Method for PET detector afterglow management

Abstract

Disclosed herein are methods and apparatus for acquiring positron emission (or PET) data in the presence of ionizing radiation that causes PET detector afterglow. In one variation, the method includes adjusting a coincidence trigger threshold of a PET detector during a treatment procedure. In one variation, the method includes adjusting a gain factor (e.g., a gain factor for multiplying and / or shifting output (s) of PET detector (s)) used in positron emission data acquisition during a treatment procedure. In some variations, a method for acquiring positron emission data during a radiotherapy procedure includes suspending communication between a PET detector and a signal processor of a controller for a predetermined period of time after emission of a radiation pulse by a linear accelerator.

Description

[0001] This application is a divisional application of Chinese Patent Application No. 202210276868.9, which in turn is a divisional application of Chinese Patent Application No. 201880057036.X. The original parent application was filed on July 11, 2018, with a priority date of July 11, 2017. The international application number is PCT / US2018 / 041700, and the date of entry into the Chinese national phase was March 3, 2020. The invention is entitled "Method for Afterglow Management of PET Detector". Cross-reference to related applications

[0002] This application claims priority to U.S. Provisional Patent Application No. 62 / 531,260, filed July 11, 2017, the entire contents of which are incorporated herein by reference. Technical Field

[0003] This disclosure relates to methods used in radiotherapy systems that include a linear accelerator (or other ionizing radiation source) and one or more positron emission tomography (PET) detectors. Background Technology

[0004] Radiation therapy systems typically have a radiation source (e.g., a linear accelerator or linear accelerator) that generates a therapeutic beam of radiation to target a tissue region, such as a patient's tumor area. Although the generated beam can be directed to the target region and can be beam-limited by one or more jaws and / or collimators, a portion of the beam may deviate from and / or scatter from the target region. This scattered radiation may interfere with the function of other components of the radiation therapy system.

[0005] For example, scattered or stray radiation can affect the ability of various detectors in a radiotherapy system, such as X-ray and / or PET detectors, to accurately acquire data. PET detectors in a radiotherapy system may be affected, making their response to scattered or stray radiation indistinguishable from genuine positron emission events. In cases of high-level radiation (e.g., during radiation pulses from a linear accelerator), PET detectors may "blind" and / or saturate. This can render them unable to meaningfully detect positron emission data.

[0006] Therefore, it may be desirable to develop methods and devices for managing the risk of equipment damage and / or data corruption due to scattered radiation from linear accelerators. Summary of the Invention

[0007] This document discloses a method and apparatus for acquiring positron emission (or PET) data in the presence of ionizing radiation that causes afterglow in a PET detector. In one variation, the method may include adjusting a re-activation threshold of the PET detector during a treatment session. The re-activation threshold may increase with increasing afterglow of the PET detector. For example, the re-activation threshold may increase with increasing and / or exceeding a threshold dark count rate of one or more PET detectors. Alternatively or additionally, the re-activation threshold may increase with increasing and / or exceeding a threshold bias current level of one or more PET detectors. The re-activation threshold may also be adjusted based on a measured temperature of the system (e.g., at or near the PET detector), wherein the re-activation threshold may increase with increasing system temperature. In some variations, the re-activation threshold may be adjusted based on the radiation output of a radiation source or linear accelerator. For example, the re-activation threshold may be adjusted when the number of emitted radiation pulses exceeds a predetermined threshold, and / or based on a pulse schedule, and / or based on the cumulative amount of radiation emitted by the linear accelerator during the treatment session. In some variations, the overlap trigger threshold can be adjusted if the synchronization offset between two system components (e.g., a linear accelerator and a collimator) exceeds a predetermined threshold.

[0008] In some variations, a method for acquiring positron emission data during a radiotherapy session may include pausing communication between the PET detector and the controller's signal processor for a predetermined time period after the linear accelerator has emitted a radiation pulse. For example, the predetermined time period may be about 100 μs or longer, or about 200 μs or longer. Alternatively or additionally, the predetermined time period may be determined at least in part by the width or duration of the linear accelerator radiation pulse. For example, the predetermined time period may be about 25 times or about 100 times longer than the duration of the linear accelerator pulse. After the predetermined time period has elapsed, communication between the PET detector and the signal processor may resume, and the positron emission data may be transferred from the detector to the signal processor and / or acquired by the signal processor for analysis and / or storage by the controller.

[0009] In other variations, the radiotherapy system may include a radiation source, multiple PET detectors (e.g., a PET detector array), and a radiation shield that can move above the multiple PET detectors. The radiation shield can be positioned above the PET detectors during the radiation interval when the radiation source is emitting radiation, and can be positioned away from the PET detectors during the detection interval when the radiation source is not emitting radiation.

[0010] A variant of the radiotherapy system may include: a radiation source configured to direct one or more radiation pulses toward a PET-avid region of interest, wherein each radiation pulse has a predetermined pulse duration; a plurality of PET detectors configured to detect positron emission paths by detecting positron annihilation photon pairs incident on a portion of the detectors within an overlap time window and generating a detector signal exceeding an overlap triggering threshold; and a controller communicating with the plurality of PET detectors, wherein the controller is configured to adjust the overlap triggering threshold during the treatment phase. The controller may be configured to adjust the overlap triggering threshold after a threshold number of radiation pulses has been directed toward the region of interest. The threshold number of radiation pulses may be approximately 1000 radiation pulses. The overlap triggering threshold may be from approximately two photon triggering to approximately five photon triggering. The overlap triggering threshold may be a first overlap triggering threshold, and the threshold number of radiation pulses may be a first threshold number of radiation pulses, and the controller may be configured to adjust the first overlap triggering threshold to a second overlap triggering threshold after a second threshold number of radiation pulses has been directed toward the region of interest. The second re-coinciding trigger threshold may be greater than the first re-coinciding trigger threshold, and the number of second threshold pulses may be greater than the number of first threshold pulses. The second re-coinciding trigger threshold may range from approximately four photon triggering to approximately six photon triggering, and the number of second threshold pulses may be approximately 2000. Alternatively, the second re-coinciding trigger threshold may be less than the first re-coinciding trigger threshold, and the number of second threshold pulses may be greater than the number of first threshold pulses. The controller may be configured to adjust the re-coinciding trigger threshold based on a timing variation greater than 10% from the baseline, and / or may be configured to adjust the re-coinciding trigger threshold when the dark count rate of one or more of the plurality of PET detectors exceeds a threshold dark count rate. The threshold dark count rate may, for example, range from approximately 3 Mcps to approximately 10 Mcps. Optionally, the controller may further include a current detector configured to measure the bias current of one or more of the plurality of PET detectors, wherein the controller is configured to adjust the re-coinciding trigger threshold when the bias current exceeds a threshold bias current value. The threshold bias current value may range from approximately 0.1 mA to approximately 5 mA, for example, approximately 1 mA, approximately 3 mA. Alternatively or additionally, the controller may be configured to adjust the re-trigger threshold when the radiation dose emitted from the radiation source exceeds a threshold radiation level. The threshold radiation level may be from about 0.1 cGy / min to about 1 cGy / min. The controller may further include a signal processor and a switch configured to selectively transmit the PET detector output signal to the signal processor. The switch may be configured to pause communication of the PET detector output signal to the signal processor for a predetermined time period after each radiation pulse, wherein the ratio of the predetermined time period to the duration of each radiation pulse may be between about 25:1 and about 100:1.The controller can be configured to pause communication between the PET detector output signal and the signal processor for the duration of each radiation pulse and a predetermined time period after each radiation pulse. The controller can be configured to pause communication between the PET detector output signal and the signal processor based on a gating signal. The gating signal may cause the controller to pause communication between the PET detector output signal and the signal processor for at least 100 μs after each radiation pulse. In some variations, the gating signal may cause the controller to pause communication between the PET detector output signal and the signal processor for at least 200 μs after each radiation pulse. Alternatively or additionally, the controller can be configured to adjust the overlap trigger threshold at least in part based on the timing schedule of the radiation pulses.

[0011] A method for automatically adjusting the overlap trigger threshold of a PET detector is also disclosed herein. The method may include: measuring a characteristic of a radiotherapy system comprising two or more PET detectors having overlap trigger thresholds; determining whether the measured characteristic exceeds a predetermined threshold for the characteristic; and adjusting the overlap trigger threshold based on the determination that the measured characteristic exceeds the threshold for the characteristic. Adjusting the overlap trigger threshold may include: increasing the overlap trigger threshold if the measured characteristic exceeds the predetermined threshold for the characteristic, or decreasing the overlap trigger threshold if the measured characteristic is at or below the predetermined threshold for the characteristic. The measured characteristic may be the dark count rate of two or more PET detectors, and the predetermined threshold may be a dark count rate threshold. The measured characteristic may be the bias current of two or more PET detectors, and the predetermined threshold may be a bias current threshold. The radiotherapy system may include a temperature sensor, and the measured characteristic may be a temperature measurement, and the predetermined threshold may be a temperature threshold. Alternatively or additionally, the radiotherapy system may include a radiation source having a pulse counter, and the measured characteristic may be a pulse count measured from the pulse counter, and the predetermined threshold may be a pulse count threshold. A radiotherapy system may include a radiation source and a collimator, wherein the radiation source and the collimator may be configured to operate together with a predetermined timing tolerance, and wherein the measured characteristic may be a deviation from the predetermined timing tolerance, and a predetermined threshold may be a timing deviation threshold.

