The present disclosure relates to detectors for detecting photons, in particular highly energetic photons. The present disclosure also provides systems and methods for monitoring metabolic activity. In particular, the present disclosure provides examples of methods and systems for determining and displaying areas showing increased metabolic activity substantially in real-time. The present disclosure further provides examples of detectors suitable for these methods.
Examples of the herein disclosed systems and methods can provide true 4D(x,y,z,t) PET scanning without the use of other imaging modalitity such as MRI, which can be of significant value for better understanding neuroscience.
One known method for measuring or determining brain activity uses electroencephalography (EEG) which relies on electrodes placed along the scalp. EEG measures voltage fluctuations resulting from ionic current within the neurons of the brain.
In some cases, it is known to use an invasive approach, in which needle-like probes are inserted inside the brain of an animal to detect signals.
Other known methods include e.g. PET scans and fMRI. Functional magnetic resonance imaging or functional MRI (fMRI) is a functional neuroimaging procedure using MRI technology that measures brain activity by detecting changes associated with blood flow. This technique relies on the fact that cerebral blood flow and neuronal activation are coupled. When an area of the brain is in use, blood flow to that region also increases.
A brain positron emission tomography (PET) scan is an imaging test that allows doctors to see how a patient's brain is functioning. The scan captures images of the activity of the brain after radioactive "tracers" have been absorbed into the bloodstream. These tracers are "attached" to compounds like glucose (sugar). Glucose is the principal fuel of the brain.
Active areas of the brain will be utilizing glucose at a higher rate than inactive areas. The nuclear decay of the radioactive tracer can be pinpointed to specific areas of the brain using a PET scan.
Like other diagnostic techniques in Nuclear Medicine, PET is based in detecting and analyzing the distribution inside the body of radioisotopes which have previously been administered to a patient. The radioisotopes may e.g. be injected into the patient's blood.
Several positron-emitting radioisotopes for medical use are known. The most commonly used is Fluorine-18, which is capable of joining a glucose tracer to get 18-fluoro-deoxy-glucose (18F-FDG). In this way, glucose that is detectable by the emission of radioactive signal is obtained.
After administration of the radioisotopes, the radioisotopes spread throughout the area of the body to be examined and tend to be taken up by e.g. cancer cells. However, the tracers also concentrate in the areas of the brain where the consumption of glucose is relatively high. The systems and methods of the present disclosure are based upon this latter phenomenon.
When the radioisotope decays, it emits a positron which after a few millimeters, or after even less than 1 mm annihilates with an electron. This produces a pair of gamma ray photons moving in opposite directions, each photon having an energy of 511 keV. This pair of gamma ray photons can be detected using a so-called PET scanner. Using the location of detection of both gamma ray photons, the Line of Response (LOR) (which is the line connecting the two locations of detection of the gamma photons) can be reconstructed. This procedure is schematically illustrated in figure 1.
Figure 1 shows a conventional PET scanner 1, in which a bed 3 is provided. Upon this bed, a body 2 of a human or an animal is schematically indicated. Along the perimeter of the PET scanner, a plurality of detectors 4 is provided. The gamma ray photons which move in opposite direction are detected respectively by detector 4a and detector 4b. Using this detection, the LOR can be reconstructed. For a scan of a body part, e.g. the brain, such a body part, e.g. a subject's head may be stabilized and immobilized in a predetermined position to make sure the subject stays still. However, from a single event the precise origin (location of radioactive decay) cannot be determined: the origin might be located anywhere along the LOR.
After collecting millions or 100 millions of such events and using dedicated image processing software for PET, a 3D image is obtained showing concentration of the radioisotope and therefore an area of brain activity. The PET scanner is usually coupled to a computer, which is responsible for measuring the amount of radioisotopes absorbed by the body, and determining the LOR's. The origin of the events, i.e. the location where the radioactive decay occurs and thus the location where the brain is assumed to be more active, may then be displayed in an image, e.g. on a screen or print-out. It would be beneficial in many different types of PET imaging if the imaging could take place substantially in real-time and with good spatial resolution and energy resolution in order to observe metabolic processes in real-time.
Each radioisotope has a characteristic decay in terms of half-life and in terms of its radiation, e.g. positron, gamma rays with specific energy level or even positron and gamma rays. When e.g. a specific type of cancer is suspected, a doctor might administer a radioisotope that is known to attach particularly well to that kind of cancer cells. To gain a better understanding of the type of cancer, various isotopes might be administered at the same time.
Depending on which body part is to be examined, and in some occasions depending on the size of a patient, a PET scanner and/or a Compton camera may be used. A Compton camera has two detection planes. Photons emitted from the source are scattered in the first plane (Compton scattering) and absorbed in the second plane (photoelectric effect). In both planes the position of the interaction and the energy deposited are measured. The detectors are operated in coincidence, so that only photons that interact with both detectors and deposit a total energy within a given time window are recorded. Using the location of detection and the energy of the photon, the point of origin of the photon can be calculated, using the so-called Compton formula.
In nuclear medical imaging (of the brain or of other body parts or organs) it is beneficial that all or most or many of the events are detected. If more of the events are detected, a smaller amount of isotopes has to be administered. The patient thus will suffer less from possible side effects from the radiation. It is of course important if the precise position of decay can be detected, i.e. high spatial resolution. Or, if the same normal dose is used, the image can be acquired in less time and hence reduce possible image smearing related to patient movements.
An advantage of using EEG for monitoring brain activity is that it can show brain activity in real-time. On the contrary, PET and fMRI have a time resolution between seconds and minutes. Particularly, in PET scans, because many events occur at the same time and one needs to link two separate impacts of gamma rays to the same event, image reconstruction algorithms are needed to reconstruct an image of the activities in the body under study.
