CN107544086B - Gamma photon detecting and imaging device and method - Google Patents
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Abstract
The invention provides a gamma photon detection device and a gamma photon detection method. The method comprises the steps of utilizing the plurality of detection probes to detect horse adding photons generated by a target object, enabling the horse adding photons to firstly act with a first layer of detector and then act with a second layer of detector, then obtaining an electric signal generated by each detection probe in at least one first time period, and respectively carrying out first time signal coincidence calculation on the electric signals generated by the plurality of scintillation crystal detectors of each detection probe so as to obtain a plurality of first time signal related data. Finally, the electrical signal generated by the first layer scintillation crystal detector of each detection probe in at least one second time period is obtained, and a second time signal coincidence calculation is carried out to obtain a plurality of pairs of second time signal related data. And performing activity distribution reconstruction by using the first and second time signal related data.
Description
Technical Field
The present invention relates to a gamma photon detection device and method, and more particularly, to a gamma photon detection imaging device and method for detecting transient gamma photons and positive photon mutual destruction gamma photons.
Background
As shown in fig. 1, Radiotherapy (RT) is a treatment method that irradiates a lesion with high-energy photons (such as X-ray) or charged particles to "kill malignant tumor" or "inhibit malignant tumor proliferation". Wherein, the curve 90 represents the relationship between the X-ray dose and the tissue depth, and because the physical characteristics of the X-ray and the substance action, the relative dose is gradually attenuated along with the depth along with the increase of the depth of the ray entering the tissue, so that more dose is released in the normal tissue before and after the tumor on the path of the single X-ray beam entering the body, which causes the normal tissue to be affected before the tumor cancer cells are destroyed by the X-ray, and even causes the normal tissue area after the tumor to have dose deposition. In contrast to the X-ray treatment, curve 91 represents the dose-depth relationship for charged particles, such as protons or heavy ions, since the mass loses less energy per unit distance as the proton travels faster, the kinetic energy initially released into the patient's body surface is relatively low, and up to a certain range depth, the kinetic energy of the protons or heavy ions decreases significantly, resulting in a maximum dose deposition, referred to as Bragg Peak 910(Bragg Peak), with almost no dose following the Bragg Peak.
Since proton or heavy ion therapy has the property that the dose is concentrated in the bragg peak, the uncertainty effect of the proton or heavy ion range is much more sensitive than X-ray. If the proton or heavy ion range prediction is uncertain, the depth of the bragg peak with highly concentrated dose is different from the expected depth, which may cause the risk that the high dose region does not cover the whole tumor or the high dose region covers the important organs around the tumor. Therefore, in clinical application, proton or heavy ion therapy must be able to accurately determine the range of proton or heavy ion range, and limit the bragg peak in the tumor region to achieve the best therapeutic effect.
However, there are still some factors that affect the inaccuracy of proton or heavy ion therapy, as follows:
(1) uncertainty caused by treatment planning: this is because the Hounsfield Unit (HU) in the current treatment plan mainly comes from X-ray CT image is the corresponding relationship between tissue and electron density, but the proton or heavy ion and biological tissue action is mainly considered the nucleus of tissue and its stopping power (stopping power), so there is uncertainty in converting the HU information of CT image into the stopping power of biological tissue. In addition, CT imaging artifacts, X-ray beam hardening (beamhardening) phenomena, etc. are also sources of uncertainty in currently evaluating proton or heavy ion range;
(2) the actual treatment location differs from the treatment plan: mainly comes from the movement of a treatment target caused by the positioning error of a patient, respiration or heartbeat and the like;
(3) tumor size changes during treatment, patient size changes, etc.; and
(4) errors from the facilities that generate the proton or heavy ion beam, such as energy degraders (energygraders), beam-forming (beam-delivery) related equipment, etc.
Due to the range inaccuracy, clinical treatment usually involves a large treatment safety margin (safety margin) in addition to the Clinical Target Volume (CTV) to be treated. The range of treatment, which increases due to range uncertainty, increases with tumor depth y (a 30cm depth tumor may increase by about 14mm of safe range of treatment). Increasing the safety margin by 1cm beyond CTV has the effect that normal healthy tissue of-80 cm3 receives the same radiation dose as the tumor, thus increasing the risk of damage to nearby normal tissue and carcinogenesis. The presence of critical organs with low radiation tolerance near the CTV will also affect the therapeutic dose delivery and the rate of malignancy control. In order to reduce the inaccuracy of the treatment process and improve the reliability of proton or heavy ion treatment, if the proton or heavy ion irradiation area of a patient can be monitored in the treatment process to confirm whether the path of a proton beam in the body of the patient needs to be corrected, the amplification of the treatment safety range can be reduced, the dosage administration is improved, and the tumor control rate and the treatment effect are further improved. Therefore, how to identify and monitor the path of the proton or heavy ion beam in the patient is a very important research issue.
In the prior art, there are several developing proton range verification techniques, and techniques that are consistent with real-time and non-invasive include proton transmission (proton transmission) imaging, prompt horse photon imaging, and pet (positron emission tomography) imaging. The proton penetration contrast takes additional proton equipment time and generates additional radiation dose, and the problems of multi-Coulomb scattering (multi-scattering) effect and the like are overcome, so that the short-term practical application chance is not high. Transient gamma Photon (PG) imaging and PET imaging are detected by using secondary particle-positron emission nuclei and transient gamma photons generated by the action of protons on diseased tissues. Both positron-destructive-horse photons and transient-horse photons can be detected through the body of a patient, and can be used for measuring and imaging the non-invasive nature and the range of photon emission of a living body (in vivo).
The proton and heavy ion are similar to the biological tissue mechanism of action and are not further distinguished from each other as represented by the proton therapy discussion below. Upon irradiation of protons, inelastic collisions of protons with tissue nuclei produce gammaphotons, which can be divided into two classes depending on the mechanism of induced generation. One is that when the proton and the nucleus of human tissue are in inelastic collision, the target nucleus is excited to an excited state, and a prompt gamma photon is generated in nanoseconds (ns). The other is that the target nuclei are fragmented by proton impact (nuclear fragmentation) to generate positive-emitting nuclei (plasma-emitting nuclei), such as O-15 (half-life: 2.037min), C-11 (half-life: 20.385min), N-13 (half-life: 9.965min), etc. (Newhausene and Zhang, 2015). Positron emission nuclear decay produces a positron (positron) that interdicts with electrons in tissue to produce a positron interdiction gammas pair of 511keV energy. The energy of the transient and positron mutual destruction photons can penetrate through the body of a patient to be detected and imaged, the activity distribution image is analyzed, and the path of the proton in the human body can be obtained, so that the method can be applied to range verification of proton or heavy ion treatment.
Positron mutual destruction and photon pair detection is generally performed by PET imaging, and is a mature and clinically applied in vivo proton range verification technique. However, the bio-metabolic (biological washout) effect of the positron emitting nuclei in the patient, the limited yield of the positron emitting nuclei, the short half-life, and the spatial alignment (co-registration) error between the treatment system and the PET image (caused by the movement of the patient between the treatment room and the PET imaging apparatus) limit the accuracy of the PET to proton range estimation.