[0012] A method for detecting positron annihilation emission paths is also disclosed herein. The method may include: directing one or more radiation beam pulses to a target region, wherein the target region is PET-avid; detecting a first positron emission path defined by a first positron annihilation photon pair, which is incident on a portion of a plurality of PET detectors within a time window and generates a detector signal exceeding a first coincidence triggering threshold; adjusting the first coincidence triggering threshold to a second coincidence triggering threshold; and detecting a second positron emission path defined by a second positron annihilation photon pair, which is incident on a portion of a plurality of PET detectors within a time window and generates a detector signal exceeding the second coincidence triggering threshold. The first coincidence triggering threshold may be adjusted to the second coincidence triggering threshold after a predetermined number of radiation beam pulses have been directed to the target region. Adjusting the first coincidence triggering threshold may be based at least in part on a timing schedule of the radiation pulses. The second coincidence triggering threshold may have a larger value than the first coincidence triggering threshold; for example, the second coincidence triggering threshold may be approximately four photon triggers, and the first coincidence triggering threshold may be approximately two photon triggers. In some variations, the predetermined number of radiation pulses can be approximately 1000. The predetermined number of radiation pulses can be a first predetermined number of radiation pulses, and the method can further include: adjusting a second coincidence triggering threshold to a third coincidence triggering threshold after a second predetermined number of radiation pulses have been directed to the target region; and detecting a third positron emission path defined by a third positron annihilation photon pair, which is incident on a portion of a plurality of PET detectors within a time window and generates a detector signal exceeding the third coincidence triggering threshold. The third coincidence triggering threshold can be greater than the second coincidence triggering threshold, and the second predetermined number of radiation pulses can be greater than the first predetermined number of radiation pulses. The third coincidence triggering threshold can be from approximately four photon triggering to approximately six photon triggering, and the second predetermined number of radiation pulses can be approximately 2000. Each radiation pulse can have a pulse width, and the plurality of PET detectors can communicate with a controller including a signal processor, and the method can further include pausing data communication from the PET detectors to the signal processor for a predetermined time period after each radiation pulse, wherein the ratio of the predetermined time period to the pulse width is between approximately 25:1 and approximately 100:1. Optionally, the suspension of data communication can be based on a gating signal. The gating signal may cause data communication from the PET detector to the signal processor to be suspended for at least 100 μs after the radiation pulse, or the gating signal may cause data communication from the PET detector to the signal processor to be suspended for at least 200 μs after each radiation pulse. When the dark count rate of one or more PET detectors exceeds a threshold dark count rate, a first overlap trigger threshold can be adjusted to a second overlap trigger threshold. The threshold dark count rate can be from approximately 3 Mcps to approximately 10 Mcps.When the bias current of one or more PET detectors among a plurality of PET detectors exceeds a threshold bias current value, the first re-coincidence trigger threshold can be adjusted to a second re-coincidence trigger threshold. For example, the threshold bias current value can be from about 0.1 mA to about 5 mA, such as about 1 mA, or about 3 mA. When the radiation dose emitted from the radiation source exceeds a threshold radiation level, the first re-coincidence trigger threshold can be adjusted to a second re-coincidence trigger threshold. For example, the threshold radiation level can be from about 0.1 cGy / min to about 1 cGy / min.

[0013] Also disclosed is a radiotherapy system comprising: a radiation source configured to deliver one or more radiation pulses to a PET-avid region of interest during one or more radiation intervals; a plurality of PET detectors configured to detect one or more positron emission paths emitted by the PET-avid region of interest during one or more detection intervals; and a radiation blocking filter movable above the plurality of PET detectors. The radiation blocking filter may be configured to be positioned above the plurality of PET detectors during one or more radiation intervals and to be positioned away from the PET detectors during one or more detection intervals.

[0014] A radiotherapy system is disclosed herein, comprising: a radiation source configured to direct one or more radiation pulses toward a PET-avid region of interest; a plurality of PET detectors configured to detect positron annihilation photons; a current detector configured to measure bias currents of the plurality of PET detectors; and a controller configured to receive photon data outputs from the plurality of PET detectors, wherein the controller is configured to detect overlapping positron annihilation photon pairs by adjusting the photon data outputs using a gain factor having a value based on the bias current measured during the treatment phase (e.g., calculated based on the measured bias current). The controller may be configured to adjust the gain factor when the bias current exceeds a threshold bias current value, for example, the threshold bias current value may be from about 0.1 mA to about 1 mA. In some variations, the gain factor may be a ratio between the measured bias current and the magnitude of the peak shift in the detection of positron annihilation photons in the photon data output. Adjusting the photon data output may include multiplying the photon data output by the gain factor or linearly shifting the photon data output by the gain factor. Alternatively or additionally, the controller can be configured to adjust the gain factor after a threshold number of emission pulses has been directed to the region of interest; for example, the threshold number of emission pulses may be approximately 1000 emission pulses. In some variations, the gain factor may be a first gain factor, and the threshold number of emission pulses may be a first threshold number of emission pulses; the controller can be configured to adjust the first gain factor to a second gain factor after a second threshold number of emission pulses has been directed to the region of interest. The second gain factor may be greater than the first gain factor, and the second threshold number of emission pulses may be greater than the first threshold number of emission pulses. Alternatively or additionally, the controller can be configured to calculate the peak position of annihilated photons based on photon data output from multiple PET detectors, and adjust the gain factor based on the offset of the peak position from the baseline level. Alternatively or additionally, the controller can be configured to adjust the gain factor when the dark count rate of one or more of the multiple PET detectors exceeds a threshold dark count rate; for example, the threshold dark count rate is from approximately 3 Mcps to approximately 10 Mcps. Alternatively or additionally, the controller may be configured to adjust the gain factor when the amount of radiation emitted from the radiation source exceeds a threshold radiation level, for example, the threshold radiation level may be from about 0.1 cGy / min to about 1 cGy / min.

[0015] In some variations, the controller may further include a signal processor and a switch configured to selectively transmit the PET detector output signal to the signal processor. The switch may be configured to pause communication of the PET detector output signal to the signal processor for a predetermined time period after each radiation pulse, wherein the ratio of the predetermined time period to the duration of each radiation pulse may be between approximately 25:1 and approximately 100:1. The controller may be configured to pause communication of the PET detector output signal to the signal processor for both the duration of each radiation pulse and the predetermined time period after each radiation pulse. For example, the controller may be configured to pause communication of the PET detector output signal to the signal processor based on a gating signal. In some variations, the gating signal may cause the controller to pause communication of the PET detector output signal to the signal processor for 100 μs or longer after each radiation pulse; for example, the gating signal may cause the controller to pause communication of the PET detector output signal to the signal processor for 200 μs or longer after each radiation pulse. Alternatively or additionally, the controller may be configured to adjust the gain factor at least in part based on the timing schedule of the radiation pulses. Attached Figure Description

[0016] Figure 1A This is a schematic diagram of a variant of a radiotherapy system, shown from the front.

[0017] Figure 1B The scattered X-rays that may cause afterglow in the PET detector and the output curves of the PET detector affected by the afterglow are schematically depicted.

[0018] Figure 2A This is a flowchart of a variation of the method used for dynamic gain adjustment.

[0019] Figure 2B This is a flowchart of a variation of a method for adjusting the threshold of a dynamic PET detector.

[0020] Figure 3 This is a flowchart of a variation of a method for dynamic PET detector threshold adjustment based on PET detector noise levels.

[0021] Figure 4 This is a flowchart of another variation of a method for dynamic PET detector threshold adjustment based on changes in component timing distribution.

[0022] Figure 5 This is a flowchart of a variation of a method for dynamic PET detector threshold adjustment based on the PET detector dark count rate.

[0023] Figure 6A This is a flowchart of a variation of a method for dynamic PET detector threshold adjustment based on PET detector bias current.

[0024] Figure 6B This is a flowchart of a variation of a method for dynamic gain adjustment based on the bias current of a PET detector.

[0025] Figure 7A This is a flowchart of a variation of a method for gating the communication of positron emission data from a PET detector to a controller.

[0026] Figure 7B This is a schematic diagram of a variant of the logic circuit used for gating communication of positron emission data from the PET detector to the controller.

[0027] Figure 7C It is a timing diagram of a variant of a method for gating communication of positron emission data from a PET detector to a controller.

[0028] Figure 8A This is a schematic diagram of a variant of the radiation filter ring in the first configuration.

[0029] Figure 8B It is in the second configuration Figure 8A A schematic diagram of the radiation filter ring.

[0030] Figure 8C It is in the first configuration Figure 8A A side view of a schematic diagram of a radiation filter ring.

[0031] Figure 9A This is a schematic diagram of another variation of the radiation filter ring in the first configuration.

[0032] Figure 9B It is in the second configuration Figure 9A A schematic diagram of the radiation filter ring.

[0033] Figure 10A The parameters and layout of the experimental setup used to measure the afterglow of a PET detector are described.

[0034] Figure 10B This is a schematic diagram of the setup for an experimental apparatus used to measure the dark count rate of a single-crystal PET detector.

[0035] Figure 10C PET detector data and dark count graphs are provided before, during, and after the linear accelerator pulse.

[0036] Figure 10D A graph showing the dark count rate of the PET detector over time after the linear accelerator pulse is provided.

[0037] Figure 10EA graph showing the dark count rate of the PET detector over time after the linear accelerator pulse is provided.

[0038] Figure 11 The parameters and setup of another experimental setup for measuring the afterglow of two overlapping polycrystalline PET detectors are described.

[0039] Figure 12 This is a graph showing the transit time of a PET detector as a function of the time following the linear accelerator pulse.

[0040] Figure 13A Experimental data curves depicting the changes in bias current and temperature as a function of the linear accelerator beam conduction time are presented.

[0041] Figure 13B The calibration curves generated by measuring the bias current and the position of the light peak are depicted.

[0042] Figure 13C A graph depicting the energy resolution of the PET detector over time (each data series interval represents a 10-minute increment, where the linear accelerator beam is turned on at data series value 1 and off at data series value 7), with afterglow corrected by gain adjustment.

[0043] Figure 13D The shift of the temporal resolution centroid over time is depicted (each data series interval represents a 10-minute increment, where the linear accelerator beam is turned on at data series value 1 and turned off at data series value 7). Detailed Implementation

[0044] Some variations of radiotherapy systems may include a therapeutic radiation source (such as a linear accelerator) and one or more PET detectors (e.g., an array of one or more PET detectors) for detecting emissions from positron emission tomography (PET-avid) tissue regions. Prior to treatment, a molecule labeled with a radioactive atom, called a PET radiotracer, may be injected into the patient, and this tracer may preferentially accumulate at one or more tumor regions. The radioactive atom in the patient's body undergoes radioactive decay and emits a positron. Once emitted from the atom, the positron rapidly collides with nearby electrons, and both annihilate. Two high-energy photons (511 keV) are emitted from the annihilation point and travel in opposite directions. When two PET detectors detect the two photons simultaneously, annihilation is known to occur somewhere along the line connecting the two PET detectors. The radiotherapy system can acquire positron emission data before or during treatment, and this data can be used to guide radiation to these tumor regions. For example, a emission-guided radiotherapy system may include multiple PET detectors and a linear accelerator mounted on a gantry that can rotate around the patient. In some variants, multiple PET detectors may include two PET detector arrays mounted opposite each other on a gantry. Emission data acquired in real time by the detectors can be analyzed by a system controller to control gantry rotation, directing radiation from the linear accelerator to the PET-avid tumor region. In some variants, real-time positron emission data can also be used to update treatment plans to address any tumor movement that may occur between treatment planning and treatment phases.