Other advantages of EEG include the fact that the subject is not exposed to either radioisotopes or a strong magnetic field. Moreover, EEG as compared to fMRI is silent, which is useful for e.g. measuring response to auditory stimulation. On the other hand, the spatial resolution with EEG is significantly lower than the obtainable spatial resolution with the other methods.
fMRI has an important disadvantage in that it is noisy and can be experienced as claustrophobic.
It is an object of the present disclosure to provide examples of systems and methods for monitoring brain activity based on PET technology that avoid or significantly reduce one of the major disadvantages of PET mentioned before, namely the poor time resolution. It is a further object of the present disclosure to provide examples of systems and methods for monitoring metabolic activity based on PET technology substantially in real-time.
Substantially in real-time herein means that metabolic activity can be monitored (and displayed) to a researcher as decay takes place. The positions of individual events of decay as an indication of a change in metabolic activity can be visualized.
For example, for monitoring brain activity, this means that brain activity can be displayed on the same time scale as stimuli used in such research. A researcher will thus be able to link a possible change in brain activity to a specific stimulus.
 US 6,484,051
discloses a Compton imager and methods for generating three-dimensional images. The Compton imager detects Compton scattering of simultaneously or nearly simultaneously emitted gamma rays produced by a radio-nuclide. A possible location of each radio-nuclide decay is defined by the intersection of Compton direction cones corresponding to the detected gamma rays. Three-dimensional images are generated by superposition of individual locations of separate radioactive decay locations.
"Development and Modeling of a Compton Camera Tomographer Based on Room Temperature Solid State Pixel Detector" (Yónatan Calderon) discloses a Compton camera.
It is a further object of the present disclosure to provide examples of detectors for detecting photons that have both high spatial and high energy resolution.
It is a further object of the present disclosure to provide examples of improved Compton cameras, Compton imaging, and Compton-PET imaging.
It is a further object of the present disclosure to provide examples of methods and systems that allow the reduction of the dose of radioisotopes administered to a patient.
In a first aspect, a detector for detecting photons according to claim 1 is provided. The detector comprises a plurality of detector modules forming a front detector and a rear detector. The detector modules comprise a plurality of detector devices, and an interface, the detector devices having a substrate extending from a front end to a rear end, and carrying a plurality of pixelated semiconductor detector slabs, and being arranged such that the front end is closer to a source of the photons than the rear end. The semiconductor detector slabs are arranged on readout circuits, and the detector devices have an input/output element at or near the rear end of the substrate. A front group of one or more of the semiconductor detector slabs arranged close to the front end of the substrate are made from a semiconductor material configured to promote photon scattering and form the front detector, and a rear group of one or more of the semiconductor detector slabs arranged closer to the second end of the substrate than the front group are made from a semiconductor material configured to promote photon absorption, and form the rear detector.
The arrangement of the front and rear detectors makes the detector suitable for Compton imaging. With the detector according to the first aspect, the front and rear detector slabs are arranged on common substrates with the readout circuit arranged directly underneath the semiconductor slabs. The readout circuit can be attached to the semiconductor via bonding or gluing or otherwise. The input/output element is provided at or near the rear end of the substrate, so that no parasitic elements or electronics are arranged between the front and the rear detector.
In some examples, the substrate may be a kapton layer, in particular a thinned kapton printed circuit board.
In some examples, a distance between the rearward most detector slab of the front detector and the forward most detector slab of the rear detector may be greater than a distance between the detector slabs of the front detector and than a distance between the detector slabs of the rear detector. An increased distance between the front and rear detector can increase the accuracy of the Compton cone calculation and lead to better spatial resolution.
In some examples, the distance between the rearward most detector slab of the front detector and the forward most detector slab of the rear detector is between 10 and 30 centimetres, specifically between 10 and 20 centimetres.
In some examples, the detector slabs of the front detector may be made from silicon. Generally speaking, low Z semiconductor material is suitable for promoting photon scattering, whereas high Z semiconductor material is suitable for promoting photon absorption. High Z semiconductor material is therefore more suitable for the outer detector. In some examples, the detector slabs of the outer detector may be made from CdTe or CdZnTe. Other suitable semiconductors can include GaAs, TIBr, PbS, CaTiO3 or Hgl2.
The atomic number or proton number (symbol Z) of a chemical element is the number of protons found in the nucleus of an atom. It is identical to the charge number of the nucleus. Low Z semiconductor material may be regarded all throughout the present disclosure as a semiconductor having Z below 20. High Z semiconductor material may herein be regarded as a semiconductor having Z above 20.
In some examples, the front detector and rear detector have a substantially circular cross-section, and the substrates of the detector devices have an isosceles trapezoidal shape. In this aspect, a particularly beneficial PET detector can be provided. The front detector and rear detector may thus be arranged as concentric annular detectors, the front detector being the inner detector ring and the rear detector being the outer detector ring.
Also disclosed (but not claimed) is a method for real-time visualization of metabolic activity in a subject using annular inner and outer ring detectors (i.e. a PET detector arrangement) is provided. The subject has previously been administered a radioactive tracer and positioned such that a portion of a body of the subject is at least partially provided within a photon detector field of view (FOV). The method comprises detecting radioactive decay of the tracer within the photon detector, calculating a position of an origin of the radioactive decay and displaying the position of the origin of the radioactive decay substantially in real-time. The radioactive tracer comprises a radioactive isotope which in decay emits a positron and a separate gamma ray, and the photon detector comprises an inner ring detector, and an outer ring detector. Impacts of photons in the inner and outer ring detectors belonging to the same event of radioactive decay, wherein the separate gamma ray is detected in both the inner and the outer ring detector can be detected. And in relation with the impacts belonging to the same event of radioactive decay Line-of-Response calculation and Compton cone calculation based on Compton scattering from the inner ring detector is used to calculate an origin of the event.