As for the transient-gamma photons, the kinetic energy of protons required for generating the transient-gamma photons is lower than that required for generating the positron emission nuclei, so that the distribution position of the transient-gamma photons is consistent with the position of the maximum dose of proton deposition, and the accuracy of evaluating the bragg peak by utilizing the transient-gamma imaging activity is higher than that by utilizing a PET method. In addition, the instantaneous horse-adding photon yield is 60-80 times more than that of the positron mutual destruction horse-adding photon, which means that in the same detection geometrical condition theoretically, the instantaneous horse-adding photon radiography can have stronger signals, and the statistical error is smaller than that of PET. In the prior art, a collimator-based gamma camera (collimator-based camera) technology common in the field of nuclear medicine is used, and a slit collimator or an edge-knit collimator of a pinhole improved version is combined with a gamma detector to perform instantaneous gamma ray detection radiography. However, this technique has poor detection efficiency due to the collimator blocking most of the photons and the limited ability of the probe to detect high-energy prompt gamma photons. More importantly, due to the existence of the collimator, a large amount of neutrons (one of the secondary radiation particles induced by protons) are generated, and the neutrons will interfere with the signal detection of the detection probe, so that the signal-to-noise ratio (S/N) is reduced, and the definition of the active image and the accuracy of determining the proton range bragg peak by using the same are affected. In addition, because of the high energy of the photons of the prompt gamma radiation induced by radiotherapy, the collimator design is rather massive (up to 80cm half-height in size) and occupies a limited space in the treatment room. The above factors result in limited applicability of the collimator-type cameral to proton clinical therapy.
At this time, another combo scatter imaging technique, which is less common in the medical imaging field, becomes a new choice for studying prompt gamma photon radiography for proton or heavy ion therapy because it is suitable for high-energy photon detection imaging. The Kangpulu scattering imaging technology is mainly applied to astronomical cosmic rays, nuclear power plants, sensitive nuclear species and horse photon detection of homeland security and the like. Currently, some international proton or heavy ion treatment teams directly adopt previous detection designs to perform conpult scatterometry imaging, such as american tumor treatment Center (peg and Charles Stephenson Cancer Center) using a double-sided silicon strip type probe (DSSD) detector technology, korean hanyang university team using gamma electron cone top imaging (GEVI) detection technology, japan kyoto university team using electron tracking Compton camera (etc) detection technology, and american maryland university team cadmium zinc telluride (CdZnTe, CZT) detection technology. The team's detection probes either use ionization chamber (ionization chamber) technology or semiconductor detector technology. In particular, the semiconductor detector has quite high energy resolution (< 1%), and is suitable for conoton scattering imaging. However, for high energy plus horse photons, the detection efficiency is poor and the detector signal response time is long (μ s level). The long signal response time also lengthens dead time (dead time) of the detecting probe, so that the detecting counting efficiency of the detecting probe per unit time is poor. Such detectors can be used in non-medical applications because non-medical imaging can tolerate long data acquisitions of hours or even days to obtain adequate detection counts, but long imaging is not feasible if performed on a medical patient. Thus, to achieve sufficient detection counts, a relatively high proton dose is required to illuminate, and the count statistics for prompt plus motor imaging are acceptably inaccurate.
According to the practical experience of Dr.J.C.Polf, university of Maryland, USA, even if the CdZnTe detection probe with the highest detection efficiency is used for proton therapy prompt horse radiography, the required dosage is still more than 100 times higher than that of clinic. In order to make up for the lack of efficiency of the detector, there is also a design of tens of layers of detectors, but the overall size of the detector and the electronic system is considerable, which is not favorable for the application in the proton treatment room with limited space. In addition, since the detector signal response time is as long as microseconds (μ s), the time resolution of the scattering detector and the absorption detector is also on the order of microseconds. The application of microsecond-level time resolution contrast equipment in proton therapy prompt masam photon monitoring may result in more than 80% of detection counts being uncorrelated non-true coincident events (non-true coherent events), which may result in the position misjudgment of a prompt masam photon source. Because the traditional Compton scattering imaging probe is applied to the detection of forbidden radiation sources in astronomy and airport customs, the microsecond time resolution is enough to meet the non-medical requirement because the flux of cosmic rays reaching the earth is limited, the forbidden radiation sources are shielded and protected, and high-intensity background interference is avoided. In a word, the direct application of the existing DSSD and CZT probes designed for astronomy, homeland and nuclear energy safety requirements to the application of the transient gamma photon radiography of proton treatment has the defects of low detection efficiency, poor time resolution, overhigh rate of non-real time signal coincidence events and the like, and the defects of system signal-to-noise ratio, image definition and proton range Bragg peak estimation accuracy are influenced, so the Kangpulu scatter imaging detection probe for the proton treatment demand dosimeter has the necessity.
In addition, the prior art, such as U.S. Pat. No. 6,484,051, discloses a device capable of simultaneously detecting positron mutual destruction and transient gamma photons generated by isotopes, which requires at least three detectors to operate simultaneously due to the simultaneous measurement. In addition, the signal of the transient amazing photon is tens times of the number of the positron mutual destruction amazing photons, and if the signal is measured at the same time, the detection counting rate (counting rate) is dominated by the number of the positron mutual destruction amazing photons with low generation rate (yield rate), so that the signal-to-noise ratio of the detection signal is influenced, and the quality of image reconstruction is reduced. In addition, U.S. Pat. No. 9,069,089 discloses a technique for simultaneously detecting positron mutual destruction and transient amazing photons, in which a sensor for sensing transient amazing photons is disposed at the periphery of a positron mutual destruction and amazing photon sensor, and two different sensors are used for sensing amazing photons.
Disclosure of Invention
The invention provides a gamma photon detection device and a method, which utilize a rapid scintillation detector (for example, scintillation crystals such as LSO, LYSO, LaBr3 and the like are taken as materials), combine PET and Kangpu swallowing scattering detection technologies, manufacture a two-in-one (2-in-1) detection system with a function of measuring positron interdestructing photon pairs and instant gamma photons, and then utilize time sequence type data acquisition to carry out radiography and a PET embedded Kangpu swallowing imaging algorithm method to carry out image reconstruction and image fusion. The two-in-one detection system has the functions of measuring the radiography of the positron mutual destruction photon pair and the transient horse-adding photon, has double purposes and high convenience, and can improve the image quality and the verification accuracy of the proton/heavy ion range by integrating the information obtained by PET/PG dual-function imaging radiography.
The invention provides a horse-added photon detection device and a horse-added photon detection method, which adopt a shared probe and time sequence type data acquisition, have the advantages of saving cost and improving the signal-to-noise ratio of an image, and the reason is that the signal of the transient horse-added photon is dozens of times of the photon pair of the positron mutual destruction.
The invention provides a gamma photon detection device and method, which can be used for double isotopes (mixed positron isotope and gamma photon isotope) or special isotopes (isotope) capable of emitting positron and gamma photon, such as124I,86Y,89Zr, etc.) decay, transient and positron-mutually-destroyed pairs of positron-photons are detected. Because the above radiation source has correlation between the positron emission nuclei and the distribution of transient amalgams. Therefore, the PET activity distribution can be used as the prior information (prior information) for reconstructing the conplet image after being processed by dose correlation. In the image reconstruction process, the image reconstruction is carried out by using the correlation of the positron emission nucleus and the photon spatial distribution of the instant gammadia and a PET embedded Kangpu imaging algorithm method, so that the possible position distribution range of the instant gammadia source is reduced to the maximum possible probability distribution area from the original Kangpu pyramid, and the PG image signal-to-noise ratio can be improved and the reconstruction speed of the Kangpu image is accelerated by the method. The PET/PG dual-function imaging can take the advantages of PET and photon imaging of the instant gamma ray into account, and can further improve the accuracy of proton or heavy ion range verification and dosage administration or the quality of nuclear medicine double isotope or special isotope radiography.
