices of the prior art, wedge-shaped recesses are present between the detectors (or modules) (see, for example, in Figure 14). Since these gaps are present at either end of the LOR, this results in a reduction in the detection efficiency of the PET scanner, which is quadratic. This means that, if the detection efficiency of a single photon of 511keV is 25%, then the probability of detecting two consecutive photons is (0.25) 2 = 0.0625, approximately 6%. With the trapezoidal isosceles shape of the detectors, the PET scanner modules can be placed in a substantially uniform ring without wedge-shaped spaces between the adjacent modules and, therefore, maintain good detection efficiency.

Figure 9b shows the same device as Figure 9a. In a preferred example, 120 trapezoidal devices that can be used in a module according to the invention together form a ring that is part of a PET scanner. The dashed arrow indicates a radial direction of the PET scanner. As can be clearly seen, the size of the pixelated solid state detector varies along this radial direction. In this example, to offset any displacement between the pixel pads of the detectors and the pixel channels of the reading elements, an intermediate kapton layer is arranged.

In this example, the height of the solid state detector H is equal to 1 cm. The length of the short base (SB) of the first detector is 19.95 mm and the length of the long base (LB) of the first detector is 20.47 mm. These measurements are based on a PET scanner with a radius of 40 cm, 120 devices along its circumference and a free space of 100 μm between adjacent devices.

Figure 10 shows a partial view of a detector 100 of a PET scanner. Said detector 100 of a PET scanner is ring-shaped and comprises a plurality of systems 80 for detecting high energy photons as described above. At least some devices of the systems comprise base layers that are in the shape of an isosceles trapezoid, which makes it possible to obtain the shape of a ring.

Basically, the systems 80 are stacked to form the ring, that is, the plurality of systems are arranged forming the ring with the lateral edges of the devices 10 adjacent to each other.

In this way, the shape of the systems defines a hermetic geometry.

Figure 11 shows a complete ring 100 (ie, the gamma ray detector) for a PET scanner. In this example, the detector 100 has a field of view (FoV) of 66 cm with a standard size gantry of 80 cm. It has 4 cm of CdTe 11 detector in the radial direction that can effectively stop 90% of the 511 keV gamma rays. This means that the detection efficiency for PET events with this device is 81%.

Accordingly, a PET scanner comprises the ring 100 of the systems 80 surrounding a platform on which the patient is located. Each system (can be understood as a sector of the ring) comprises at least one module 60, and each module comprises a plurality of devices 10.

The operation of the PET scanner comprising the devices 10 of the invention is as follows. The PET detector (100) is formed by many devices (10). After the mother nuclide disintegrates, positrons are released with energy of approximately 1 MeV. These positrons travel an average of about 1 mm before annihilation with electrons in the human body. This process of annihilation produces 2 gamma photons that move in the opposite direction, each with an energy of 511 keV (the resting mass of the pair of electrons and positrons). The two gamma rays are recorded by a solid-state detector and the Response Line (LOR) can be constructed. Due to the fact that the detector according to the invention is three-dimensional, the so-called parallax effect can be avoided. You can place a vertex point where there is an additional accumulation of radionuclide and, therefore, additional positron emitters, since many lOr will cross at the "same point". Thanks to the improved detector, the reconstruction of the LOR can be improved and the detection quality of a radionuclide concentration (which can indicate the presence of cancer cells) can also be improved.

From the above explanation and from Figures 9 to 11, it should be clear that gamma rays will mainly affect the solid state detectors on the thin side of the detectors, that is, in a direction perpendicular to the electric field applied to the detector .

In the case where the large faces of the devices 10 have a parallelogram shape (for example, a square or a rectangle), it is possible to obtain a gamma ray detector for a gamma camera.

Therefore, Figure 12 illustrates a gamma ray detector 120 for a gamma camera. Said detector 120 comprises a plurality of systems 80 as described above. The plurality of systems is arranged to form a rectangular prism with the side edges of the devices adjacent to each other. In the case of Figure 12, the gamma ray detector 120 comprises thirty systems and this will result in a detector with dimensions close to 66 cm x 66 cm. The size of the detector can be made larger without problem. This is possible due to the modularity of the devices 10, the modules 60 and the systems 80.

The operation of the gamma camera comprising the devices 10 of the invention is as follows.

It is expected that the gamma camera will detect photons within a range of energy between 140 keV (99Tcm) and 777 keV (82Rb) depending on the type of radionuclide used. A suitable collimator can be placed in front of the detector to select the direction of the photons that impinge on the camera. Photons with energy interact with the device (10) and only the information about the deposited energy and the voxel coordinates are used. The time information in this mode is not applicable.

In the SPECT representation, a gamma camera and a radionuclide that emits a positron and only a 511 keV photon are used. The SPECT image is usually obtained by taking multiple images at different angles using the gamma camera.

Figure 13 shows an example of a Compton 130 camera. The Compton scattering takes place in a close up and the second plane is the gamma ray detector 120 as described above, where the X-rays are captured after dispersing.

The advantages of this detector design for the Compton camera are:

- the detector 120 for gamma rays can be made as thick as necessary to improve its absorption efficiency for high energy gamma rays;

- the dimensions of the gamma ray detector 120 can be made as large as necessary;

- the response of detector 120 is good for a wide range of energy;

- the parallax effect in the detector 120 can be avoided since the detector has a small voxel volume. In the illustrated example, it is 1 mm x 1 mm x 2 mm. It could be much smaller if necessary, since this has to do with the pixel electrode pattern that practically has a spatial resolution of 1 pm.