In this aspect, real-time visualization of metabolic activity is made possible by making use of specific radioactive isotopes (e.g. 22
V) that in decay emit a positron and an additional gamma ray. The positron quickly annihilates with an electron to produce two back-to-back gamma rays moving in opposite directions. The Line of Response of these back-to-back gamma rays may be calculated in a "classical" way. In the present method, the additional information that the additional gamma ray can provide is used to more quickly determine the origin of radioactive decay (i.e. the location of the isotope, which can indicate an area of increased metabolic activity). In a classical way for any two back-to-back gamma rays, there are an infinite number of positions along the Line of Response from which the impacts may originate. It is only by combining thousands of Lines of Response that the origin of decay can be determined. In the method herein provided, the additional gamma ray can provide for a single event the information to determine the point along the Line of Response that is the origin of the decay.
In accordance with this aspect, also events based on random coincidences (i.e. two separate gamma rays are detected at the same time but actually do not belong to the same event: the LOR calculation in this case would render a wrong result) can be excluded. The angle measured in Compton scattering cannot be reconciled with the apparent LOR. The extra information from Compton scattering can thus be used to discard such random events.
In some examples, specific events complying with specific rules of energy deposits may be selected and linked together. For example, events in which two back-to-back gamma rays are detected in the outer ring or rear detector, and a separate gamma ray scatters in the inner or front detector and is detected both in the inner/front and in the outer/rear detector may be selected and separated from other events. The calculation of these specific events comprises LOR calculation for the back-to-back gamma rays, and Compton cone calculation for the separate gamma ray. The intersection of the Compton Cone with the LOR determines the point of origin.
In some examples, calculating the position of the origin of the radioactive decay may furthermore comprise determining energy deposits of the at least four impacts. The energy deposits may be used to determine impacts that belong to the same event. For example, the energy of the separate gamma ray is known for specific isotopes, e.g. 1274.5 keV for 22
Na. If the separate gamma ray scatters in the inner detector and is detected in both the inner and outer detector, the photons of these impacts must have a total energy corresponding to approximately 1274.5 keV. Separate impacts can thus be linked to the same event, based also on a timestamp of the registered impacts. Similarly, back-to-back gamma rays only detected in the outer detector each must have an energy deposit of approximately 511 keV to qualify as the same event. Similarly again, if one of the back-to-back gamma rays is scattered and detected by both the inner and the outer detectors, the energy deposits must have a sum of approximately 511 keV in order to be recognized as originating from the same event.
However, the use of this radioisotope is merely one example of the use of a detector as herein described. In yet a further aspect, a method for real-time visualization of metabolic activity in a subject using front/inner and rear/outer detectors is provided. The subject has previously been administered a radioactive tracer which in decay emits three separate gamma rays and the subject is positioned such that a portion of a body of the subject is at least partially provided within a photon detector field of view (FOV). The method comprises detecting radioactive decay of the tracer within the photon detector, calculating a position of an origin of the radioactive decay and displaying the position of the origin of the radioactive decay event by event in quasi real-time. The photon detector comprises an inner ring / front detector, and an outer ring / rear detector. Scattering and absorption of photons in the inner/front and outer/rear detectors respectively belonging to the same event of radioactive decay can be detected. And in relation with the impacts that belong to the same event of radioactive decay Compton cone calculation for the three gamma rays based on Compton scattering from the inner ring / front detector and absorption in the outer ring/rear detector is used to calculate an origin of the event.
In this aspect, real-time visualization of metabolic activity is made possible by making use of specific radioactive isotopes e.g. 94
Tc that in decay emits three gamma rays. 94
Tc in decay emits three gamma rays at 701 keV, 849keV and 871 keV. 94
Tc is however merely one example of a radioisotope that may be used in this aspect.
The three separate gamma rays, in some cases, may scatter in the front/inner detector and be absorbed in the rear/outer detector. The events in the front/inner detector and rear/outer detector linked to the same gamma ray may be linked together based on the energy deposits in the separate detectors. The origin of the decay can immediately be determined by Compton cone calculation for each of the three gamma rays. The three Compton cones will intersect at a single point, which is the origin of the nuclear decay. Again, as compared to "classic" PET scanners or "classic" Compton cameras, the need to combine many events together to determine the origin of decay is obviated with the detector as herein provided.
It should be clear however, that the detectors as herein provided may be used in combination with many different radioisotopes with different decay characteristics.
Also disclosed (but not claimed) is a method for real-time monitoring of activity in a brain of a subject is provided. The subject has previously been administered a radioactive tracer, and the method comprises positioning the subject such that the brain of the subject is at least partially provided within a photon detector, providing a stimulus to the subject and a method for real-time visualization according to examples of the previous aspect.
In this aspect, a researcher may provide different kinds of stimuli to a subject. Thanks to the fact that the position of radioactive decay can be calculated and visualized substantially in real-time, event-by-event, a researcher may see or monitor different areas of the brain that become active in response to the stimuli substantially in real-time. For this type of imaging even relatively low activity can be used. In response to what is displayed to the researcher, he/she can adapt the stimuli (which might be e.g. tactile, visual, auditory or combinations thereof) in real-time to further the investigation. The spatial resolution of PET scanners may be combined with a temporal resolution to e.g. EEG.
In a further aspect, a system for real-time visualization of metabolic activity in a subject which has previously been administered a radioactive tracer is provided. The system comprises a detector according to any of the examples herein disclosed, and a computing system for determining impacts of photons in the front/inner and rear/outer detectors, for calculating a position of an origin of the radioactive decay. In relation with the impacts Line-of-Response calculation and/or Compton cone calculation based on Compton scattering from the front/inner detector, and for generating a video signal reflecting the origin of the radioactive decay, and a device capable of receiving the video signal and reproducing the video signal on a screen.