The invention provides a gamma photon detection device, which uses a probe composed of a scintillation crystal. The scintillation detector has the characteristic of quick time response (time response), and the scintillation crystal with quick response (such as LSO, LYSO, LaBr3 and the like) is selected to match with the quick photoelectric sensor and the preamplification circuit, so that the response time and the time resolution of a detection signal of the system can reach nanosecond level. The invention detects the Gama photon pair generated by the mutual destruction of the positive photons through the same probe, excites the atomic nucleus to an excited state because of the inelastic collision of the proton and the tissue atomic nucleus, and generates a high-energy instantaneous Gama photon in nanosecond time. Compared with microsecond response of semiconductor detecting probes such as a bilateral silicon band type probe (DSSD) and a cadmium zinc tellurium type detecting probe (CdZnTe, CZT), the nanosecond-level scintillation detecting system has fast signal response time, a coincidence time window can be shortened by hundreds of times, the detecting counting rate can be improved, the influence of non-true coincidence detecting counting can be greatly reduced, and the proton range image definition and the Bragg peak estimation accuracy can be further improved.
The invention provides a gamma photon detection method, which utilizes a gamma photon detection device with a one-machine dual-function design to achieve the dual-radiography functions of positron emission nucleus distribution and instantaneous gamma photons by using the same probe. The scheme adopts a rapid scintillation detector as a substrate to assemble an imaging probe, and data acquisition of a Prompt Gamma Imaging (PGI) mode and a positron emission imaging (PAI) mode of two periods of time is realized, so that the two-in-one detection system has the functions of detecting a positron emission cross-destruction photon pair and a prompt gamma photon. With this system, in-room PET for proton/heavy ion therapy may also have prompt gamma photon detection imaging applications. During the proton/heavy ion beam treatment period (t 0-t 1), the detecting probe performs data acquisition in a PGI mode, and during the period from t 1-t 2 after the beam is stopped, the detecting probe performs data acquisition in a PAI mode. The obtained original data is subjected to PET image reconstruction and PET embedded image reconstruction to obtain a PET image and a PG image respectively. Compared with the prior art with single machine and single function and double functions, the single machine and double function design of the invention can have the functions of in-room PET and instantaneous horse photon distribution imaging, is economic and convenient, and improves the range evaluation accuracy by reconstructing and fusing PET/PG double images by PET embedded Kangpu images due to fully utilizing the spatial distribution relevance of instantaneous horse photons and positive photon mutual destruction and horse photons.
In one embodiment, the present invention provides a gamma photon detection device, which includes a plurality of detection probes and a signal processing circuit. Each detection probe is provided with a plurality of layers of scintillation crystal detectors which are arranged along the axial direction of the corresponding detection probe, the adjacent scintillation crystal detectors have a distance, and each detection probe is used for capturing gamma photons to generate corresponding electric signals. The signal processing circuit is electrically connected with the plurality of detection probes and is used for acquiring an electric signal generated by each detection probe in a first time period, performing first time signal coincidence calculation on the electric signal generated by the plurality of scintillation crystal detectors of each detection probe to acquire a plurality of first time signal related data corresponding to the plurality of detection probes, acquiring an electric signal generated by the first layer scintillation crystal detector of each detection probe in a second time period, and performing second time signal coincidence calculation to acquire a plurality of pairs of second time signal related data.
In another embodiment, the present invention provides a gamma photon detection method, which comprises the steps of providing a gamma photon detection device having a plurality of detection probes and a signal processing circuit electrically connected to the detection probes, wherein each detection probe has a plurality of layers of scintillation crystal detectors arranged along an axial direction of the corresponding detection probe, and each adjacent scintillation crystal detector has a distance. Then, the plurality of detecting probes detect the gamma photons generated by a target object. Then, the electrical signals generated by each detecting probe are independently captured in a first time period to form a plurality of first electrical signal groups. Then, the signal processing circuit is used for obtaining the electric signal generated by each detection probe in at least one first time period, and the first time signal coincidence calculation is carried out on the electric signals generated by a plurality of scintillation crystal detectors of each detection probe so as to obtain a plurality of first time signal coincidence related data corresponding to the plurality of detection probes, and the first time signal coincidence related data are stored; and capturing the electric signals generated by the first layer of scintillation crystal detectors of each detection probe in a second time period to form a plurality of second electric signal groups for storage. Then, the signal processing circuit is used for obtaining the electric signals generated by the first layer of scintillation crystal detector of each detection probe in at least one second time period, and a second time signal coincidence calculation is carried out to obtain a plurality of pairs of second time signal coincidence related data.
Drawings
FIG. 1 is a graph of dose intensity versus depth for charged particles and conventional photons;
FIG. 2A is a schematic diagram of an embodiment of a gamma photon detection device according to the present invention;
FIGS. 2B to 2D are schematic views of different embodiments of the detecting probe of the present invention;
FIGS. 2E to 2H are schematic views of different embodiments of the detecting probe and the arc track assembly of the present invention;
FIGS. 2I to 2K are schematic views of different embodiments of the detecting probe of the present invention;
FIG. 3A is a schematic flow chart of a horse-added photon detection method according to the present invention;
FIG. 3B is a schematic diagram of the process of the acquisition of the electrical signal of the detecting probe and the calculation of the time information;
FIG. 3C is a schematic flow chart of an embodiment of activity distribution reconstruction according to the present invention;
FIGS. 4A-4C are schematic diagrams of embodiments of detecting transient and positron mutual destruction photons using the Zebra photon detection apparatus of FIG. 2A;
FIG. 5 is a schematic diagram of acquiring a conplet scattering signal and a PET plus horse photon signal at different time periods according to the present invention;
FIG. 6 and FIG. 7 are schematic diagrams of different embodiments of time-domain synchronization algorithms;
FIG. 8A is a schematic view of a conpulson congolon;
FIG. 8B is a schematic diagram of a prior information reconstruction using a PET image-based dose distribution as a conplet image;
FIG. 9 is a schematic representation of the correlation between conpulson cones and PET image-based dose distribution intersections between multiple probes.
Description of reference numerals: 2-gamma photon detection means; 20a,20b,20c,20d,20 e-detection probe; 200a,200b,200c,200d,200e,200f,200g,200 h-scintillation crystal detector; 201-an array of scintillation crystals; 202-a photosensor; 203-a readout circuit; 204-analog and digital conversion circuitry; 205-an optical fiber; 21-a signal processing circuit; 22-a reconstruction unit; 220-PET reconstruction module; 221-conpu swallow reconstruction module; 222-first activity distribution information; 223-dose distribution information; 224-second activity distribution information; 23-a verification unit; 24-a charged particle generating means; 25-a bearing table; 26-an arc track; 3-gamma photon detection; 30-35-step; 340-343-step; 350-352-step; 8-a charged particle beam; 80. 80' -conpuloton cone; 81-activity distribution; 82a, 82 b-region; 9-a target; 90. 91-curve; 910 bragg peak; 92-reaction site.