[0045] PET detectors include a scintillation material (e.g., a scintillation crystal such as bismuth germanium oxide, gadolinium silicate, or lutetium silicate) coupled to a sensor (e.g., any photodetector, photomultiplier tube, such as a silicon photomultiplier tube). When a high-energy photon strikes the PET detector, the energy from that photon induces a scintillation event in the scintillation material, which may generate one or more lower-energy (e.g., visible light) photons detected by the photodetector device. The photodetector device may have a baseline dark count rate or dark current, where random fluctuations in the output may be indistinguishable from fluctuations indicating the presence of photons. Dark counting causes detector pixels to be excited by discharge. When a pixel discharges, it draws current from a power source, and the current drawn from the power source can be called the bias current. The bias current may be proportional to the average number of dark counts excited over a period of time plus other constants or slowly varying terms; that is, the bias current may be proportional to the dark current. The dark current may be proportional to the afterglow photocurrent plus the thermal noise current of the PET photodetector. The bias current can be measured using a current measurement device or module that may be included in the PET detector array. Alternatively or additionally, an ammeter connected in series with the PET detector photodetector and power supply can be used to measure the bias current. Measuring the bias current to the photodetector and / or variations in the bias current within a selected or set operating range (e.g., gain and / or sensitivity) can provide an indication of the dark count rate and / or variations in the dark count rate (i.e., variations in the bias current indicate a shift in the dark count rate). For example, as the dark count rate increases, the bias current to the PET detector photodetector may also increase because more current is drawn from the power supply as a greater number of random fluctuations cause the detector pixels to discharge more frequently. Under normal operating conditions, the dark count rate may be relatively low, for example, approximately 2 million dark counts per second (cps). Increased ambient temperature and / or elevated radiation levels can cause an increase in the dark count rate or dark current of the photodetector.

[0046] A radiotherapy system may include at least two PET detector arrays positioned opposite each other on a gantry. For example, a PET detector on the first array may have a corresponding PET detector on a second array located on the opposite side, enabling the detection of two high-energy photons from a positron annihilation event. In one variant, the radiotherapy system may include two PET detector arrays, each comprising 32 PET detector modules (64 PET detector modules in total). Each PET detector module may include a 6×12 subarray of PET detectors, where each PET detector has its own photodetector. In some variants, each PET detector module may measure and output the bias current of all photodetectors in the 6×12 array of PET detectors, and the gain of all photodetectors in the PET detector module can be set via a single gain input value. Because positron emission and annihilation events are random events, the system's PET detectors may detect multiple high-energy photons within a very short time interval, and the controller uses the timing information of each detected photon (e.g., detection time) and the location of the PET detectors detecting these photons to determine which two photons are part of a positron annihilation photon pair. For example, if two high-energy photons are detected by two PET detectors placed opposite each other within a specific time interval (e.g., an overlap time window), the controller can pair the two photons together as originating from the same positron annihilation event, which occurs somewhere along the line connecting the two PET detectors. The overlap time window is the time interval within which the detected photons can be considered to overlap (and processed as if they came from the same positron annihilation event). The overlap trigger threshold can be a trigger threshold that distinguishes between the signal generated by the detection of annihilated photons and the signal generated by scattered radiation and / or other noise sources (e.g., random detector noise, afterglow, thermal noise, etc.). If the annihilation event is located closer to one of the PET detectors than the other, one photon in the pair will have a shorter travel distance (i.e., one photon will have a shorter transit time) and will therefore strike the first PET detector before the second photon strikes the second PET detector. The controller can use the time difference between the detection of photons in a positron annihilation pair to determine the location of an annihilation event occurring on the line between two events detected by the PET. PET detectors with sufficient time accuracy to sense the time-of-flight (TOF) difference of positron annihilation photons can transmit the TOF data to the system controller to calculate the location of the positron annihilation event.

[0047] During the treatment process, a linear accelerator generates high-flux X-ray pulses directed toward a target area. Beam-limiting devices, such as one or more jaws and / or collimators (e.g., multi-leaf collimators), help limit the spread of X-rays and guide them to the target tissue area. These X-rays can interact with the patient, with a portion radiating to the patient's target area (e.g., the tumor area) and a portion being scattered by the patient. The scattered X-rays can interact with components of the radiotherapy system, such as X-ray detectors (e.g., MV or kV detectors) and / or PET detectors. This effect occurs in… Figure 1A The diagram schematically illustrates that the body of patient 120 can scatter X-rays from linear accelerator 130 and target 132. The X-rays from the linear accelerator and target can be shaped by beam-limiting devices (such as a multi-leaf collimator 134) to form a treatment beam 122. The scattered X-rays or radiation 124 can be incident on PET detector 126, triggering a scintillation event (e.g., lower-energy photons) indistinguishable from a scintillation event caused by positron emission tomography (PET), which is then sensed by a photodetector of the PET detector. Other radiotherapy systems, such as proton therapy systems, can also generate scattered X-rays or neutrons. Scattered radiation from a proton source can also excite a scintillation crystal. The afterglow of PET detector 126 caused by scattered radiation (and / or other radiation sources) can accumulate over time and cause the detector to saturate or “blank” for a period of time, making it unable to detect positron emission event data during that blanking interval. Figure 1B An example of the output trajectory 110 from the PET detector 100 is depicted, where at t pulse A linear accelerator pulse 101 is applied at a location. The linear accelerator pulse may have a pulse width from about 1 μs to about 10 μs (e.g., from about 3 μs to about 5 μs, about 3 μs, about 5 μs, about 8 μs, etc.) and an inter-pulse interval P from about 2 ms to about 20 ms (e.g., from about 4 ms to about 10 ms, from about 5 ms to about 15 ms, about 4 ms, about 10 ms, etc.). intervalAnd / or a pulse frequency from about 100 Hz to about 250 Hz. Scattered X-rays / radiation 102 from the linear accelerator pulse can irradiate the PET detector 100, thereby generating afterglow photons 104 in the scintillator material 103, which are then detected by the photodetector 105. As seen in the output trajectory 110, the afterglow photons cause a large number of short-term artifacts in the period immediately following to about 50 µs or longer (e.g., 100 µs), during which the PET detector's responsiveness to positron emission events is reduced or degraded (e.g., PET detector saturation or blanking can be a result of photodetector saturation and / or the scintillator reaching its maximum photon output and / or electrical and / or magnetic interference from the linear accelerator, etc.). This period can be referred to as the blanking interval 112 and is the short-term effect of the detector afterglow. The blanking interval 112 can last from the start of the pulse to about 50 μs (or longer), depending on, for example, the duration and energy of the linear accelerator pulse. After the initial blanking interval 112, the scintillation material of the PET detector can continue to scintillate, thus continuing to generate afterglow photons, but possibly at a lower rate than during the blanking interval 112. These afterglow photons can be generated, for example, by continuous excitation and / or increased energy levels of the scintillation material of the PET detector. Successive incidentness of these afterglow photons on the photodetector 105 can result in a higher level of noise 114 in the output trajectory 110 after the linear accelerator pulse is applied than before. This increased level of noise 114 may take approximately 1–5 hours to decay to the pulse level before the linear accelerator pulse and can be a long-term effect of the afterglow. In the presence of high levels of scattered radiation, afterglow photons can saturate the photodetector (e.g., a silicon photomultiplier tube). Since more than one linear accelerator pulse is emitted during the treatment cycle (e.g., approximately 2 ms to approximately 10 ms between each pulse), the afterglow noise of later pulses may accumulate in the afterglow noise of earlier pulses, potentially leading to increasingly noisy noise on the output trajectory 110 of the PET detector. This can impair the PET detector's ability to acquire accurate and precise positron emission data throughout the duration of one or more treatment sessions. In particular, due to short-term or long-term afterglow effects, it may impair the PET detector's ability to perform transit-time analysis with sufficient precision to detect coinciding positron annihilation photon pairs.

[0048] Another way the afterglow effect can impair the ability of a PET detector to acquire accurate and precise positron emission data throughout the entire duration of one or more treatment sessions is due to degradation of the photodetector's energy resolution. As mentioned above, photodetectors can be saturated by afterglow photons. Photodetectors, such as silicon photomultiplier tubes, can comprise hundreds to thousands of discrete Geiger avalanche photodiodes (which may be called micropixels). Optical photons interacting with a single Geiger avalanche photodiode or micropixel can cause the micropixel to discharge. After discharge, the micropixel requires a finite amount of time to recover. This finite amount of time can be from about 10 ns to about 100 ns. If significant afterglow is present (e.g., determined by an increased bias current exceeding a threshold), the total number of discrete micropixels available for detecting positron emission data may be reduced because they are excited by afterglow photons and the scintillation signal generated by positron annihilation photons cannot be detected. As the photodetector saturates due to afterglow, its effective gain, or cumulative gain, decreases. In other words, compared to the signal output from the photodetector under normal (i.e., non-persistent) conditions, the signal output from the photodetector is reduced due to the persistence effect for a specific flicker event. If the gain of the photodetector decreases, the quantitative accuracy of measuring the total energy of the incident photons (e.g., flicker events) may degrade, which could impair the ability to reject scattered photons. Although the persistence effect may not degrade the sensitivity of the PET detector, it may reduce the quantitative accuracy of energy and the timing resolution for each flicker event.

[0049] Afterglow can also enable photodetectors to detect or record positron annihilation photons at lower energy levels (i.e., 511 keV photons); that is, instead of the 511 keV photon peak being located at the 511 keV level in the energy spectrum, the peak of the 511 keV photon is located at an energy level lower than 511 keV. Since the coincidence detection controller or processor is configured to detect positron annihilation events based on 511 keV photons (e.g., setting the detection window to be centered at the 511 keV level), shifting the 511 keV photon peak to a lower energy level (e.g., outside the detection window) may cause the PET detection system controller or processor to miss the detection of positron annihilation events.