In yet a further aspect, a system for real-time visualization of metabolic activity in a subject is provided. The subject has previously been administered a radioactive tracer. The system comprises a photon detector comprising an inner ring detector, and an outer ring detector, and a computing system for determining at least four impacts of photons in the inner and outer ring detectors, for calculating a position of an origin of the radioactive decay using in relation with the at least four impacts Line-of-Response calculation and Compton cone calculation based on Compton scattering from the first detector, and for generating a video signal reflecting the origin of the radioactive decay. Alternatively, in the case of decay with three separate gamma rays as previously described, Compton calculation may be done for three separate gamma rays. The system further comprises a device capable of receiving the video signal and reproducing the video signal on a screen.
In some examples, the screen may show a 3D scatter plot in the volume of interest of a subject (this volume of interest could be the brain as an example) where every dot represents an event of decay. Each dot may last a fraction of a second to make it possible for the human eye to register such a signal. One can also use computer algorithms to sense a small signal and visualize it, which cannot be detected by the naked eye. Such algorithms may be based on changes of statistical significance, e.g. more than 5 standard deviations above nominal conditions or as differential change, over an area of interest
In some examples, the screen may show the events of decay (for example in the scatter plot) in combination with a representation of the part of the body of the subject. Such a representation may have been obtained previously in a CT or MRI scan.
In some examples, a digital signal processor connected with the inner and the outer ring detectors may be used for calculating the origin of the radioactive decay, i.e. an area of increased brain activity. Digital signal processors (DSP) may be configured for a specific task and perform this task very rapidly. The calculation of the Line-of-Response, and/or the Compton cones and intersections between them are covered by relatively simple mathematical equations. DSP's may thus provide the speed of calculation needed to be able to monitor brain activity in real-time.
In some examples, a distance between an inner rim of the outer ring detector (or innermost detecting portion of the outer ring detector) and an outer rim of the inner ring detector (or outermost detecting portion of the inner ring detector) is at least 5 cm, and preferably 10 cm or more. Increasing the distance between the inner and the outer detector reduces the error on the angle that defines the Compton cone, due to the voxel size. This improves the imaging resolution, since this will at the same time reduce the uncertainty or error in the intersection of the LOR and the Compton Cone
BRIEF DESCRIPTION OF THE DRAWINGS
Non-limiting examples of the present disclosure will be described in the following, with reference to the appended drawings, in which:
Figure 1 schematically illustrates the calculation of Line of Response of back-to-back gamma rays as used in PET scanners;
Figure 2a schematically illustrates a photon detector that may be used in examples of the method and systems for real-time visualization of metabolic activity;
Figure 2b schematically illustrates a calculation of an origin of nuclear decay combining LOR calculation and Compton cone calculation;
Figures 3a - 3d schematically illustrate details of devices and modules that built-up a detector and an example of a detector having an inner and outer ring detector;
Figure 3e schematically illustrates an alternative detector device which may be used in examples of detector systems disclosed herein and figure 3f schematically results in a PET detector built up from such detector devices;
Figures 3g - 3k schematically illustrate a build-up of devices into modules, and of modules into detection systems and into detectors according to examples of the present disclosure;
Figure 4a schematically illustrates an example of a system for real-time visualization of metabolic activity;
Figure 4b schematically illustrates a process which may be used in a digital signal processor in examples of the system of figure 4a;
Figure 4c schematically illustrates a process which may be used to calculate an origin of radioactive decay in examples of the present disclosure;
Figure 4d schematically illustrates an alternative process which may be used to calculate an origin of radioactive decay in examples of the present disclosure;
Figures 5a and 5b schematically illustrate a detector arrangement involving a plurality of detectors; and
Figures 6a - 6c schematically illustrate a further example of a detector device which may be used to build up detectors and may be used in methods disclosed herein.
DETAILED DESCRIPTION OF EXAMPLES
Figure 1 has been previously discussed.
Figure 2a schematically illustrates a photon detector that may be used in examples of the method and systems for real-time visualization of brain activity. In figure 2a, the photon detector includes an inner ring detector and an outer ring detector, circumferentially arranged outside the inner ring detector.
The spatial resolution is related to the pixel/voxel size of semiconductor detector. Examples of such detectors are described e.g. in PCT/EP2009/061633
, published as WO 2010/034619
. The herein provided detectors are based on the use of detector devices and modules that have certain aspects in common with the detector devices and modules disclosed in this document.
Displayed in figure 2a is an event in which the back-to-back gamma rays resulting from positron-electron annihilation are detected in the outer detector. The corresponding additional gamma ray is scattered in the inner detector and is detected in both the inner and the outer detector. The inner detector effectively works as a Compton Camera scatterer. Specific radioisotopes show this behaviour at decay.
Figure 2b schematically illustrates a calculation of an origin of nuclear decay combining LOR calculation and Compton cone calculation. For the specific type of event displayed in figure 2a, a Line of Response may be calculated for the back-to-back gamma rays. A Compton cone may be calculated for the photons detected from the Compton scattering. It may be seen in figure 2b, that the Compton cone can intersect in more than one point with the Line of Response. However, only one of these points can actually be the origin of the radioactive decay, namely the point inside the inner ring detector.
Figure 3a schematically illustrates an example of a detector device which may be advantageously used to build a photon detector, e.g combining the inner and/or the outer ring detector as illustrated in figures 2a and 2b. Figures 3g - 3i illustrate more details of the build-up of the detector devices, whereas figures 3j and 3k illustrate the build-up of the detector modules and detector systems respectively.
The device 10 for detecting highly energetic photons may comprise a plurality of modular pixelated room temperature semiconductor detector slabs 11, 21 in a tiled/stacked scheme. The device 10 comprises an isosceles trapezoidal substrate 13 carrying the semiconductor detector slabs. The substrate has a first front end 13a, and a second rear end 13b. The detector device configured to be positioned with respect to a subject in such a way that the front end 13a is closer to the subject.