Detailed Description
Please refer to fig. 2A, which is a schematic diagram of an embodiment of a gamma photon detection apparatus according to the present invention. In this embodiment, the gamma photon detecting device 2 comprisesA plurality of detecting probes 20a and 20b, a signal processing circuit 21, a reconstruction unit 22 and a verification unit 23. Each of the detecting probes 20a and 20b is a probe having nanosecond response grade, and the detecting surfaces thereof correspond to each other and are parallel to each other. Each of the detecting probes 20a and 20b has a plurality of layers of scintillation crystal detectors 200a to 200b and 200d to 200e arranged along the axial direction of the corresponding detecting probe, and adjacent scintillation crystal detectors 200a to 200b and 200d to 200e have a distance therebetween. In one embodiment, each of the detecting probes 20a and 20b is configured to capture the gamma photons generated by a charged particle beam along a traveling direction to generate a corresponding electrical signal. The material of each layer of the scintillation crystal detectors 200 a-200 b is a fast scintillation crystal such as lutetium orthosilicate (LSO), lutetium yttrium orthosilicate (LYSO), or lanthanum bromide (LaBr3), but not limited thereto. The charged particle beam is a proton beam or a heavy ion beam, and is generated by using a cyclotron (cyclotron) or a synchrotron (synchrotron) to generate a continuous or pulsed charged particle beam for irradiating a malignant tumor, and the tumor is destroyed by the maximum dose of its bragg peak. In addition, in another embodiment, the detection probes 20a and 20b can also be used in special isotopes (isotopes that emit both orthoions and gammadam photons, such as124I,86Y,89Zr, etc.) decay, detecting the resulting positron-destroying and positron-adding pairs. In one embodiment, the detection probes 20a and 20b are constructed as a conplet probe.
In one embodiment, as shown in FIG. 2B, each layer of scintillation crystal detectors 200 a-200B has a scintillation crystal array 201 connected to a photosensor 202 and readout circuitry 203 with preamplification circuitry, respectively, such that the response time and time resolution of the detection signals of the system can reach nanosecond level. In one embodiment, the scintillation crystal array 201 is composed of a plurality of scintillation crystals 201a, and the photosensor 202 is a multi-positive position-sensitive photomultiplier tube (PSPMT) design for converting the optical signals received by the scintillation crystal array 201 into analog electrical signals. Further, the photosensor 202 may be a photomultiplier tube (PMT) or a silicon photomultiplier (Si — PM) array. The readout circuit 203 is used to shunt, amplify and convert the current signal output by the photosensor 202 into an analog voltage signal.
In addition, as shown in FIG. 2C, it is a schematic diagram of another embodiment of the scintillator crystal detector. The scintillation crystal array 201 and the photosensor 202 are conducted through the optical fiber 205, and an included angle exists between the scintillation crystal array 201 and the photosensor 202. As shown in fig. 2D, which is a schematic diagram of a three-layer detector, the false signals caused by the optical fiber transmission can be reduced, i.e., extra compton scattered photons generated by the photomultiplier tube and scattered photons generated by the scintillation crystal react with the photomultiplier tube before reaching the next layer of detector. In addition, the distance between the detection layers can be reduced as much as possible, and the geometric efficiency is improved. Referring back to fig. 2A, the analog electrical signals output from the detecting probes 20a and 20b pass through an integrating and analog-to-digital circuit 204 for converting the output analog electrical signals into digital electrical signals. The signal processing circuit 21 is electrically connected to the integrating and analog-to-digital converting circuit 204 for receiving the digital electrical signal. In one embodiment, the signal processing circuit 21 has an fpga (field Programmable Gate array) Programmable circuit, which can calculate the received electrical signals by a program, such as: the timing is calculated, and then the output data of the circuit 21 is transmitted and stored in a computer hard disk or a memory to be used as a subsequent reconstruction unit 22. In this embodiment, the signal processing circuit 21 obtains the electrical signal generated by each detection probe in a first time period, performs a first time matching operation on the electrical signal generated by the multi-layer scintillation crystal detector of each detection probe to obtain a plurality of first time matching related data corresponding to the plurality of detection probes, obtains the electrical signal generated by the first-layer scintillation crystal detector of each detection probe in a second time period, and performs a second time matching operation on the electrical signal to obtain a plurality of pairs of second time matching related data. The first time-signal coincidence algorithm is a time-signal coincidence algorithm performed on signals obtained by each layer of the scintillation crystal detectors, and is used for judging whether the electric signals detected by each layer of the scintillation crystal detectors of each detection probe are from the same gamma photon, namely, a transient gamma photon. In this embodiment, taking the detection probe 20a as an example, after a certain gamma photon first acts on the first layer of the scintillation crystal detector 200a, the generated scattered photon then acts on the second layer of the scintillation crystal detector 200b, and the two have a slight time difference, so that the first time coincidence algorithm means that the time coincidence algorithm is performed on the signals obtained by each layer of the scintillation crystal detector 200a,200b or 200d,200e, and further, whether the electrical signals passing through each layer of the scintillation crystal detector 200a,200b belong to the same gamma photon event or not, and whether each layer of the scintillation crystal detector 200d,200e belongs to the same gamma photon event or not. The second time signal is calculated according to the algorithm, namely, the paired horse adding photons generated by mutual destruction of the positive ions and the electrons are respectively calculated with the electric signals generated by the action of the first layer of detectors of the opposite detection probes so as to judge whether the electric signals detected by the first layer of scintillation crystal detectors of each detection probe are from the same pair of horse adding photons, namely, the paired positive ions and the electrons are mutually destroyed to generate the horse adding photons. For example, in the present embodiment, the electrical signals generated by the pair of horse-added photons respectively acting on the first layer scintillation crystal detectors 200a and 200d of the pair of detection probes 20a and 20b can be determined whether the pair of horse-added photons 200a and 200d are the electrical signals generated by the pair of horse-added photons generated by the mutual destruction of the same group of positive photons and electrons after the second time signal coincidence operation.
The above-mentioned detecting probes are all combinations of top-illuminated (top-on) detectors, and in another embodiment, the detecting probes may also be combinations of side-illuminated (side-on) detectors and top-illuminated detectors. For example, as shown in fig. 2I, the detection probe 20c has two scintillation crystal detectors 200a and 200f, wherein the scintillation crystal detector 200a is a top-on detector, and the scintillation crystal detector 200f is a side-on detector. In addition, as shown in fig. 2J, the detecting probe 20d has two scintillation crystal detectors 200f and 200g, which are side-illuminated detectors. In addition, it is noted that the photosensors 202 at both ends of the side-illuminated detector can be disposed at both ends of the scintillation crystal, as shown in FIG. 2I, or at a single end, as shown in FIG. 2K, as may be desired by the user.
The reconstruction unit 22, which is disposed in a computer or a system with computing capability in one embodiment, includes a PET reconstruction module 220 and a compton reconstruction module 221 for retrieving the first and second time-signal related data from a storage device or a memory stored in a hard disk of the computer. The conpu reconstruction module 221 processes data from the PGI mode, which is generated by the probing probes 20a and 20b sensing transient amalgamated photons during the proton/heavy beam treatment period (t 0-t 1). The PET reconstruction 220 module processes data from pairs of 511keV energy Zebra photons acquired by the detection probes 20a and 20b during the time periods t 1-t 2 after the therapeutic beam is stopped in the PAI mode. The first activity distribution information 222 reconstructed by the PET reconstruction 220 module is further subjected to deconvolution (deconvo) processing to obtain a dose distribution information 223, and then the dose distribution information is transmitted to the compton reconstruction module 221, so that the compton reconstruction module 221 can reconstruct and obtain a second activity distribution information 224 about transient gammadam photons by using the first dose distribution as previous information, and further estimate the actual range of the charged particles in the living body. The PET/PG dual-function imaging can take the advantages of both PET and photon imaging of the instant gamma ray into account, and can further improve the accuracy of proton or heavy particle range verification and dose application. In an embodiment of radiation therapy application, the reconstruction unit 22 may be connected to a verification unit 23, after the verification unit 23 receives the activity distribution output by the reconstruction unit 22, the verification unit 23 establishes and displays an actual dose range distribution of the charged particles in the living body through activity distributions of PET and PG, so as to provide a medical practitioner to compare and confirm an actual range result and a predicted range of a treatment plan, and further use the difference as a basis for next charged particle irradiation parameter correction. In addition, in one embodiment, the gamma photon detecting and imaging device 2 further comprises a charged particle generating device 24 to form a gamma photon detecting device for therapy and simultaneous in vivo therapy monitoring, wherein the charged particle generating device 24 is used to generate charged particles required for therapy.