[0050] method

[0051] A method for acquiring positron emission data from a PET detector in the presence of scattered radiation may include adjusting the gain of the PET detector's photodetector (e.g., a photomultiplier) as the detector's afterglow increases, such as... Figure 2AAs shown in the flowchart, method 220 may include setting an initial gain value 222 for the photodetector of the PET detector of the radiotherapy system. This step may be performed during system assembly and / or manufacturing, or may be performed only before the start of the radiotherapy phase. After the gain value has been set, the method may include performing 224 radiotherapy, which may include: injecting a PET tracer into a patient; and activating a linear accelerator to generate and excite radiation pulses to one or more target areas. During the radiotherapy phase, the system controller may monitor one or more parameters and / or characteristics of the linear accelerator and / or the PET detector and / or any other detectors or sensors (e.g., current or voltage sensors, temperature sensors, radiation sensors, etc.). The controller may determine 226 whether one or more of those characteristics meet the criteria for adjusting the gain value of the photodetector of the PET detector. If one or more criteria for adjusting the gain value have been met, the controller may adjust 228 the gain value of the PET detector, for example, by adjusting the bias voltage of the photodetector and / or by adjusting the gain factor used by the controller's processor in data acquisition or analysis (e.g., adjusting the acquisition or analysis software gain factor). For example, one or more system parameters exceeding a predetermined threshold can indicate that the PET detector afterglow has increased to a certain level, and increasing the gain value of the photodetector and / or the data acquisition gain factor can help reduce false detections of overlapping high-energy photons. At increased levels of afterglow, the scintillating material can generate more photons. These afterglow photons can cause the PET detector photodetector to record the detection of 511 keV photons at a lower energy level. That is, the output from the PET detector can indicate that a photon at an energy level lower than 511 keV was detected when a 511 keV photon was actually detected, but the magnitude / energy of the PET detector photodetector output is reduced due to the afterglow. Increasing the gain value of the PET detector photodetector (e.g., by increasing the bias voltage of the photodetector) may help increase the output of the PET detector photodetector so that it accurately reflects the detection of 511 keV photons, which may help improve the detection rate of true overlapping high-energy photons. Alternatively or additionally, the system processor can use a gain factor in data acquisition to compensate for the reduced PET detector output. For example, the system processor can multiply the output of the PET detector by and / or an offset gain factor, the value of which depends on the persistence level. In some variations, Figure 2A The methods described in the middle and Figure 3The method described in section -7 can be implemented using a machine-readable instruction set, which can be stored in the memory of a controller communicating with the PET detector. Data from the radiotherapy system (such as from various sensors, the PET detector, linear accelerator, etc.) can be transmitted to the controller, which can perform calculations (e.g., analyses) based on those measurements and / or store the results of those calculations and / or system data in one or more controller memories. Command signals generated by the controller can be transmitted to components of the radiotherapy system (e.g., the PET detector and / or linear accelerator) to control the operation of those components (e.g., adjusting the gain value of the photodetector of the PET detector).

[0052] A method for acquiring positron emission data from a PET detector in the presence of scattered radiation may include: adjusting the overlap trigger threshold of the PET detector as the detector's afterglow increases, such as... Figure 2B As shown in the flowchart, method 200 may include setting an initial coincidence trigger threshold 202 for the PET detector of the radiotherapy system. This step may be performed during system assembly and / or manufacturing, or it may be performed only before the start of the radiotherapy phase. After the coincidence trigger threshold has been set, the method may include performing radiotherapy 204, which may include injecting a PET tracer into a patient and activating a linear accelerator to generate and excite radiation pulses to one or more target areas. During the radiotherapy phase, the system controller may monitor one or more parameters and / or characteristics of the linear accelerator and / or the PET detector and / or any other detectors or sensors (e.g., current or voltage sensors, temperature sensors, radiation sensors, etc.). The controller may determine 206 whether one or more of those characteristics meet the criteria used to adjust the coincidence trigger threshold of the PET detector. If one or more criteria for adjusting the coincidence trigger threshold have been met, the controller may adjust 208 the coincidence trigger threshold of the PET detector. For example, one or more system parameters exceeding a predetermined threshold may indicate that the PET detector afterglow has increased to a certain level, and increasing the coincidence trigger threshold may help reduce false detection of coinciding high-energy photons. In other words, at higher levels of afterglow, the scintillating material generates more photons. These afterglow photons can degrade or reduce the ability of a PET detector to detect overlapping high-energy photons. Increasing the overlap trigger threshold of the PET detector can help ignore afterglow photons and help improve the detection rate of truly overlapping high-energy photons. In some variations, Figure 2B The methods described in the middle and Figure 3The method described in section -7 can be implemented using a machine-readable instruction set, which can be stored in the memory of a controller communicating with the PET detector. Data from the radiotherapy system (such as from various sensors, the PET detector, linear accelerator, etc.) can be transmitted to the controller, which can perform calculations based on those measurements and / or store the results of those calculations and / or system data in one or more controller memories. Command signals generated by the controller can be transmitted to components of the radiotherapy system (e.g., the PET detector and / or linear accelerator) to control the operation of those components (e.g., adjusting the overlap trigger threshold of the PET detector).

[0053] Standards used for PET detector photodetector gain adjustment (e.g., adjusting the gain value of the PET detector photodetector and / or the gain factor used in positron emission tomography data acquisition) and / or overlap threshold adjustment can be measured on the entire PET detector array and / or PET detector module (i.e., a subarray with PET detectors) and / or on a single PET detector. For example, in a radiotherapy system with two PET detector arrays, each PET detector array includes multiple PET detector modules (e.g., 32 PET detector modules), each PET detector module includes a subarray of PET detectors (e.g., a 6×12 subarray of PET detectors), and each PET detector has its own photodetector. Standards (and / or temperature, bias current, noise level, overlap timing distribution, light peak, dark count rate, etc.) can be measured on the entire PET detector array and / or on each PET detector module and / or on each PET detector. Similarly, gain and / or overlap trigger thresholds can be adjusted for the entire PET detector array and / or each PET detector module and / or each PET detector. For example, all PET detectors in a PET detector module can have the same photodetector gain value (i.e., the bias voltage applied to the module is applied to all PET detector photodetectors), and the bias current measurement can be the cumulative bias current of all PET detectors in the module. The bias current, bias voltage, and / or gain factor of each PET detector module can be different from each other. That is, different levels of afterglow correction can be applied to different PET detector modules. For example, in a radiotherapy system with two PET detector arrays, each with 32 PET detector modules, the afterglow effect of each of the 64 PET detector modules can be corrected by measuring 64 bias currents (and / or temperature, noise level, coincidence timing distribution, light peak, dark count rate, etc.) of the 64 PET detector modules and then applying afterglow correction to each of the 64 PET detector modules individually (e.g., applying 64 possibly different gain and / or coincidence threshold adjustments). Alternatively or additionally, bias current (and / or temperature, noise level, coincidence timing distribution, light peak, dark count rate, etc.) can be measured for each PET detector photodetector and / or across the entire PET detector array having multiple PET detector modules.While the following description and variations may relate to measuring bias currents (and / or temperature, noise level, coincidence timing distribution, light peak, dark count rate, etc.) for a single PET detector and / or photodetector (or for multiple PET detectors and / or photodetectors) and adjusting the gain and / or gain factor and / or coincidence threshold for a single PET detector and / or photodetector (or for multiple PET detectors and / or photodetectors respectively), it should be understood that the description is also applicable to measuring multiple bias currents (and / or temperature, noise level, coincidence timing distribution, light peak, dark count rate, etc.) for multiple PET detectors and / or photodetectors (or for each PET detector and / or photodetector respectively) and adjusting the gain and / or gain factor and / or coincidence threshold for such multiple PET detectors and / or photodetectors (or for each PET detector and / or photodetector respectively).

[0054] exist Figure 3A variation of a method for acquiring positron emission data in the presence of scattering or stray radiation is described. Method 300 may include generating a calibration table 302 between detector noise levels and a coincidence triggering threshold of the PET detector. One method for generating the calibration table may include: creating an environment that induces different degrees or levels of noise on the PET detector; providing a positron emission source (e.g., a seed emitting positrons) that emits positrons at a known rate; and adjusting the coincidence triggering threshold of the PET detector at each noise level until the output of the PET detector corresponds to a predetermined temporal resolution quality or metric. The temporal resolution quality or metric may be determined during the manufacture and / or calibration of the radiotherapy system. The temporal resolution quality may be measured using a calibration source, and the temporal spectrum of coincidentally detected photons may be analyzed. For example, a positron emission point source may have a temporal spectrum following a Gaussian distribution, in which the mean is related to the spatial offset of the point source between PET detectors, while the variance is related to the quality of the temporal resolution capability. One method for quantifying the temporal resolution quality may include calculating the full width at half maximum (FWHM) of the temporal spectrum. Method 300 may further include measuring 304 the noise level of the PET detector during the treatment phase and comparing 306 the measured noise level with a noise level in a calibration table to identify a coincidence trigger threshold corresponding to the measured noise level. 308 The coincidence trigger threshold may be adjusted based on changes in the measured noise level. For example, the coincidence trigger threshold may increase as the noise level on the PET detector increases. Alternatively or additionally, method 300 may be used to adjust the gain value of the PET detector photodetector and / or the gain factor used in positron emission data acquisition (e.g., a gain factor used to multiply and / or offset the outputs of (multiple) PET detectors). For example, variations of method 300 may include: generating a calibration table between detector noise levels and gain values ​​and / or gain factors; measuring the noise level of the PET detector during the treatment phase; and comparing the measured noise level with a noise level in the calibration table to identify a gain value and / or gain factor corresponding to the measured noise level. The gain value and / or gain factor may be adjusted based on changes in the measured noise level.