In the device of this example, six pixelated detector slabs are shown. Four detector slabs 21 that are closer to the rear end 13b (i.e. and thus further away from the subject) are grouped together to ultimately form the rear detector (and in the case of figure 2, the outer detector ring). Two detector slabs 11 closer to the front end 13a are grouped together to ultimately form a front detector (and in the case of figure 2, the inner detector ring).
In the example of figure 3a, a distance between the rearward most detector slab of the front detector (the second detector slab when seen from the front end 13a towards the rear end 13b) and the forward most detector slab of the rear detector (the third detector slab when seen from the front end 13a towards the rear end 13b) is greater than a distance between the detector slabs of the front detector and also than a distance between the detector slabs of the rear detector.
The distance between the rearward most detector slab of the front detector and the forward most detector slab of the rear detector may be between approximately 10 and 30 centimetres, specifically between 10 and 20 centimetres. The calculation of the Compton cone as previously illustrated with reference to figure 2b can be improved with increased distance between the inner detector and the outer detector (i.e. the front detector and the rear detector).
A distance between consecutive detector slabs of the front/inner detector may be between 0.5 and 2 centimetres. In the example of figure 3, a length of the semiconductor slabs 11 in the direction from the front end to the rear end may be 1 cm. And a distance between the consecutive detector slabs may be approximately 0.5 cm. A distance between the front/inner detector and the rear/outer detector in the example of figure 3a may be 10 - 15 cm approximately.
The front detector slabs 11 are made from a semiconductor material that is configured to promote scattering of photons. A low Z semiconductor material (i.e. Z value below 20) promotes scattering, rather than absorption. The outer detector slabs 21 is made from a semiconductor material configured to promote absorbing photons. A high Z (Z above 20) semiconductor material is configured to promote absorption.
Room temperature semiconductor detectors may be e.g. Si, GaAs, CdTe, CdZnTe, TIBr, PbS, CaTiO3 or Hgl2. The inner detector slabs 11 may be made in particular from Silicon. The outer detector slabs in the example of figure 3a may be made e.g. from CdTe, or CdZnTe.
Pixelated detector slabs may be used. The approximate size can be of 2cm x 1cm x 0.2cm for the outermost/most rearward detector slab 21; the pixel size may be approximately of 1 mm x 1 mm. Since the shape of the devices 10 is trapezoidal, the dimensions of the detector slabs can vary within the same device. The size of the detectors may also be chosen differently. 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. This detector will however not suffer from problems such as time collection in thick conventional semiconductor detectors.
The trapezoidal shape of the detector device means that the width of the substrate at the first end is smaller than the width of the substrate at the second end.
Six ASICs (application-specific integrated circuit), one for each semiconductor detector slab 11, 21 can act as readout elements for them. In other examples, it is possible to use more than one ASIC to act as readout element for a single slab. The ASICs and (pixelated) semiconductor slabs may be attached with bump bonding or conductive glue, so that each pixel pad can be connected independently to its own front-end readout channel.
An input/output element connector 12, at the backend of element 10, is connected to the ASICs for data input and output (that is, mainly for obtaining the values generated by the ASICs from the semiconductor detector captured parameters) and a kapton layer 13 which may be a kapton PCB acts as a base layer, upon which the semiconductor detector slabs 11,21, the ASICs, and the input/output element 12 are mounted.
With reference to figures 3g - 3i, wire bonding pads 20 may be provided for connecting each individual ASIC 30 to the kapton layer 13, and then to input/output element 12. More specifically, the top plane of the kapton layer 13 is in this example used for mounting the ASICs 30, wire bonding 20, the semiconductor detector slabs 11, and the input/output element 12. In the bottom plane, the kapton layer 13 may comprise connections to a power supply that can provide High voltage to polarize semiconductor detector slabs 11 of a neighbouring second device arranged next to the shown device. The slabs may be piled on the top of each other and the back side of one element 10 that provides the High Voltage and may be connected to the top of the slab in the slab beneath element 10 via double adhesive conducting thin tape.
Figures 3b and 3j schematically illustrate a module 60 comprising a plurality of detector devices 10 for detecting highly energetic photons (as illustrated in figure 3a) which are connected through their input/output elements 12 to an interface, which in this case may be a printed circuit board 61 (PCB) The devices are arranged in a row with their large faces adjacent to each other.
The PCB 61 may comprises connectors (e.g. plugs) 62, for example, with low profile, for accessing the ASICs 30, that is, the connectors of the PCB are connected with the input/output signal connectors 12, which have access (through the kapton and wire bonding) to the ASICs. This way, the processed data in the ASICs 30 can be obtained in the connectors and read from them.
Each pixelated semiconductor detector slab essentially can give two dimensional information about where impact with a gamma ray occurred, because they are pixelated In the module, a plurality of devices is arranged in such a way that a three-dimensional semiconductor detector is obtained. Every device is a two-dimensional detector, but by providing a number of devices on top of each other, a three-dimensional detector is obtained.
The working principle of the module 60 is as follows: A gamma ray impacts a pixelated detector. At the point of impact of the captured photon, electron-holes (e-h) are created. Due to the applied high voltage, the e-h drift inducing a signal on the pixel electrode, which is later amplified and processed by the ASIC. The ASIC will indicate the position of the impact point and hence the coordinate of the pixel where impact occurred. Additionally, the ASIC can provide the energy and the time stamp for the event relative to a global time clock. The interface data bus or PCB 61 has data on which ASIC the event was registered. The PCB thus has data about the coordinate of the voxel where impact occurred.