As shown in fig. 2E, in an embodiment, the detecting probes 20a and 20b are disposed on an arc-shaped track 26, and can perform displacement motion on the arc-shaped track 26, so that the detecting probe 20a or 20b can rotate around a center to adjust the detecting position, in this embodiment, the center of the arc-shaped track 26 is used as a rotation axis to rotate. It should be noted that, since the present invention has the functions of PET and PG detection imaging, the detecting probes 20a and 20b must be arranged at 180 degrees with respect to each other in the PET imaging mode, and the detecting probe 20a or 20b can have any included angle in the PG imaging mode. In another embodiment, as shown in fig. 2F, when the charged particle generating device 24 generates the charged particles 8, the detecting probes 20a and 20b have an included angle greater than 0 degree, which is about 90 degrees in this embodiment, so as to increase the uniformity of the three-directional resolution of the snapshot-amalgamated photon image. It should be noted that although the detecting probes in fig. 2A, 2E and 2F are a pair, the invention is not limited thereto, and they may be implemented according to the requirements of the application, for example: resolution, cost, etc. considerations, set the required number, for example: in fig. 2G, there are two detection modules, each having a plurality of detection probes 20a,20c,20e and 20b,20d,20 f. In addition, as shown in fig. 2H, each of the detecting probes 20a and 20b can also rotate around its own axis and move in a translational manner toward or away from the center of the circle.
Please refer to fig. 3A, which is a schematic flow chart of a gamma photon detection method according to the present invention. In this embodiment, the method 3 includes the following steps: first, step 30 is performed to determine a reaction site in a target according to a photon detection. In this step, one embodiment of the photon detection is a Computed Tomography (CT) scan performed for treatment planning, which is performed on an organism by mainly X-rays, such as: a human body, which is the target, is scanned, and the reaction site, in one embodiment, a tumor. The location of the tumor is determined by computed tomography. After the reaction site is found, step 31 is performed, in accordance with the condition of the organism, for example: the surrounding soft tissue distribution, and the response position information, are converted into an irradiation parameter of the charged particle beam, which includes irradiation energy, dose, and traveling direction and depth. Since the reaction sites are spatially distributed in three dimensions within the target, the energy required for the charged particle beam varies with the depth of travel of the beam onto the target. It should be noted that the traveling direction and depth of the charged particle beam are not limited to a single one, and a plurality of traveling directions and depths, that is, a plurality of directions and a plurality of range depths, may be irradiated onto the reaction site according to the treatment requirement.
The charged particles in this embodiment are protons, and since the initial dose released by the protons entering the body surface of the patient is relatively low, the kinetic energy of the protons is greatly reduced to a certain range depth, and then the maximum dose deposition is generated, and the dose released by this range depth is called bragg peak, and there is almost no dose behind the bragg peak (as shown in fig. 1). The energy or dose determined by the reaction position in step 31 is determined by utilizing the characteristic that the Bragg peak with the highest dose is located in the target position area through the regulation and control of proton energy and proton range, so that the normal tissue behind the reaction position can avoid the radiation irradiation risk.
After step 31, step 32 is performed, in which the gamma photon detection device 2 shown in fig. 4 is provided. In order to ensure that the dose or energy determined in step 31 can be concentrated at the position corresponding to the bragg peak, the gammadam photon detection device 2 provided by the present invention can be used for real-time synchronous in-vivo monitoring during the treatment process, so as to ensure the treatment efficiency and reduce the error, thereby protecting the normal tissue from the radiation damage. As shown in FIG. 4A, the target 9 is a living body, and lies on the carrier 25, and the target 9 has a reaction site 92 therein, which is a tumor in this embodiment. The detecting probes 20 are disposed around the periphery of the target 9, and although there are two detecting probes 20a and 20b in the embodiment, the number of the detecting probes can be determined according to the requirement. Then, step 33 is performed to enable the charged particle generating device 24 to generate a charged particle beam 8. During treatment, according to the irradiation information generated in step 31, the projected charged particle beam 8 is projected toward the target 9 and enters the target, and after inelastic collision between protons and tissue nuclei, two types of addam photons are mainly generated, wherein one type of addam photons is generated by mutual destruction of positrons generated by decay of secondary particles, namely, positron emission nuclei, generated by interaction between protons and diseased tissue and electrons in the tissue, and a pair of addam photons with 511keV energy is generated. The other is that when the proton and the atomic nucleus of the human tissue generate inelastic collision, the target nucleus is excited to an excited state, and a transient amazing photon is generated in nanosecond time.
To detect these two gammaonic photons, an extraction electrical signal is detected for a first time period and a second time period, via step 34. For the present embodiment of generating charged particles, the first time period represents a time period during which the charged particles are turned on, and the second time period represents a time period during which the charged particles are turned off. As shown in fig. 4A, during the first time period when the charged particle beam 8 is activated, each layer of the scintillation crystal detectors 200a and 200b and 200d and 200e of each of the detection probes 20a and 20b will detect the incoming signal of the transient hamama photon, taking the detection probe 20a as an example, that is, after a certain hamama photon first acts on the first layer of the scintillation crystal detector 200a, the generated scattered photon then acts on the second layer of the scintillation crystal detector 200b, and there is a slight time difference between the two, and similarly, the detection probe 20b also does. On the contrary, as shown in fig. 4B, during the second time period when the charged particle beam 8 is turned off, the first layer scintillation crystal detectors 200a and 200d of each of the detection probes 20a and 20B detect the incident signals of the pairs of gamma photons generated by the mutual destruction of the positive photons, i.e., the pairs of gamma photons generated by the mutual destruction of the positive photons and the electrons respectively interact with the first layer scintillation crystal detectors 200a and 200d of the opposite detection probes 20a and 20B. The main reason for segmented detection is that transient amazing photons are generated only during the period when charged particles are started, and the existence time is short, whereas positive destruction amazing photons have long reaction time, so that the characteristics of the two kinds of amazing photons are different. In addition, it has clinical application maturity to measure positive mutual destruction and horse photon pair and estimate charged particle range with PET activity image, separately measures instantaneous horse photon signal and positive mutual destruction and horse photon pair, can promote the accuracy that instantaneous horse photon image rebuild through more ripe PET technique, and the signal of instantaneous horse photon is many tens times of positive mutual destruction and horse photon, and instantaneous horse photon radiography can promote the signal-to-noise ratio of image, and then strengthens the precision that charged particle range verified the application. It should be noted that although the positive sub-pair of fig. 4B is sensed by the first layer of scintillation crystal detectors 200a and 200d, in another embodiment, as shown in fig. 4C, sensing may also be performed by the second layer of scintillation crystal detectors 200B and 200e based on the positive sub-pair.