Another advantageous application of the module, the system and the detectors according to the present invention may be in the CT scan. In the scan ^t C a X - ray source is placed in a side of the body of a patient and a detector is placed on the other side. The X-ray source and the detector move with respect to a patient. The photons captured by the detector give an indication of the attenuation of the X-rays by the tissue of the patient's body parts (organs, bones, etc.). By using the different levels of attenuation of different tissues, an image of a patient can be obtained.

The sensor used in CT scanning typically has to detect photons within an energy range of 80 keV - 140 keV. One of the advantages of using a pixelated solid state detector in the device that can be used in a module according to the invention is that the pixel size can vary within the same detector (which is impossible with the prior art detectors). based on bright crystals). In a preferred embodiment of the invention, a detector for CT scanning is formed from a plurality of modules or systems in which the pixel size varies within a single detector. Preferably, the pixel size in the "front part" of the device (where the photon enters the detector) is reduced, for example, 100 μm x 100 μm. The pixel size gradually increases towards the "back" of the detector. For example, the next row of pixels can be 100 pm x 200 pm, the next row 100 pm x 300 pm.

This configuration has the advantage that, with a reduced number of pixel channels, photons with lower energy can still be captured with greater precision and less parallax, while photons with higher energy are likely to be captured towards the back of the device.

Since the energy range (80 keV - 140 keV) of the photons in the CT scan is small compared to, for example, the PET scan (511 keV), the height (H as defined in Figure 9b) ) of the detector may also be smaller (if the same detector material is used). As described above, all the devices that were used to build the module to be applied in a PET scanner were each provided with 4 pixelized solid-state detectors (in the configuration shown in Figure 1a; solid state with the configuration of figure 1b). Each solid state detector had a size of 1 cm x 2 cm x 0.2 cm. The total height of the detector in this case was 4 cm. In the case of a CT scanner, for example, only one CdTe detector with a size of 0.5 cm x 2 cm x 0.2 cm would have to be arranged. (using the configuration shown in Figure 1a) to capture approximately 100% of the photons emitted by the X-ray source of the CT scanner. Clearly, in the configuration shown in Figure 1b, two detectors of this size would be available.

Claims (15)

1. Module for detecting high energy photons, comprising a plurality of devices (10) for detecting high energy photons, each of said devices comprising one or more pixelated solid state detectors (11) for detecting high energy photons; means for providing a high voltage for biasing said solid-state detectors; at least one reading element (30) connected to said pixelated solid-state detectors (11); an input / output element (12) connected to said at least one reading element (30) for input and output of data; and a base layer (13) for mounting the pixelated solid state detectors (11), said at least one reading element (30) and the input / output element (12), wherein each solid state detector pixelated is mounted on a reading element, and the one said devices are arranged one above the other and connected to each other to support each other and in such a way that the combined pixelated solid state detectors (11) of the devices (10) can provide the three-dimensional coordinates of high-photon incident points. Energy. Module (60) according to claim 1, characterized in that the base layers (13) of the devices have an isosceles trapezoidal shape with exactly one pair of parallel sides. Module (60) according to claim 2, characterized in that at least one pixelated solid state detector (11) has an isosceles trapezoidal shape with exactly one pair of parallel sides. Module (60) according to claim 1, characterized in that the base layers (13) of the devices have a rectangular shape. Module (60) according to any one of claims 1-4, characterized in that it comprises devices (10), in which one or more pixelated solid state detectors (11), said at least one reading element (30) and the input / output signal element (12) are housed in an upper side of the base layer (13) and in which the lower side of the base layer (13) comprises means for providing a high voltage for biasing the pixelated solid-state detectors (11) that are mounted on an adjacent device (10). Module (60) according to any of claims 1-5, characterized in that the devices (10) further comprise an intermediate layer mounted between the solid state detectors (11) and the reading elements (30) to adapt to any displacement between the pixel pads of the solid state detectors (11) and corresponding pixel channels in the reading elements (30). Module according to any of claims 1-6, characterized in that it comprises at least one device (10) comprising at least one pixelated solid-state detector (11), in which the size of pixel varies. Module (60) for detecting high energy photons according to any of claims 1-7, characterized in that the input / output elements (12) of the plurality of devices (10) are connected to a only interface (61). Module (60) for detecting high energy photons according to claim 8, characterized in that at least part of the solid state detectors (11) of the devices (10) are arranged in a stepped manner with respect to the other. 10. System (80) for detecting high energy photons comprising a plurality of modules (60) according to claim 8 or 9, characterized in that the plurality of interfaces (61) of the modules are connected through of suitable connectors (62, 71) to a single interface bus (70). System (80) according to claim 10, characterized in that it comprises only modules (60) in which the base layers (13) of the devices and / or the pixelated solid-state detectors (11) ) have an isosceles trapezoidal shape with exactly one pair of parallel sides. System (80) according to claim 10, characterized in that it comprises only modules (60) in which the base layers (13) of the devices have a rectangular shape. 13. Ring-shaped detector (100) for a PET scanner, comprising a plurality of systems (80) according to claim 11. 14. A gamma camera comprising a detector (120), comprising a plurality of systems (80) according to claim 12. 15. Compton camera (130) comprising a detector (120), comprising a plurality of systems (80) according to claim 12.