An impact in this sense can be a scattering, or absorption. Whether or not there is scattering or absorption can be determined from the energy deposited. Scatter events can be linked to each other by checking whether the sum of energy deposits, within the coincidence time window adds up to an amount of energy expected from radiation emitted.
In some examples, spaces between different detector devices may be filled up with lightweight material which is transparent to radiation. In particular, the spaces that may be present between the most rearward semiconductor slab of the front detector and the most forward semiconductor slab of the rear detector may be filled up with such material.
In one example, a cardboard honeycomb material may be inserted in between detector devices to increase the stiffness of the modules. In a further example, aerogel may be used instead.
From a plurality of modules, a detector system 80 can be made as illustrated in figures 3j and 3k. A detector system 80 comprises a plurality of modules 60. The interface 61 of each of the modules is connected with connectors to interface bus 70. The interface bus 70 may have suitable receptors/sockets for the connectors/plugs of the modules.
Figure 3c shows a partial view of a detector 100 including both an inner ring detector and an outer ring detector. The detector 100 comprises a plurality of systems 80 for detecting highly energetic photons. Each of these systems comprises a plurality of the modules 60 schematically illustrated in figure 3b arranged axially one behind the other. At least some devices of the systems comprise base layers having a shape of an isosceles trapezoid, which allows obtaining the shape of a ring.
Basically, the detector systems 80 are piled to form the ring, that is, the plurality of systems are arranged forming the outer ring detector and inner ring detector with the side edges of the devices 10 adjacent to each other. This way, the shape of the systems forms a hermetic geometry. The detector systems 80 are held together by a support structure comprising front and rear plates 90 clamping the devices/modules together in an axial direction. The plates are held by support bars 95. The support bars may be suspended in a suitable support structure
As a result a detector is formed wherein the space between the inner detector ring and the outer detector ring is essentially formed by the layers of kapton (or an alternative suitable substrate for the semiconductor detector slabs 11, 21). In this example, there is minimum of passive material, between the inner and the outer detector, in order not to alter the energy and the direction of the scattered photon.
Figure 3d schematically illustrates a frontal view of the same detector.
Figure 3e shows a further example of a detector device 10a, that may be used to build modules, and systems, and ultimately a detector as illustrated in figures 3b - 3d. Contrary to what was shown in figure 3a, in this case, the detectors 11, 21 are evenly spaced on the substrate. In figure 3a, groups of detectors ultimately forming the inner detector and groups of detectors ultimately forming the outer detector are spaced more closely together.
In the example of figure 3e, the three inner detector slabs 11 ultimately will form the inner ring detector 400 and are made from a low Z semiconductor material promoting scattering. The three outer detector slabs 21 will ultimately form the outer ring detector 300 and are made from a high Z semiconductor material promoting absorption. It will be clear that different number of detector slabs may be chosen, and they may be grouped together differently as well.
Figure 3f illustrates a further view of the PET detector 100 built up from the detector device illustrated in figure 3e. Figure 3f shows how semiconductor slabs 11 together form the inner detector, whereas the outer semiconductor slabs 21 together form the outer detector. Several of the detector devices may be grouped together to form a module. An interface 70 collects data from several modules.
Figure 4a schematically illustrates an example of a system for real-time visualization of brain activity. In the system of figure 4a, a detector comprises an inner ring detector 400 and an outer ring detector 300. The outer ring detector 300 is arranged circumferentially around the inner ring detector 400. The combined detector including inner ring 400 and outer ring 300 may be made in accordance with the example shown in figures 3a - 3f.
A subject may be positioned with a portion of his/her body (e.g. his/her head) at least partially in the inner ring detector 400. The subject may have previously been administered a radioactive tracer comprising a radioisotope having a specific decay in which a positron and an additional gamma ray are emitted. The positron almost immediately annihilates with an electron to produce two back-to-back gamma rays. In total, there are thus three gamma rays, of which two move in opposite directions.
A researcher may expose a subject to different stimuli which may include visual, auditory and tactile stimuli or combinations thereof. In response to different stimuli, different parts or areas of the brain may be more active, attracting glucose. The radioactive tracers comprise glucose, and thus the radioactive tracers are concentrated in areas of increased brain activity. However, the same systems and methods may be useful for other PET imaging applications e.g. in oncology or pharmacokinetics.
Each of the systems 80 as illustrated in figure 3c may be connected to Digital Signal Processor 500. When using the built-up of the detector in accordance with the examples of figures 3a - 3i, no separate connections between the DSP and the inner detector are needed. All signals are transmitted at the second ends 13b of substrates 13.
The signals may comprise information as to in which voxel an impact was detected. The DSP 500 can from this information calculate the origin of radioactive decay and send the information to display 600. Substantially in real time, 3D scatter plots of the intersection points between the LOR and Compton cone can be continuously reproduced.
Thanks to the additional information derived from the additional gamma ray, the origin of increased radioactive decay (and thus the area of increased brain activity) may be calculated more rapidly with the DSP and visualized substantially in real-time on display 600.
A researcher may thus adapt the stimuli to further investigate. In prior art systems, because of the delay involved in image reconstruction algorithms the valuable information is not available as a researcher is investigating.
In some examples, a display for which the user can define the time window in which the data is accumulated and then displayed in real time. In order for the human eyes to see real time changes from one frame to another, the accumulation period may be around 20msec, or 50Hz.
Figure 4b schematically illustrates a process which may be used in a digital signal processor in examples of the system of figure 4a. At block 510, the signals may be received from the inner and outer detectors. These signals comprise information as to which voxel within each detector has detected an event and the corresponding time stamp. These signals may also comprise information as to the energy deposition of each of the impacts. At block 520, the geometric location of the points of impact may be calculated (this may be implemented e.g. as a look-up table linking each voxel to a specific geometric location) as well as the corresponding energy deposits.