Referring to fig. 3B, in step 34, a step 340 of independently capturing the electrical signals generated by each of the detecting probes 20a and 20B in a first time period forms two first electrical signal sets corresponding to the detecting probes 20a and 20B. In this step, taking the detecting probe 20a as an example, the signal processing circuit 21 is mainly used to record the corresponding gamma photons E0A first position for Compton scattering (Compton scattering) interaction with the first layer of scintillation crystal detectors 200a, a first time and a first deposition energy Δ E, and a scattered and stimulated photon energy E1Second position, second deposition energy Δ E for interaction with second layer of scintillation crystal detector 200b2And the second time to form the first electric signal group. Ideally the second layer scintillation crystal detector 200b completely absorbs the scattered plus horse photon energy E1So that the second deposition energy Δ E2＝E1. Then, in step 341, the first electrical signal sets generated by each detection probe in the first time period in step 340 are obtained, and a first time signal coincidence algorithm is performed on the electrical signals generated by the scintillation crystal detectors of each detection probe to filter non-true non-correlated signals, so as to obtain first time signal related data of the transient gamma photons corresponding to the detection probes. Step 340 is illustrated in fig. 6, which is a schematic diagram of the first time-domain signal matching algorithm. Taking the detection probe 20a as an example, the first to second layers of scintillation crystal detectors 200a to 200b provided therein are respectively used to extract signals corresponding to detected events, and fig. 6 shows three events I1, I2, and I3. Each event has two electrical signals, one for each layer of the scintillation crystal detectors 200 a-200 b, each set of electrical signals containing position, time, and energy information. Comparing the first and second time signals, if the first and second time signals fall within the predetermined time rangeAnd in the time window, the electric signal measured by the second layer of scintillation crystal detector comes from the gamma photon acted with the first layer of scintillation crystal detector with a large chance. In the present embodiment, there are three time intervals W1, W2 and W3 with the same time window width, and it is noted that the size of the time window can be determined according to the requirement. As can be seen from fig. 6, in the period W2, only the scintillation crystal detector 200a has a signal, the scintillation crystal detector 200b does not have a signal, and in the periods W1 and W3, the scintillation crystal detectors 200a and 200b have signals simultaneously, so that the signals extracted in the periods W1 and W3 belong to signals of time-series coincidence, that is, it is determined whether the signals are the same group of incident evanescence photons and the electric signals generated by the evanescence photons. If so, the record is stored, otherwise the record is not stored.
Then, in step 342, the electrical signals generated by the first layer of scintillation crystal detectors 200a and 200d of each of the detection probes 20a and 20b are captured during the second time period to form a plurality of second electrical signal sets. In step 342, the signal processing circuit 21 is mainly used to process and digitize the horse-added photons at the first position, the first time and the first deposition energy of the detection probe 20a representing the action of the first layer of the scintillation crystal detector 200a, and the second position, the second deposition energy and the second time of the detection probe 20b representing the action of the horse-added photons at the first layer of the scintillation crystal detector 200d, so as to form the second electrical signal group. As shown in fig. 5, during the detecting period from t0 to t2, during the first time period from t0 to t1 of the detecting period, a plurality of electrical signal sets a1 of the detecting probe 20a and a plurality of electrical signal sets B1 of the detecting probe 20B are obtained in step 340, and similarly during the second time period from t1 to t2, a plurality of electrical signals C1 of the first layer of the scintillation crystal detectors 200a and 200d of each of the detecting probes 20a and 20B are obtained by executing step 341. According to the design of the existing nuclear medicine PET scintillation detector, the detector detects circuit signals, so that the effect of the gamma photon on which scintillation crystal can be known, the effect position and the occurrence time can be further known, the time for detecting the signals can be recorded, and the energy deposited on the detector can also be known from the corresponding electric signal intensity. Thus, each of the signals Pa 1-Pan, Pb 1-Pbn, Pd 1-Pdn and Pe 1-Pen contains information on the position, time and deposition energy.
Then, step 343 is performed to obtain a plurality of second electrical signal sets generated by the first layer of scintillation crystal detectors of each detection probe in the second time period in step 342, and a second time signal coincidence algorithm is performed to filter the non-true non-correlated signals to obtain a plurality of pairs of second time signal correlated data. In this step, the object of the timing coincidence algorithm is the first layer of the scintillation crystal detector of each detection probe, as exemplified in fig. 4A, i.e., the scintillation crystal detectors 200a and 200d of the detection probes 20a and 20 b. Referring to FIG. 7, in FIG. 7, a plurality of second electrical signal sets corresponding to three events I4-I6 generated by the scintillation crystal detectors 200a and 200d are shown, and in the present embodiment, whether these signals belong to a pair of electrical signals generated by a pair of 511keV addends generated by mutual destruction of the same positron is determined by three time intervals W4-W6 of the same time window width and energy. If so, the record is stored, otherwise the record is not stored. It should be noted that, in another embodiment, steps 340-341 and steps 342-343 can be performed independently and synchronously.
Referring back to fig. 3A, after step 34, the reconstruction unit 22 shown in fig. 4A may be further utilized to perform step 35 activity distribution reconstruction, and further convert the result of the final activity distribution into an image. The reconstruction of the step is based on the fact that PET activity distribution is used as the previous information of the conplet image, and then the reconstruction of the conplet image is carried out. Because proton or heavy ion range pathways in the patient's body generate photons of instant gammadam in addition to the positron emitting nuclei, the positron emitting nuclei are related to the proton or heavy ion dose distribution and also have a correlation with the photon distribution of instant gammadam. Therefore, the PET activity distribution as the basis can be used as the prior information for reconstructing the conplet image after the dose correlation treatment. As shown in FIG. 3C, step 350 is performed first, and the second time signal coincidence data obtained in step 343 is pre-processed, including the calibration of the detector, the correlation between the position signal outputted from the detector and the crystal position, and the correlation is converted into the position and energy dataPET reconstruction is performed to obtain a first activity distribution for positron interdestructions plus photons. The activity distribution of the positron mutual destruction and gamma photons is matched with the distribution state of proton or heavy ion range approaches in the patient body, which is approximately in a linear beam in space and represents the incident range condition of the charged particle beam obtained according to the positron mutual destruction and gamma photons. Step 351 is then performed to reconvert the first activity distribution of positron interdestructions plus photons from step 350 into dose distribution information obtained from PET estimation. Finally, step 352 is performed to correlate the dose distribution information of the positron mutual destruction and horse photons obtained in step 351, that is, the dose distribution information of step 351 is used as the previous information of step 352 to perform the PET embedded type compton image reconstruction, that is, the compton conical surface and the dose distribution obtained by the activity distribution conversion estimation of the positron mutual destruction and horse photons in step 351 are compared to obtain
And performing intersection to further obtain a second activity distribution.