[0055] Scattered X-rays can interfere with the ability of a PET detector to accurately measure the arrival time of high-energy photons. In the absence of scattered X-rays, the timing accuracy of a PET detector can be characterized by a coincident timing distribution with a certain range of timing errors. As mentioned above, the coincident timing distribution can be measured, for example, using a point calibration source. Time differences with thousands of coincident events can be analyzed, and the coincident timing distributions can be merged and / or histogramized to generate a timing distribution. The full width at half maximum (FWHM) of the timing distribution can be used to characterize the timing resolution of the PET detector or the entire PET system. As the level of scattered radiation increases, the coincident timing distribution can change, causing the timing error range to increase. For example, in the absence of X-ray interference, a PET detector may have a coincident timing distribution with a timing error range of 300 ps FWHM, but in the presence of scattered X-rays, the coincident timing distribution can change, causing the timing error range to increase to 550 ps FWHM. Figure 4A method for acquiring positron emission data in the presence of scattered radiation based on a coincidence timing distribution is described. Method 400 may include: measuring the coincidence timing distribution of a PET detector 402 before the linear accelerator is activated (e.g., before a treatment phase, and / or during manufacturing and / or calibration phases); measuring the coincidence timing distribution of a PET detector 404 during a time period when the linear accelerator has been activated (e.g., during a treatment phase); and comparing 406 the coincidence timing distribution measured in step 404 with the coincidence timing distribution measured in step 402. If the timing distribution changes by more than about 10% compared to the previously measured timing distribution, a coincidence trigger threshold for the PET detector 408 can be adjusted. One way to change the coincidence trigger threshold is by sending a command to readout circuitry (e.g., an ASIC) to increase the voltage of a timing comparator. In another method, the coincidence trigger threshold may be a predetermined number of optical photons counted on a photodetector. In this method, the coincidence trigger threshold can be adjusted by changing (e.g., increasing or decreasing) the number of photons that need to be detected to signal a coincidence event. Alternatively or additionally, method 400 can be used to adjust the gain value of the photodetector of the PET detector and / or the gain factor used in positron emission data acquisition (e.g., a gain factor used to multiply and / or offset the outputs of (multiple) PET detectors). For example, variations of method 400 may include measuring the coincident timing distribution of the PET detector before activating the linear accelerator (e.g., before the treatment phase, and / or during manufacturing and / or calibration phases), measuring the coincident timing distribution of the PET detector during a period of time when the linear accelerator has been activated (e.g., during the treatment phase), and comparing the coincident timing distribution measured during treatment with the coincident timing distribution measured before treatment. If the timing distribution changes by more than about 10% compared to the previously measured timing distribution, the gain value and / or gain factor can be adjusted based on the change in timing distribution.

[0056] The afterglow of the PET detector can lead to an increase in the dark count rate of the photodetector, which may interfere with the accurate detection of positron emission events. Figure 5Another variation of the method for acquiring positron emission data in the presence of scattered radiation is described. In this method, the coincidence triggering threshold can be adjusted based on the change in the dark count rate of the PET photodetector. Method 500 may include measuring the dark count rate of the PET detector 502 before activating the linear accelerator (e.g., before the treatment phase, and / or during the manufacturing and / or calibration phases), measuring the dark count rate of the PET detector 504 during a period when the linear accelerator has been activated (e.g., during the treatment phase), and comparing the dark count rate measured in step 504 with the dark count rate measured in step 502 506. In some variations, the dark count rate can be measured by measuring the bias current of the photodetector, and the comparison in step 506 can be between a dark count rate calculated based on the bias current and / or the bias current measurement itself. Alternatively or additionally, the dark count rate (i.e., measuring the number of low-energy photon triggers) can be measured by counting low-photon triggers. The dark count rate can be measured across the entire PET detector array or module / subarray, and / or on a per-photodetector basis. If the dark count rate measured in steps 504 and 506 deviates from about 2 Mcps to about 10 Mcps (e.g., about 3 Mcps), the overlap triggering threshold of the PET detector(s) can be adjusted in step 508. For example, if the dark count rate increases or exceeds a threshold (e.g., more than about 2 Mcps, more than about 3 Mcps, and / or more than about 10 Mcps), the overlap triggering threshold of the PET detector can be increased. Steps 504-508 can be repeated throughout the treatment phase and / or when using a linear accelerator. Alternatively or additionally, method 500 can be used to adjust the gain value of the PET detector photodetectors and / or the gain factor used in positron emission data acquisition (e.g., a gain factor used to multiply and / or offset the output(s) of the PET detector(s)). For example, a variation of method 500 may include measuring the dark count rate of the PET detector before activating the linear accelerator (e.g., before the treatment phase, and / or during manufacturing and / or calibration phases), measuring the dark count rate of the PET detector during a time period after the linear accelerator has been activated (e.g., during the treatment phase), and comparing the dark count rate measured during treatment with the dark count rate measured before treatment. If the dark count rate measured during treatment deviates from the dark count rate measured before treatment by more than about 2 Mcps to about 10 Mcps (e.g., about 3 Mcps), the gain value and / or gain factor may be adjusted based on the change in timing distribution.

[0057] The effect of PET detector afterglow can also be measured in the bias current of the photodetector. Changes in the bias current can indicate a decrease in the PET detector's ability to acquire positron emission data, and adjusting the re-activation threshold (e.g., increasing the re-activation threshold as the afterglow effect increases) can help improve the accuracy of emission data acquisition. Figure 6A A variation of a method for acquiring positron emission data in the presence of scattered radiation is illustrated. In this method, the re-coincidence trigger threshold can be adjusted based on changes in the bias current of the photodetector. The bias current can be measured by monitoring the voltage source of the photodetector, such as a high-voltage power supply. Method 600 may include measuring the bias current of photodetector 602 before linear accelerator activation (e.g., before treatment, and / or during manufacturing and / or calibration), measuring the bias current of photodetector 604 during the period in which the linear accelerator has been activated (e.g., during treatment), and comparing 606 the bias current measured in step 604 with the bias current measured in 602. If the bias current measured in steps 604 and 606 deviates from about 0.1 mA to about 5 mA, the re-coincidence trigger threshold of PET detector 608 can be adjusted. Alternatively or additionally, the bias current may be measured across the entire PET detector array or module / subarray, and / or on a per-photodetector basis. Steps 604-608 can be repeated throughout the treatment process and / or when using a linear accelerator.

[0058] Alternatively or additionally, the overlap trigger threshold of the PET detector and / or the gain value of the PET detector photodetector and / or the gain factor used in positron emission tomography (PET) data acquisition can be adjusted based on temperature and / or radiation measurements at or around the linear accelerator (or any therapeutic radiation source) and / or the PET detector array. For example, a radiotherapy system may include one or more temperature sensors located at or near the PET detector array and / or the linear accelerator. Temperature data from these sensors can be transmitted to a controller, and the overlap trigger threshold of the PET detector can be adjusted if the temperature at the linear accelerator and / or the PET detector array exceeds one or more thresholds. Similarly, one or more dosimeters (e.g., MOSFET dosimeters, thermoluminescent dosimeters, etc.) may be located at or near the PET detector array and / or the linear accelerator. Radiation data from these dosimeters can be transmitted to a controller, and the overlap trigger threshold of the PET detector can be adjusted if the radiation level at the linear accelerator and / or the PET detector exceeds one or more thresholds. Some methods can also adjust the overlap trigger threshold and / or the gain value of the PET detector photodetector and / or the gain factor used in positron emission tomography (PET) data acquisition based on the radiation output of the linear accelerator (e.g., a gain factor used to multiply and / or offset the output of (multiple) PET detectors). For example, a radiotherapy system may include a dose chamber or ionization chamber disposed in the beam path of the linear accelerator. The ionization chamber can send the amount of radiation emitted by the linear accelerator to a controller that can adjust the overlap trigger threshold of the PET detector and / or the gain value of the PET detector photodetector and / or the gain factor used in PET emission tomography based on the radiation output of the linear accelerator. For example, a table mapping various radiation output thresholds to various overlap trigger thresholds and / or the gain values ​​of the PET detector photodetectors and / or the gain factors used in PET emission tomography can be stored in the controller's memory, and the controller can compare real-time ionization chamber measurements with the thresholds in the table to determine whether to adjust the overlap trigger threshold and / or the gain value of the PET detector photodetector and / or the gain factor used in PET emission tomography. The threshold can be based on the cumulative radiation output from the first pulse emitted by the linear accelerator up to the current time point, and / or it can be based on the radiation output over a predetermined time interval (e.g., the pulse rate during a treatment session). For example, a radiation output level from a linear accelerator to the human torso greater than 0.1 Gy / min can generate a sufficient level of scattered radiation to cause afterglow in a PET detector.

[0059] In some variations, a table mapping the linear accelerator pulse count to various coincidence trigger thresholds and / or the gain values ​​of the PET detector photodetectors and / or the gain factors used in positron emission data acquisition (e.g., gain factors used to multiply and / or offset the outputs of (multiple) PET detectors) can be stored in the controller memory. The number of emission pulses emitted by the linear accelerator can be used by the controller to adjust the coincidence trigger threshold of the PET detector. For example, after the linear accelerator has emitted a first number of pulses (e.g., 10,000 pulses), the controller can adjust the coincidence trigger threshold of the PET detector. When the linear accelerator has emitted an additional number of pulses (e.g., another 10,000 pulses, bringing the cumulative pulse count to 20,000), the controller can again adjust the coincidence trigger threshold and / or the gain values ​​of the PET detector photodetectors and / or the gain factors used in positron emission data acquisition. Before adjusting the coincidence trigger threshold and / or the gain value of the PET detector photodetector and / or the gain factor used in positron emission data acquisition, the number of pulses emitted by the linear accelerator (i.e., the threshold number of radiation pulses) may be approximately 1000, 2000, 4000, 7500, or 12000 pulses, depending on the level of scattered or stray radiation present in the particular treatment system. That is, for systems with higher levels of scattered or background radiation, the number of linear accelerator pulses may be lower than for systems with lower levels of scattered or background radiation before adjusting the coincidence trigger threshold and / or the gain value of the PET detector photodetector and / or the gain factor used in positron emission data acquisition. In some variations, this table can map the linear accelerator pulse rate or pulse schedule (i.e., the number of pulses and / or the pulse timing schedule within a specific time interval) to the PET detector coincidence trigger threshold and / or the gain value of the PET detector photodetector and / or the gain factor used in positron emission data acquisition. One or more of these parameters may be used individually and / or in combination with one or more methods described herein to determine when to adjust the PET detector coincidence trigger threshold and / or the gain value of the PET detector photodetector and / or the gain factor used in positron emission tomography (PET) data acquisition, and / or how much to adjust the coincidence trigger threshold (e.g., increasing or decreasing a specific value, etc.). As an example, the initial coincidence trigger threshold used for the PET detector at the start of a treatment session could be approximately 2 photon triggers. A photon trigger can be a voltage, charge, or count representing the detected photons. For example, a 2-photon trigger means that the timing discriminator of the PET detector is excited when it detects the arrival of two or more photons. After 10,000 emission pulses have been emitted, the coincidence trigger threshold can be increased to approximately 5 photon triggers. After 10,000 emission pulses have been emitted (i.e., a cumulative total of 20,000 emission pulses), the coincidence trigger threshold can be increased to approximately 6 photon triggers.As expected, the number of radiation pulse thresholds before changing the overlap trigger threshold and the increment of the overlap trigger threshold change can be varied from this example.