At block 530, the origin of decay of specific events may be calculated. An example for doing this will be explained with reference to figure 4b. A video signal reflecting the origin of decay for these events can be generated at block 540. This video signal may then be sent at block 550 to a device capable of reproducing the video signal, e.g. a computer screen.
Figure 4c schematically illustrates a process which may be used to calculate an origin of radioactive decay at block 530, also illustrated in figure 4b. At block 531, photons detected by the inner and outer detectors may be linked to each other by taking into account the energy deposits and the time coincidence. For example, back-to-back gamma rays each have an energy of 511 keV. A gamma ray that is scattered in the inner detector and is detected by both the inner and the outer detector will divide its energy between the inner and outer detector. The sum of the energy deposits thus has to be 511 keV if the gamma ray corresponds to one of the back-to-back gamma rays. On the other hand, if the gamma ray corresponds to the additional gamma ray as a direct result of the decay, its total energy will depend on the isotope used. For example, isotope 22
Na has an additional high energy photon with an energy of 1274.5 keV, for isotope 44
SC this is 1157 keV and for 48
V, this is 1312.1 keV. It will be clear that other isotopes might be used and that the corresponding energy for the chosen isotope will be known. Since the composition of the radioactive tracer is known, events that sum up to the correct amounts may be linked together. For example, in the case of 22
Na, events for which the energy deposition in the inner detector and the outer detector sum up to 1274.5 keV can be linked together.
At block 533 photons that are detected within a short time span from each other, i.e. the same time window and satisfying the corresponding energy equation may be linked to each other and linked to the same event.
It is not necessary to calculate the origin of radioactive decay for every single event. In an example, only the events wherein the additional separate gamma ray is scattered in the inner detector and the back-to-back gamma rays are not scattered and only detected in the outer detector are selected. It is advantageous to select inter alia
these events because it is known that the photons detected in the inner and outer detector must satisfy the equations for Compton scattering, whereas the back-to-back gamma rays satisfy the equation governing the Line of Response. But
The intersection of a Compton cone calculated at block 537 and the Line of Response calculated at 536 corresponds to the position of the origin of radioactive decay. This position may thus be calculated at block 539.
In other examples, other events may (additionally) be selected. Events for which one or both of the two back-to-back gamma rays are detected in the inner ring and the outer ring may (also) be selected. The information from these events may also be used for the visualization of brain activity.
In yet further examples, yet other events may (additionally) be selected. Each type of event may satisfy different energy requirements and different geometric equations for the calculation of the origin of radioactive decay.
Figure 4d schematically illustrates another example of visualization of radioactive decay in which a radioisotope with a different type of decay is shown. The detector used in this example may be a front detector and a rear detector, or an inner and outer detector as in the previous example.
In the example of figure 4d, a subject may have been administered a radioisotope which in decay emits three separate gamma rays. The three gamma rays may have different energies. If the gamma rays scatter in the front detector and absorbed in the rear detector, events in the same time window can be linked together at block 523 by checking the energy deposits. Photons belonging to the same event can thus be determined at block 533.
At block 536, events wherein all three gamma rays are scattered in the front detector and absorbed in the rear detector can be selected. For each of the three gamma rays, Compton cone calculation for each of the three gamma rays can be performed at block 537. The three gamma rays that belong to a single event of decay can be combined to calculate the intersection of the three Compton cones at block 538. The intersection of the three cones gives the origin of a single event of decay. As the decay occurs, the event may be visualized.
Figures 5a and 5b schematically illustrate an example of a detector arrangement. Figure 5a provides an isometric view, whereas figure 5b provides a side view. In figure 5a, some continuous lines and some interrupted lines are used to indicate separate detector devices, detector modules, and detector systems which in reality are not necessarily visible in this way.
The detector arrangement of figure 5a includes a top detector 1000 a bottom detector 1100 and a side detector 1200. It is noted that even though a plurality of detectors is used here in combination to provide more robust angle coverage, a single one of these detectors can function as a stand-alone Compton camera.
In the example arrangement of figure 5a, a subject or part of a subject of which imaging is to be performed is to be positioned in a space between the detectors 100, 1100 and 1200. The left side in the view of figure 5a is open for easier introduction of the body (part) of the subject for imaging. In an alternative example, a fourth detector may be arranged to provide a detector which is substantially closed in cross-section.
Each of the detectors 1000, 1100 and 1200 include an inner detector 400 and an outer detector, the inner detector arranged closer to the subject. The inner detector 400 and outer detector 300 are formed by semiconductor slabs of detector devices which are similar to the detector devices illustrated with reference to figure 3a. In the example of figure 5a, the detector devices are substantially rectangular, rather than trapezoidal. Alternatively, also in this arrangement, trapezoidal detector devices could be used. When using trapezoidal detector devices, detector 1000 will take the shape of sector of hollow cylinder.
A group of semiconductor slabs 11 closer to the subject (i.e. closer to the first end 13a of the substrate) form the inner detector 400. These semiconductor slabs may be made from silicon. A group of semiconductor slabs 21 further away from the subject form the outer detector 300.
Each of the detectors 1000, 1100 and 1200 are constructed in a similar manner. A plurality of devices 10 comprising semiconductor slabs 11, 21 and input/output elements 12 are arranged on a substrate together form a module 60 having a single interface. A number of modules 60 may be grouped together to from a system 80 sharing an interface bus, which may be a printed circuit board (PCB).
Figures 6a - 6c schematically illustrate a further example of a detector device 10' which may be used to build up detectors and may be used in methods disclosed herein. Figure 6a provides an isometric view, figure 6b a side view and figure 6c a top view.