The following description uses the principle of step 352, as shown in fig. 8A and 8B, where fig. 8A is an electrical signal that would simply represent the location of action and the deposited energy of the prompt masa photon and the detector, for example, if the single event is obtained in step 343, for example: pa1 and Pb1 in fig. 6, for example: the method comprises the steps of correcting a detector, correlating a position signal output by the detector with the position of a crystal, and reconstructing to obtain an activity distribution map. According to the conplet scattering principle, the possible origin positions of the gamma photons can be estimated from the gamma photons incident on the scintillation detector and the scattering angle between the scattered gamma photons. The intermediate calculation can be used to match the incident and scattered photons by the time-domain coincidence calculation. The gamma photons correlated by the algorithm from the first and second layer scintillation crystal detectors in step 341 are then detected by the detection algorithm, and the likely region of origin of the gamma photon is known from the detected energy and equation (1).
where E0＝E1+ΔE1
Wherein the original energy is E0The high-energy gamma photon and the first layer of scintillation detectors 200a and 200d generate conplet scattering effect to release partial energy delta E1After the first layer of scintillation detectors is E1The conpulson of the energy scatters photons. For a single snapshot gammagram ray measured, the possible locations of this origin point are distributed over the surface of a conPont conus (Comptone cone) 80. However, since the conpul cone 80 only knows the action position (vertex) and the scattering angle θ, any position on the cone surface may be the position of the target, and thus the uncertainty is not favorable for monitoring the treatment range by only using the conpul cone to infer the position of the target. In order to solve this problem, a method combining step 351 and step 352 is adopted, that is, the range approach of protons or heavy ions in the body of the patient is utilized to generate secondary particles, namely, positron emission nuclei and transient gammadam photons, so that the characteristics of both the positron emission nuclei and the transient gammadam photon distribution, which are related to the ranges and the doses of protons/heavy ions, are converted into dose distributions after dose correlation processing based on PET activity distributions, and the dose distributions are used as the previous information of conpul imaging to reconstruct the activity distributions. In this embodiment, the positron emission nuclear activity distribution is formed by the PET image after step 351
Converted into dose distribution
The conversion method may convert the positive electron emission nuclear activity distribution into a dose distribution (s.remmele et al.2011phys.med.biol.) by using, for example, a filter function inverse convolution method, but is not limited thereto. As shown in FIG. 8B, the line beam labeled 81 represents the spatial distribution of the dose estimated from the PET image
It He kangThe intersection of the conus surfaces 80 forms a pair of regions 82a and 82 b. The most likely location of the instant gamma photon origin is then reduced from the surface of the conplet cone 80 to the areas 82a and 82b, where the most likely location of the instant gamma photon origin is
Can be represented by the following formula,
thus, at step 352, after the intersection, the distribution of possible locations of origin of the single prompt masa photons may be narrowed from the original conpu conus 80 surface to the regions 82a and 82 b. In the PET embedded type Kangpu swallow image reconstruction process, the possible position distribution range of the radiation source is reduced, so that the signal-to-noise ratio quality of the reconstructed image can be greatly improved, the reconstruction convergence speed of the transient-Canma photon image can be accelerated, and more importantly, the accuracy of range evaluation is improved.
Fig. 8B shows the result of intersection between the distribution of the second temporal correlation data of step 343 after the PET image reconstruction and dose conversion of steps 350-351 and the activity distribution of the first temporal correlation data of step 341 formed in step 352. Since the embodiment of fig. 4A has two detection probes 20a and 20b, the activity distribution intersection shown in fig. 9 is formed after step 353 for each detection probe 20a and 20 b. Through a PET architecture formed by the conpu cones 80 and 80' of the detection probe 20a and the detection probe 20b and the first layer scintillation crystal detectors 200a and 200d of the detection probe 20a and the detection probe 20b, the intersection of the obtained activity distribution and the distribution 81 after dose conversion can reduce the possible positions of gamma photons to the intersection region, thereby improving the accuracy of subsequent range evaluation.
For example, tumor treatment is performed because treatment plan evaluation is performed before treatment, and the position of the tumor, the angle at which the charged particle beam should be irradiated, the dose applied, and the like are evaluated. In order to confirm the difference between the actual range of the charged particle beam in the living body and the position of the generated Bragg peak and the estimated range of the treatment plan in the treatment process, the device and the method can obtain the actual range of the charged particle beam in real time in each treatment, and the user can verify and compare the difference between the estimated range of the treatment plan and the actual range determined by the activity distribution generated in the step 352 by providing displayed information through the verification unit 23, so that the difference can be used as a basis for adjusting the irradiation parameters of the charged particle beam in the subsequent treatment, the double effects of monitoring and adjusting the recommended living body treatment can be achieved, the effectiveness of the treatment is improved, and the side effect of damaging healthy tissues due to the uncertain range is reduced.
While the foregoing embodiments are examples with two layers of scintillation crystals, in another embodiment, there may be three layers of scintillation crystals, such as the structure shown in fig. 2D, wherein the third layer of scintillation crystal detectors is arranged to consider the two-layer scintillation crystal detector design of conpult scatter imaging being too ideal, because the second layer of detectors (also called absorption detectors) must compensate all the energy of the scattered photons to obtain the initial energy E of the hamaran photons required by equation (1) or equation (2)0. However, the proton treatment induced transient gamma photons have a wide energy distribution range and high energy, and the energy is usually several keV to 10MeV, even 15 MeV. Thus, the initial Zeeman photon energy at actual detection is unknown, while the scattered Zeeman photon energy E1The probability of being totally blocked by the absorption detector is limited, want to pass E1And energy Δ E deposited on the first layer detector1The sum of (a) to obtain the initial gamma photon total energy E0The chance of (a) is not high. To enhance this part, in this embodiment, the detection probe is replaced by a three-layer scintillation detector, and the third layer detector only needs to measure the action position of the second scattered photon but does not need to record the deposition energy of the third layer detector. Theta2The calculation can be obtained by the positions of the incident gamma photons and the scattered gamma photons recorded by the three scintillation crystal detectors respectively, and then all known energy records are substituted to calculate the conpulson cone angle theta of the possible position of the incident photon source (namely the initial position of the transient gamma photons generated by the action of the protons and the tissues)1As shown in equation (3). As for the step361, the most probable location of the prompt gamma photon origin in the case of the three-layer scintillation detector
Again calculated by equation (2).
E0＝E1+ΔE1
E1＝E2+ΔE2
In addition, it should be noted that, in the above embodiments, the transient and the pair of positron destructions are detected by the transient and the positron emission photons generated by the charged particle beam. In another embodiment, the detection device of the present invention can also be applied to dual isotopes (mixed positron isotope and gamma photon isotope) or special isotopes (isotope capable of emitting both positron and gamma photon, such as124I,86Y,89Zr, etc.) and the like. Therefore, the detection of the gamma photons generated by the charged particle beam is not a limitation. In the present embodiment, similarly, the first time period detects the gammadia photons, and the second time period detects the gammadia photons in which the positive photons are mutually destroyed, so that the first time period and the second time period are alternately performed, and the obtained electrical signals are processed according to the steps shown in fig. 3B and fig. 3C, so as to reconstruct the corresponding activity distribution and image.
The above description is only for the purpose of describing preferred embodiments or examples of the present invention by means of solving the problems, and is not intended to limit the scope of the present invention. The scope of the invention is to be determined by the following claims and their equivalents.