[0060] Figure 6BA variation of a method is described in which the gain value of the PET detector photodetector can be adjusted based on changes in the bias current of the photodetector, which can compensate for saturation of the photodetector (e.g., silicon photomultiplier tube) due to afterglow. The bias current can be measured by monitoring the voltage source of the photodetector (such as a high-voltage power supply). Method 620 may include: measuring the bias current of the photodetector 622 before activation of the linear accelerator (e.g., before the treatment phase, and / or during manufacturing and / or calibration phases), measuring the bias current of the photodetector 624 during a period when the linear accelerator is activated (e.g., during the treatment phase), and comparing the bias current measured in step 624 with the bias current measured in step 622 626. If the bias current measured in steps 624 and 626 deviates from about 0.1 mA to about 5 mA, the gain value of the PET detector photodetector 628 and / or the gain factor used in positron emission data acquisition can be adjusted. The bias current can be measured across the entire PET detector array or module / subarray, and / or on a per-photodetector basis. Steps 624-628 can be repeated throughout the treatment process and / or when using a linear accelerator. In some variations, the gain values ​​of the photodetectors(s) can be adjusted by regulating the bias voltage of the photodetectors and / or by the gain factor used by the controller's processor in positron emission data acquisition (e.g., adjusting the acquisition or analysis software gain factor). For example, the system processor can use a gain factor in data acquisition to help compensate for reduced PET detector output due to afterglow effects. For example, the system processor can multiply and / or offset the output of the PET detectors by a gain factor whose value depends on the afterglow level (e.g., indicated by the measured bias current). The gain factor can be calculated, for example, by measuring the PET detector output corresponding to 511 keV photons (corresponding to different values ​​of the measured bias current) at different afterglow levels, obtaining the difference between the measured PET detector output value and the nominal output value corresponding to the detection of 511 keV photons (i.e., in the absence of afterglow), and calculating the ratio of this difference to the nominal output value. Alternatively or additionally, the gain factor can be calculated by measuring the peak shift of 511 keV photons at different afterglow levels (corresponding to different values ​​of the measured bias current), obtaining the difference between the shifted peak(s) and the nominal peak at 511 keV (i.e., in the absence of afterglow), and calculating the ratio of this difference to each of the shifted peaks. A calibration table or graph can be generated that maps the measured bias current to gain factor values. Changing the gain value and / or gain factor used in positron emission data acquisition can adjust the energy and timestamp of positron annihilation events recorded by the PET detector to correct for detector saturation caused by afterglow.In some variations, the method for adjusting the gain value and / or gain factor may include: generating a calibration table between a bias current value and a gain value (and / or gain factor); measuring the bias current of the PET detector photodetector during the treatment phase; comparing the measured bias current with the bias current value in the calibration table to identify the gain value and / or gain factor corresponding to the measured bias current; and then adjusting the gain value and / or gain factor according to the calibration table.

[0061] Figure 13A Experimental data curves depicting the changes in bias current and temperature as a function of the linear accelerator beam over time while keeping the bias voltage constant are presented. As depicted there, it can be seen that after the beam is turned on at time point 0, the bias current increases from a low level of about 0.2 mA to about 4.2 mA in about one hour. When the beam is turned off at time point 69 (i.e., 69 minutes after the beam is turned on), the bias current drifts back to its baseline value within about 100 minutes. The temperature also drifts upward as the PET detector heats up, as it must dissipate more power in the photodetector. The heat accumulated in the PET detector can be expressed as the bias current multiplied by the bias voltage. In this experiment, the heat generated in the detector starts at about 11 mW at approximately time point 0 (0.2 mA * 55 V), but increases significantly to about 231 mW at approximately peak time point 69 (4.2 mA * 55 V). Temperature variations also affect the gain of the photodetector, which, as described herein, can be corrected or compensated for by adjusting the gain factor used in positron emission data acquisition.

[0062] A calibration table can be generated by measuring the positron emission data of a calibrated positron emission point source (e.g., Na-22) using a PET detector and tracking (i.e. quantifying) how the measurement varies at different afterglow levels. Figure 13B A calibration plot (which can be presented as a calibration table) is depicted, generated by measuring the bias current and peak position on a PET detector module with a 6×12 subarray of PET detectors. The peak is the position of the 511 keV peak along the energy spectrum of the photodetector as detected by the PET detector photodetector. Photons emitted by positron annihilation events have the same energy, and therefore, each valid event has the same energy value. The "peak" can be the 511 keV peak on the energy spectrum histogram of all detected events. Figure 13BAs can be seen, as the photodetector of the PET detector experiences greater afterglow, the bias current increases and a peak for detecting 511 keV photons is detected, just as if lower-energy photons were detected (e.g., a linear downward shift from 511 keV (nominal) to 400 keV). This downward shift of the 511 keV peak degrades the performance of the photodetector and hinders its ability to identify or report 511 keV photons resulting from positron annihilation events. Figure 13C A graph depicting the energy resolution of the PET detector over time (each data series interval represents a 10-minute increment, where the linear accelerator beam is on at data series value 1 and off at data series value 7) is plotted, wherein the gain of the PET detector photodetector has been adjusted to compensate for the downward shift of the 511 keV peak. As depicted there, the energy resolution remains stable over time through gain adjustment or correction. As previously mentioned, the gain of the photodetector can be adjusted or changed by adjusting the detector's bias voltage. Alternatively or additionally, a gain factor that can be used by the system processor during data acquisition can be adjusted such that the system processor multiplies and / or shifts the PET detector output by this gain factor. In one variation, the gain factor for each bias current value can be the slope of a curve or line versus bias current graph at that bias current value at the peak position. Figure 13B In the example, the gain factor can be the slope of the line (fitted to the data points in the curve), and the gain factor can be multiplied by the PET detector module output to identify positron annihilation events. This can be based on... Figure 13B The curve is used to generate a calibration table, which maps the measured bias current level to the position of the light peak on the energy spectrum. Figure 13C The results of applying gain correction under different afterglow conditions are shown. As illustrated here, adjusting the gain factor under different afterglow conditions / levels helps maintain a relatively constant energy resolution or quality for measuring incident photon energy. Monitoring the position of the light peak in the energy spectrum (e.g., during treatment) can provide an indication of whether the gain has been appropriately adjusted to compensate for or correct for afterglow effects (e.g., the gain is not too high or too low). For example, as an alternative or supplement to a calibration table that maps bias current levels to gain values ​​(e.g., bias voltage levels, software gain factors), a calibration table can be generated that maps the light peak position (e.g., the 511 keV light peak position) to a gain value (and / or the gain factor used in positron emission data acquisition), allowing the gain value and / or gain factor to be adjusted based on the light peak position during treatment.

[0063] Another metric that can be used to determine whether the photodetector gain adjustment properly corrects for the persistence effect is the photodetector's temporal resolution. The time resolution of a photodetector, which can be shifted due to the persistence effect, represents the smallest time interval between two photon detection events that can be distinguished as two separate events by the photodetector. Figure 13D This depicts the shift of the timing centroid over time as the PET detector photodetector is subjected to scattered linear accelerator X-rays that cause afterglow (each data series interval represents a 10-minute increment, where the linear accelerator beam is on at data series value 1, off at data series value 7, and recovery periods are from data series values ​​8–16). In some variations, the timing centroid shift value can be used to predict or estimate the afterglow level. Some methods for adjusting the photodetector gain to correct for afterglow effects may include calculating a calibration table that maps the timing centroid shift to a function of bias current. Alternatively or additionally, a calibration table can be generated that maps the timing centroid shift to gain values ​​(e.g., bias voltage level, software gain factor). During the treatment phase, the timing resolution (e.g., timing centroid) can be measured / monitored, and an indication can be provided as to whether the gain has been appropriately adjusted to compensate for or correct for afterglow effects (e.g., gain is not too high or too low). In some variations, the gain of the PET detector photodetector can be adjusted to correct for any timing centroid drift.

[0064] As previously described, in the presence of scattered radiation, any method for acquiring positron emission data from a PET detector may include measuring and monitoring one or more parameters and / or characteristics of the linear accelerator and / or the PET detector and / or any other detector or sensor (e.g., current or voltage sensors, temperature sensors, radiation sensors, etc.) and determining the afterglow level or severity based on these one or more parameters. That is, parameters such as temperature, bias current, radiation emission level, and / or pulse count can serve as alternatives for quantifying the afterglow level or effect. Based on these measurements, the treatment system can alter the gain of the PET detector photodetector by applying variations in the bias voltage and / or correcting or changing the gain value used by the controller's processor in data acquisition or analysis (e.g., adjusting the acquisition or analysis software gain factor).

[0065] Alternatively or additionally, some methods may include delaying the acquisition of PET data by a controller during and after a specified time interval of linear accelerator radiation pulse. Delaying or pausing PET data acquisition and / or transmission during and after the linear accelerator pulse can help reduce or eliminate the storage and processing of positron emission data with afterglow noise and / or radiation pulse artifacts. The amount of radiation artifacts generated by the linear accelerator pulse and / or afterglow effect can be maximum during and immediately after the pulse, and processing positron emission data with elevated levels of noise or artifacts can lead to erroneous or inaccurate coincidence detection. In some variations, the width of the radiation pulse from the linear accelerator can be about 5 μs or less, and pulsed at a frequency of about 100 Hz to about 300 Hz. In this configuration, the duty cycle of the actual radiation beam on time is from about 0.05% to about 0.15% (i.e., the radiation beam off time is from about 99.85% to about 99.95%). PET detectors and / or controllers can delay and / or gate the acquisition of PET data during and / or after a linear accelerator pulse that has little or no impact on PET sensitivity (e.g., delaying and / or gated positron emission data acquisition during a 5 μs linear accelerator beam pulse). For time-of-flight PET systems, reducing or eliminating relatively noisy positron emission data from time-of-flight calculations can facilitate more accurate position calculations and / or help reduce the error range.

[0066] The duration of the delay time interval can be determined at least in part based on the amount of afterglow from the PET detector, which can be qualitatively and / or quantitatively determined based on one or more of the following: PET detector noise level, detector timing distribution, dark count rate, bias current, temperature, ambient radiation level, etc. (including any parameters described above). For example, the delay time interval can be from about 85 μs to about 500 μs, for example, at least about 100 μs, at least about 200 μs, etc. In some variations, delaying the acquisition of positron emission data by the controller can include gating the reception of positron emission data by the controller such that if the data is detected by the PET detector during a specified time interval following the linear accelerator pulse, the controller will not store the positron emission data. Alternatively or additionally, the transmission of positron emission data from the PET detector to the controller can be delayed such that PET data detected by the PET detector is not transmitted during a specified time interval following the linear accelerator pulse. For example, data transmission from the PET detector to the controller can be paused during a specified time interval following the linear accelerator pulse, and data transmission can be resumed after the specified time interval has elapsed. In some variations, the acquisition of positron emission data delayed by the controller may include reading out the positron emission data stored by the controller after a specified time interval following the linear accelerator pulse. For example, positron emission data may be acquired and stored in the controller memory even during the specified time interval following the linear accelerator pulse; however, the controller will not read the positron emission data from the memory until after the specified time interval has elapsed, and the positron emission data stored in the controller memory reflects the data acquired after the specified time interval.