In contrast to the detector devices illustrated in figure 3, the detector device 10b of this example has a substrate 13 and semiconductor slabs are provided on top side 18 and on bottom side 19. A single detector device 10b is formed in this case by providing two rows of solid-state detectors 11, 21 on a single device. The semiconductor detectors are mounted on a base layer 13. For each row of semiconductor detectors, an input/output element 12 is provided. An advantage of this design of the device is that slab 11 can be made thinner for better charge collection, than when two devices 10 according to fig. 3a are combined.
The semiconductor slabs 11 closer to first end 13a together form the inner detector 400, and semiconductor slabs 21 closer to the second end together from the outer selector 300. Figure 6c schematically illustrates interface 61 which may be shared by a plurality of devices 10b.
Although only a number of examples have been disclosed herein, other alternatives, modifications, uses and/or equivalents thereof are possible. Thus, the scope of the present disclosure should not be limited by particular examples, but should be determined only by a fair reading of the claims that follow.
1. Detektor (100; 1000, 1100, 1200) zum Detektieren von Photonen, umfassend
mehrere Detektormodule, die einen vorderen Detektor (400) und einen hinteren Detektor (300) bilden,
die Detektormodule mehrere Detektorvorrichtungen (10; 10b) und eine Schnittstelle (61) umfassen,
die Detektorvorrichtungen (10; 10b) ein Substrat (13) aufweisen, das sich von einem vorderen Ende (13a) zu einem hinteren Ende (13b) erstreckt, und mehrere pixelierte Halbleiterdetektorplatten (11, 21) trägt und derart angeordnet ist, dass das vordere Ende (13a) näher an einer Photonenquelle liegt als das hintere Ende (13b),
die Halbleiterdetektorplatten auf Ausleseschaltungen (30) angeordnet sind, und die Detektorvorrichtungen ein Eingabe- /Ausgabeelement (12) am oder nahe dem hinteren Ende (13b) des Substrats aufweisen, dadurch gekennzeichnet, dass
eine vordere Gruppe einer oder mehrerer der Halbleiterdetektorplatten (11), die nahe dem vorderen Ende (13a) des Substrats angeordnet ist, aus einem Halbleitermaterial hergestellt ist, das zum Fördern von Photonenstreuung und Bilden des vorderen Detektors (400) ausgelegt ist, und dadurch, dass
eine hintere Gruppe einer oder mehrerer der Halbleiterdetektorplatten (21), die näher am hinteren Ende (13b) des Substrats angeordnet ist als die vordere Gruppe, aus einem Halbleitermaterial hergestellt ist, das zum Fördern von Photonenabsorption und Bilden des hinteren Detektors (300) ausgelegt ist, und
wobei die vordere Gruppe der Halbleiterdetektorplatten und die hintere Gruppe der Halbleiterdetektorplatten derart angeordnet sind, dass der Detektor für Compton-Bildgebung geeignet ist.
2. Detektor nach Anspruch 1, wobei das Substrat (13) eine Kapton-Schicht ist.
3. Detektor nach Anspruch 1 oder 2, wobei ein Abstand zwischen der hintersten Detektorplatte des vorderen Detektors (400) und der vordersten Detektorplatte des hinteren Detektors (300) größer ist als ein Abstand zwischen den Detektorplatten des vorderen Detektors und als ein Abstand zwischen den Detektorplatten des hinteren Detektors.
4. Detektor nach Anspruch 3, wobei der Abstand zwischen der hintersten Detektorplatte des vorderen Detektors (400) und der vordersten Detektorplatte des hinteren Detektors (300) zwischen 5 und 30 cm, insbesondere zwischen 10 und 20 cm liegt.
5. Detektor nach einem der Ansprüche 1 - 4, wobei ein Abstand zwischen den Detektorplatten des vorderen Detektors zwischen 0,5 und 2 cm liegt.
6. Detektor nach einem der Ansprüche 1 - 5, wobei ein Abstand zwischen den Detektorplatten des hinteren Detektors zwischen 0,5 und 2 cm liegt.
7. Detektor nach einem der Ansprüche 1 - 6, wobei die Detektorplatten (11) des vorderen Detektors (400) aus Silizium hergestellt sind.
8. Detektor nach einem der Ansprüche 1 - 7, wobei die Detektorplatten (21) des hinteren Detektors (300) aus CdTe oder CdZnTe hergestellt sind.
9. Detektor nach einem der Ansprüche 1 - 8, wobei die Ausleseschaltungen ASICs sind.
10. Detektor nach einem der Ansprüche 1 - 9, wobei der vordere Detektor (400) und der hintere Detektor (300) einen im Wesentlichen kreisförmigen Querschnitt aufweisen, und wobei die Substrate die Form eines gleichschenkligen Trapezes aufweisen.
11. Detektor nach einem der Ansprüche 1 - 9, wobei die Substrate eine im Wesentlichen rechteckige Form aufweisen.
12. Compton-Kamera umfassend den Detektor nach Anspruch 11.
System zur Echtzeit-Visualisierung der Stoffwechselaktivität bei einem Probanden, dem zuvor ein radioaktiver Tracer verabreicht wurde, umfassend
einen Detektor nach einem der Ansprüche 1-12,
ein Computersystem zum Bestimmen des Aufpralls von Photonen in den vorderen (400) und hinteren (300) Detektoren, zum Berechnen einer Position eines Ursprungs des radioaktiven Zerfalls unter Verwendung von Antwortlinienberechnung und/oder Compton-Kegelberechnung in Verbindung mit dem Aufprall basierend auf Compton-Streuung vom vorderen Detektor und zum Erzeugen eines Videosignals, das den Ursprung des radioaktiven Zerfalls widerspiegelt, und
eine Vorrichtung (600), die in der Lage ist, das Videosignal zu empfangen und das Videosignal auf einem Bildschirm wiederzugeben.
14. System nach Anspruch 13, wobei das Computersystem ein digitaler Signalprozessor (500) ist, der mit den Schnittstellen der Detektormodule verbunden ist.