Claims (19)
1. A horse-added photon detection device, comprising:
the detection probes are provided with a plurality of layers of scintillation crystal detectors, each layer of scintillation crystal detector is arranged along the axial direction of the corresponding detection probe, the scintillation crystal detectors of adjacent layers have a distance, each detection probe is used for capturing the gamma photons to generate a corresponding electric signal, each layer of scintillation crystal detector of each detection probe detects the electric signal generated by the incidence of the transient gamma photons in a first time period, and the first layer of scintillation crystal detector of each detection probe detects the signals of the incidence of the paired positive photons mutually destroying the gamma photons in a second time period;
a signal processing circuit electrically connected to the plurality of detecting probes for obtaining an electrical signal generated by each detecting probe during the first time period, and performing a first time signal matching algorithm on the electrical signal generated by the multi-layer scintillation crystal detector of each detection probe, so as to determine whether the electrical signals detected by the scintillation crystal detectors of each layer of each detection probe are from the same transient gamma photon, so as to obtain a plurality of first time-related data corresponding to a plurality of detection probes, and obtaining the electrical signal generated by the first layer scintillation crystal detector of each detection probe in a second time period, and performing a second time signal matching algorithm, so as to obtain a plurality of pairs of second time signal related data to judge whether the electric signal detected by the first layer scintillation crystal detector of each detection probe is from the same pair of positive mutual destruction and horse photon; and
and the reconstruction unit comprises a PET reconstruction module and a Compton reconstruction module, wherein the PET reconstruction module is used for reconstructing a first activity distribution according to the plurality of pairs of second time signal related data, then the first activity distribution is converted to obtain a dose distribution, and the dose distribution is transmitted to the Compton reconstruction module, so that the Compton reconstruction module uses the dose distribution as previous information, associates the dose distribution with the plurality of first time signal related data to reconstruct a second activity distribution, and further estimates the actual range of the charged particles in the organism.
2. The apparatus of claim 1 wherein the prompt or positron-destroying amalgamated photon is generated by a decay of at least one isotope.
3. The gamma photon detection device of claim 1 further comprising a charged particle generation device for generating a charged particle beam that generates the transient or positron destructing gamma photons when interacting with a target along a direction of travel.
4. The apparatus of claim 3, wherein the first time period is a time period when the charged particles are turned on, and the second time period is a time period when the charged particles are turned off.
5. The apparatus of claim 3, wherein the charged particles are protons or heavy ions.
6. The gamma photon detection device of claim 1 wherein said plurality of probes are independently rotatable about a central position or relatively movable to vary the detection position.
7. The equanized photon detector device of claim 1, wherein each layer of scintillation crystal detectors further comprises:
a scintillation crystal array;
a photoelectric sensor coupled to the scintillation crystal for converting the light signal generated by the scintillation crystal array into an electrical signal; and
and the reading circuit comprises a preamplifier circuit which is electrically connected with the photoelectric sensor.
8. The gamma photon detection device of claim 7 wherein said scintillation crystal array is coupled to said photosensor by an optical fiber.
9. The gamma photon detection device of claim 7 wherein the scintillation crystal array is a side-illuminated scintillation crystal array or a front-illuminated scintillation crystal array.
10. A gamma photon detection method, comprising the steps of:
providing a horse-added photon detection device which is provided with a plurality of detection probes and a signal processing circuit electrically connected with the detection probes, wherein each detection probe is provided with a plurality of layers of scintillation crystal detectors, each layer of scintillation crystal detectors are arranged along the axial direction of the corresponding detection probe, and the scintillation crystal detectors of adjacent layers have a distance;
the detection probes detect transient or positive mutual destruction and horse-added photons generated by a target object;
capturing electric signals of the detection probes in a period, wherein the period consists of a first time period and a second time period, each layer of scintillation crystal detector of each detection probe detects electric signals generated by instantaneous mason photon incidence in the first time period, and the first layer of scintillation crystal detector of each detection probe detects paired positive photons mutually destroy mason photon incidence signals in the second time period;
acquiring an electric signal generated by each detection probe in the first time period, and performing first time signal coincidence calculation on the electric signal generated by the multi-layer scintillation crystal detector of each detection probe to obtain a plurality of first time signal related data corresponding to the plurality of detection probes;
obtaining the electrical signal generated by the first layer scintillation crystal detector of each detection probe in the second time period, and performing a second time signal coincidence calculation to obtain a plurality of pairs of second time signal related data; and
providing a reconstruction unit which comprises a PET reconstruction module and a Compton reconstruction module, wherein the PET reconstruction module is used for reconstructing a first activity distribution according to the plurality of pairs of second time-signal related data, performing conversion on the first activity distribution to obtain a dose distribution, and transmitting the dose distribution to the Compton reconstruction module, so that the Compton reconstruction module uses the dose distribution as previous information, associates the dose distribution with the plurality of first time-signal related data to reconstruct a second activity distribution, and further estimates the actual range of the charged particles in the organism.
11. The method for detecting hamaran photons of claim 10, wherein the prompt or positron-destruction hamaran photons are generated by a decay of at least one isotope.
12. The method of claim 10, wherein the transient or positron-destroying amalgamated photons are generated by a charged particle beam interacting with a target along a direction of travel, an irradiation parameter of the charged particle beam corresponding to a maximum range evaluation position.
13. The method of claim 12, wherein the first time period is a time period when the charged particle is turned on, and the second time period is a time period when the charged particle is turned off.
14. The gamma photon detection method of claim 12, wherein said charged particles are protons or heavy ions.
15. The method of claim 10, wherein the extracting the electrical signal during the first time period further comprises:
the multi-layer scintillation crystal detector in each detection probe respectively detects two events related to transient gamma photons and scattered gamma photons; and
the electric signals generated by the multiple layers of scintillation crystal detectors in the same detection probe are processed to obtain a plurality of first digital electric signals related to the multiple layers of scintillation crystal detectors.
16. The gamma photon detection method of claim 15 wherein said first time alignment algorithm further comprises the steps of:
obtaining a plurality of first digital electrical signals related to the multi-layer scintillation crystal detector; and
and comparing a plurality of first digital electric signals respectively generated by the multiple layers of scintillation crystal detectors in each detection probe in sequence by using a time window, and if the time difference of the first digital signals of the scintillation crystal detectors on different layers is within a set coincidence time window, indicating that the first digital signals measured by the scintillation crystal detectors on different layers come from transient horse photons acted by the scintillation crystal detectors on the first layer.
17. The method of claim 10, wherein the extracting the electrical signal during the second time period further comprises:
a first layer of scintillation crystal detector in each detection probe respectively detects a positive-neutron mutual destruction and gamma photon event; and
the electrical signals generated by the first layer of scintillation crystal detectors in each detection probe are processed to obtain a plurality of second digital electrical signals related to each first layer of scintillation crystal detector.
18. The method of claim 17, wherein the second time-aligned algorithm further comprises the steps of:
obtaining a plurality of second digital electrical signals of the first layer of scintillation crystal detectors of different detection probes; and
and comparing a plurality of second digital electric signals respectively generated by the first layer of scintillation crystal detectors in each detection probe in sequence by using a time window, and if the time difference of the second digital electric signals respectively generated by the first layer of scintillation crystal detectors in different detection probes is in a set coincidence time window and the energy of the second digital electric signals in the time window is in a preset energy range, indicating that the second digital electric signals measured by each first layer of scintillation crystal detectors are from the same pair of positron mutual destruction and horse addition photons.
19. The gamma photon detection method of claim 10 wherein each scintillation crystal detector further comprises:
a scintillation crystal array;
a photoelectric sensor coupled to the scintillation crystal for converting the optical signal generated by the scintillation crystal array into an electrical signal; and
and the reading circuit comprises a pre-amplifying circuit which is electrically connected with the photoelectric sensor, wherein the scintillation crystal array is a side-illuminated scintillation crystal array or a front-illuminated scintillation crystal array.
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| US10908086B2 (en) | 2018-12-26 | 2021-02-02 | Industrial Technology Research Institute | Signal processing system and method thereof |
| CN110477942B (en) * | 2019-08-20 | 2023-06-23 | 上海联影医疗科技股份有限公司 | PET detector and medical imaging equipment |
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