[0067] Figure 7AA flowchart illustrating a variation of the method for gating positron emission data is provided. As shown, method 700 may include measuring 702 the effect of afterglow on the PET detector. The amount of afterglow may be determined qualitatively and / or quantitatively based on one or more metrics described herein, including but not limited to PET detector noise level, detector timing distribution, dark count rate, bias current, temperature, ambient radiation level, etc. Method 700 may include setting 704 a gate trigger threshold based on the amount of afterglow. The gate trigger threshold is a time interval during which no positron emission data is transmitted from the PET detector and / or the controller does not store positron emission data. For example, if any parameter indicating the afterglow level exceeds a predetermined threshold (e.g., noise level, timing, dark count rate, bias current level, temperature level, radiation level), the gate trigger threshold may be increased to cause a longer period of pause in positron emission data communication between the PET detector and the controller. The gate trigger threshold may be maintained at the increased level until the afterglow returns to the PET detector's pre-afterglow level (or within approximately 5% of the pre-afterglow level). Method 700 may further include gating positron emission data of 706 based on a gate trigger threshold during and / or after the linear accelerator pulse. Method 700 may be executed continuously by the controller during the treatment phase, and / or may be executed at predetermined time intervals. In some variations, method 700 may further include changing or updating the gate trigger threshold from an initial level if the measured afterglow effect exceeds a selected threshold, and / or if the change in afterglow from a previous measurement exceeds a selected threshold. For example, if the amount of detector afterglow remains relatively constant during the treatment phase, the gate trigger threshold may not be updated. However, if there is a substantial change in the amount of detector afterglow (e.g., an increase or decrease in detector afterglow), the gate trigger threshold may be updated. Gating may increase from approximately 500 µs to approximately 10 ms. Alternatively, the gate trigger threshold may remain constant.

[0068] Figure 7BA schematic diagram of a variant of the logic circuitry 709 for gating PET data communication between a PET detector and a processor of a controller is depicted. One or more PET detectors 710 (in one or more PET detector arrays) can output data 711 to a comparator 712. The comparator 712 compares the timing characteristics of the positron emission data 711 with a re-encoder threshold 720 from a processor 716 of the controller. As described above, the re-encoder threshold 720 can be determined at system startup (e.g., it can be started at a preset default value) and / or can be updated when the system is used for treatment. If the positron emission data is within the re-encoder threshold, the positron emission data is output to a gate 714. Gate 714 can be, for example, an AND logic gate. The positron emission data can only be transmitted to the processor 716 when a gating signal 718 from the processor to gate 714 is activated (e.g., "high"). The timing of the gate signal 718 can make it inactive (e.g., "low") during the linear accelerator pulse and / or during the time interval (e.g., delay time) after the time interval after the linear accelerator pulse has passed and / or before the linear accelerator pulse (e.g., "high"). Figure 7C An example of timing diagram 730 is shown, which depicts signal timing between linear accelerator pulses 732, the inversion of a gating signal (which can be considered a "reset" signal), and a common clock signal 736 shared between the linear accelerator, PET detector, and processor. In this example, the linear accelerator pulse 732 may have a pulse width of approximately 5 μs. In a variation shown in timing diagram 730, there is a setup period t approximately 10 µs before the start (i.e., rising edge) of the linear accelerator pulse 733. setup The gate signal 734 is inverted to "high" during the period of linear accelerator pulse 733 and the time interval (e.g., a delay time) following the end of linear accelerator pulse 733 (i.e., the falling edge). The linear accelerator pulse 733 can range from approximately 2 μs to approximately 15 μs, with a duty cycle less than approximately 0.001. The total duration t during which the gate signal is inverted to "high" is... gate The time interval can range from approximately 100 μs to approximately 3 ms. When the inverted phase of the gate signal is "high," positron emission data transmission / acquisition is paused. When the inverted phase of the gate signal is "low," positron emission data transmission / acquisition can resume. The duration of the time between gate signal pulses (i.e., the inter-pulse period t) is the time between the two pulses. interpulse The interval can be as low as about 1 ms, but in some variations, the inter-pulse period between linear accelerator pulses can be between about 3 ms and about 10 ms. As mentioned above, the time interval after the end of the linear accelerator pulse 733 can be adjusted based on the degree of PET detector afterglow.

[0069] Alternatively or additionally, the system can interleave positron emission data acquisition with radiotherapy delivery (e.g., linear accelerator activation). In this method, the PET detector first acquires positron emission data. In some variations, the positron emission data from the PET detector can be used to generate images. The PET detector can be deactivated or disabled after the positron emission data has been acquired and stored (e.g., stored in the controller memory), and / or after a PET image has been generated using the positron emission data. The radiation source (e.g., a linear accelerator or proton source) can be activated after the PET detector has been deactivated, and a radiation pulse can be fired at the target (e.g., the tumor region). In this interleaved mode, positron emission data acquisition and radiotherapy beam emission do not overlap significantly in time. In some variations, the activation of the PET detector and the radiation source can be at a 50 / 50 duty cycle. While this may allow for longer positron emission data acquisition cycles, it may extend the overall length of the treatment cycle.

[0070] The PET system can interlock or stop positron emission tomography (PET) data acquisition when the afterglow level exceeds a predetermined threshold. In one variant, the bias current of the PET detector photodetector can be measured, and if the bias current exceeds a predetermined bias current interlock threshold, the PET detectors can interlock and stop data acquisition. The system controller can continue to query the bias current at regular intervals, optionally generating a notification to a clinician or technician indicating the bias current level and / or afterglow level. When the bias current value falls below the interlock release threshold, the PET detectors can resume data collection (i.e., clear the interlock). In some variants, the interlock release threshold can be the same as the interlock threshold, while in others, the interlock release threshold can be less than (e.g., below) the interlock threshold.

[0071] Oscillation Scattering Shield

[0072] Some variations of radiotherapy systems may include a movable radiation shield or filter that can be positioned above the PET detector during the radiation pulse and removed from the PET detector after the pulse. The radiation shield or filter can absorb scattered radiation and / or deflect it away from the PET detector, which may help reduce the amount of PET detector afterglow. A physical shield or filter that conceals the PET detector during a linear accelerator pulse may also help reduce or eliminate backprojection information associated with the linear accelerator pulse. In some variations, the shield can conceal the PET detector during the linear accelerator pulse and expose it before or after the linear accelerator pulse. Due to the inertia associated with the physical shield or filter, some PET detectors may be concealed for a longer period than the linear accelerator pulse.

[0073] For example, a emission-guided radiotherapy system may include multiple PET detectors and linear accelerators mounted on a gantry that can rotate around the patient. Emission data acquired in real time by the detectors can be analyzed by a system controller. Based on this emission data, the system controller can rotate the gantry to guide radiation from the linear accelerators to the PET-avid tumor region from various excitation angles. In some variations, the linear accelerators and PET detectors may be mounted on a rotatable annular or circular gantry, and the patient treatment area may be positioned along the center of the circular gantry (e.g., along the axis of rotation). The radiotherapy system may include a radiation filter ring, which may include one or more radiation shields or filters. In some variations, the radiation filter ring may include a closed-loop structure, while in other variations, the radiation filter ring may include one or more loop segments (e.g., open loops, partial segments of loops, or arcs, etc.). The radiation filter ring may be sized to fit within the inner diameter of a first circular gantry. The radiation filter ring may have the same axis of rotation as the first circular gantry and may move independently of the circular gantry. In some variations, the radiation filter ring can rotate, while in others, it can oscillate relative to a circular frame, wherein the radiation filter ring moves laterally into and out of the frame's inner diameter along the frame's axis of rotation. In some variations, the radiation shield or filter may include one or more radiation-blocking or non-transmissive components (e.g., panels comprising high-Z materials) circumferentially locating along the radiation filter ring. Other portions of the radiation filter ring may be transmissive or radiolucent (e.g., comprising low-Z materials). The radiation-blocking portions(i.e., the portions of the radiation filter ring containing the radiation-blocking components) may have a size and shape corresponding to the size and shape of the PET sensor array. In a first configuration (e.g., a radiation-blocking configuration), the radiation filter ring may be positioned such that the radiation-blocking portions are positioned above the PET detector array. In a second configuration (e.g., a radiolucent configuration), the radiation filter ring may be positioned such that the radiolucent portions (i.e., the portions of the radiation filter ring without the radiation-blocking components) are positioned above the PET detector array. A motion controller coupled to the radiation filter ring (which may be separate from and / or independent of the motion controller of the first circular frame) can rotate or oscillate the radiation filter ring to transition between the first and second configurations. The motion controller for the radiation filter ring may include an actuator, an electric motor, and / or a drive mechanism coupled to the filter ring, which provides sufficient power to change the position of the filter ring according to specific time intervals or schedules (as further described below). In some variations, the motion controller may include a spring mechanism and actuator system or mechanism with one or more springs (e.g., pneumatic or hydraulic actuators, cam-based electric motors, slotted linkage motors, electromagnetic actuators, etc.).Spring mechanisms can assist actuator systems or mechanisms by providing additional power to help accelerate filter ring movement and / or counteract any energy loss in the motion system due to friction and / or resistance.

[0074] In some variations, the rotatable radiation filter ring can rotate halfway for each linear accelerator pulse. Alternatively or additionally, the oscillating radiation filter ring can form half a cycle for each linear accelerator pulse. The oscillating radiation filter ring can be centered above the PET detector array, such that its speed when the radiation blocking portion conceals the PET detector (e.g., its lateral speed across the inner diameter of the PET detector array as it moves in and out of the rack) is higher than the speed when the radiation transmission portion is positioned above the PET detector. The time during which the PET detector array is blocked or shielded by scattered radiation (e.g., linear accelerator pulses) is relatively short compared to the time during which the PET detector array is not blocked. That is, for a given linear accelerator pulse duty cycle, the PET detector is in PET data acquisition mode and therefore can be unobstructed by the radiation blocking portion of the radiation filter ring. For example, the motion controller can be synchronized with the linear accelerator such that it moves the radiation filter ring to a first configuration during the linear accelerator pulse (and optionally, during time periods before and / or after the pulse), and moves the radiation filter ring to a second configuration after the pulse (e.g., during the inter-pulse interval). The duty cycle, pulse width, and pulse frequency of the linear accelerator pulses can be transmitted to the radiation filter ring motion controller, causing the radiation filter ring to be in a first configuration during the linear accelerator pulses. In some variations, the motion controller can cause the radiation filter ring to oscillate to cover several times the width of the PET detector array in its path. The length of the oscillation displacement of the radiation filter ring (e.g., the circumferential length or arc length swept by the radiation filter ring when oscillating between the first and second configurations) can be selected such that it is greater than the width of the PET detector array. This can help reduce the proportion of PET detectors that are blocked when the oscillating radiation filter ring is in the first configuration. For example, an oscillation displacement with a length nine times the width of the PET detector array only blocks 5% of the available PET events.

[0075] Figures 8A-8BA variant of a radiotherapy system 800 is depicted, comprising PET detector arrays 802a and 802b and a rotatable radiation filter ring 804 surrounding an aperture 803 of the system. The rotatable radiation filter ring 804 includes a first radiation-blocking portion 806a and a second radiation-blocking portion 806b. The diameter of the rotatable radiation filter ring can be such that it approximates the inner diameter of a circular frame on which the PET detector arrays are mounted. In some variants, the diameter of the rotatable radiation filter ring can be smaller than the inner diameter of the circular frame. The PET detector arrays 802a and 802b can be positioned relative to each other (e.g., relative to each other such that the center of the first PET detector array 802a is positioned approximately 180 degrees to the center of the second PET detector array 802b). Similarly, to correspond to the arrangement of the PET detector arrays 802a and 802b, the radiation-blocking portions 806a and 806b can also be positioned relative to each other. The lengths of the radiation blocking portions 806a and 806b can be selected such that their circumferential lengths correspond to the circumferential lengths of the PET detector arrays 802a and 802b. Figure 8A A first configuration of a rotatable radiation filter ring 804 is depicted, wherein radiation blocking portions 806a, 806b (which may include one or more high-Z materials) are positioned or aligned above PET detector arrays 802a, 802b. For example, the rotatable radiation filter ring may be in this first configuration during linear accelerator pulses. Figure 8B A second configuration of the rotatable radiation filter ring 804 is depicted, wherein when the linear accelerator is inactive (e.g., not emitting pulses, or within the inter-pulse interval), the radiation blocking portions 806a, 806b are not positioned above the PET detector array (i.e., misaligned). Figure 8C yes Figures 8A-8B A schematic cross-sectional view of the system 800. Figure 8C A rotatable radiation filter ring in a first configuration is depicted, wherein the radiation blocking portion is positioned above the PET detector array, such that the PET detector array 802a, 802b is aligned with the radiation shield along the treatment plane 801 through the orifice 803.

[0076] Figures 9A to 9B A schematic cross-sectional view depicting another variation of the radiation filter ring 900 is shown, wherein instead of rotating the radiation filter ring within the inner diameter of a circular frame (which may be arranged around the aperture 903), the radiation filter ring 900 oscillates laterally into and out of the inner diameter of the circular frame. The oscillating radiation filter ring 900 may include first and second radiation blocking portions 902a, 902b, which may have dimensions and shapes corresponding to the size and shape of the PET detector arrays 904a, 904b. Figure 9AA first configuration of an oscillating radiation filter ring 900 is depicted, wherein the oscillating radiation filter ring is positioned within the inner diameter of a circular frame (i.e., the radiation blocking portion is disposed above the PET detector array). As described above, the radiation blocking portion may include one or more high-Z material panels for reflecting and / or absorbing scattered radiation, and when disposed above the PET detector array, it helps reduce the afterglow of the PET detector by reducing or eliminating the incident scattered radiation on the PET detector. Figure 9B A second configuration of the oscillating radiation filter ring is depicted, wherein the radiation blocking portions 902a, 902b are not positioned above the PET detector array (i.e., misaligned). Lateral movement of the oscillating radiation filter ring is indicated by arrow 901. During linear accelerator pulses, the oscillating radiation filter ring can be in the first configuration, such that the PET detector arrays 904a, 904b are protected from radiation along the processing plane 901 through the aperture 903, and can be in the second configuration when the linear accelerator is inactive (e.g., not emitting pulses or during inter-pulse intervals).

[0077] While the aforementioned variations of the rotatable or oscillating radiation filter ring are circular or annular, in other variations, the radiation filter may be a block of radiation-blocking or non-transmissive material that moves above the PET detector during time intervals of desired high levels of scattered radiation and away from the PET detector during time intervals of desired relatively low levels of scattered radiation. For example, the radiation filter may be mounted on an arm, track, etc., and / or may be coupled to an actuator or motor that moves it above or away from the PET detector.

[0078] Example

[0079] Figure 10AAn example of an experimental setup is depicted in which the afterglow of a PET detector was measured and characterized. The short-term and long-term afterglow of the PET detector were measured. An acrylic scattering target 1000 was used to scatter radiation from a 6 MeV linear accelerator 1002. The target 1000 was an acrylic block measuring 40 cm × 40 cm × 20 cm. A PET detector 1004 was used to measure the scattered radiation incident on two simultaneously operating PET detectors to simulate radiation scattering from a patient. The center of the target 1000 was 53 cm from the PET detector 1004. The 6 MeV linear accelerator 1002 generated a 10 cm × 10 cm X-ray radiation field on the acrylic phantom 1000 with a pulse width of 3 μs at 153 Hz. The entire setup was surrounded by multiple lead blocks 1006, each approximately 20 cm in length. A first dose chamber 1008 exists in the path of the linear accelerator beam 1010, a second dose chamber 1008 is located behind the PET detector 1004, and a dose chamber 1008 is located on the wall of the target 1000 furthest from the linear accelerator 1002. A membrane 1009 is located on the wall of the target 1000 closest to the linear accelerator 1002. Figure 10B Is with Figure 10A A schematic diagram of two single-crystal PET detectors 1012 (with an LYSO scintillation crystal) and 1014 (with an LFS scintillation crystal) used together in the device. Both PET detectors 1012 and 1014 use a solid-state photodetector or photomultiplier 1016 (e.g., an MPPC (SiPM) photodetector), which is capable of transit time measurement and used to measure the effect of scattered radiation on PET data acquisition.

[0080] Silicon photomultipliers (SiPMs) used in PET detectors can be characterized by their dark count rate performance. SiPMs may be sensitive to single photons, and dark counting is a hot-electron noise event in the detector. Scintillation detectors produce short-lived afterglows, and these generate photons that persist between scintillation pulses. From the perspective of the detection system, these are approximate dark counts. The dark count rate (DCR) of the sensor is characterized as a function of when the LINAC pulse occurs. Figure 10C The DCR was characterized exactly before the linear accelerator pulse and from 50 µs to 200 µs after the linear accelerator pulse. Based on these analyses, the controller delayed PET data acquisition by approximately 200 µs, which helped reduce the DCR to an acceptable level.

[0081] The initial DCR of the photoelectric sensor is 2M dark counts per second. This matches the specifications provided by the equipment supplier. Figure 10DThe DCR for the LYSO scintillation crystal as a function of the total exposure time to the linear accelerator was plotted. With the linear accelerator pulsed, as expected ( Figure 10D The minimum DCR (at the bottom of the trajectory) occurs exactly before the linear accelerator pulse. If the DCR is measured immediately after the linear accelerator pulse, the DCR is significantly higher. Therefore, there is a short-term constant decay, which contributes to the afterglow.

[0082] Figure 10E The DCR as a function of the total exposure time of the linear accelerator is depicted for two different crystals: LFS and LYSO. For the LYSO scintillation crystal, the DCR is shown. The linear accelerator runs continuously for 3600 seconds or 1 hour. As the linear accelerator runs over time, afterglow tends to accumulate in the scintillation crystal. This afterglow cannot be distinguished from the dark count of the sensor. It significantly degrades the signal-to-noise ratio of the SiPM. Note that the sensor starts with a DCR of approximately 2 M dark counts / second. After 1 hour of afterglow, the dark count increases to over 10 M dark counts / second, or worse by a factor of 5.

[0083] Figure 11 Another example of an experimental setup for measuring afterglow using two polycrystalline silicon PET detectors is depicted.

[0084] Figure 12 The graphs depict the transit time of the PET module at various time points after the linear accelerator pulse, with each curve representing the temporal resolution at different coincidence trigger thresholds. As shown, the afterglow noise can be reduced by increasing the threshold of the PET detector module. Figure 11 The afterglow effect can be mitigated by changing the trigger threshold of the PET detector's electronics. The system's DCR may be measured during treatment, and the threshold can be adjusted to help improve the system's timing performance.

[0085] The radiotherapy system described herein may include a controller having a processor and one or more memories. The controller may include one or more processors and one or more machine-readable memories communicating with the processors. The controller may be connected to the radiotherapy system and / or other systems via wired or wireless communication channels. In some variations, the controller of the radiotherapy system may be located in the same or a different room from the patient. For example, the controller may be coupled to a patient platform or mounted on a trolley or medical cart adjacent to the patient and / or operator.

[0086] The controller can be implemented consistently with a wide range of general-purpose or special-purpose computing systems or configurations. Various exemplary computing systems, environments, and / or configurations suitable for use with the systems and devices disclosed herein may include, but are not limited to, software or other components embodied within or on personal computing devices, network devices, servers, or server computing devices, such as routing / connectivity components, portable (e.g., handheld) or laptop computers, multiprocessor systems, microprocessor-based systems, and distributed computing networks.

[0087] Examples of portable computing devices include smartphones, personal digital assistants (PDAs), mobile phones, tablet computers, phablets (personal computing devices larger than smartphones but smaller than tablet computers), wearable computers in the form of smartwatches, portable music devices, etc.

[0088] In some embodiments, the processor can be any suitable processing device configured to run and / or execute a set of instructions or code, and may include one or more data processors, image processors, graphics processing units, physical processing units, digital signal processors, and / or central processing units. The processor can be, for example, a general-purpose processor, a field-programmable gate array (FPGA), an application-specific integrated circuit (ASIC), etc. The processor can be configured to run and / or execute application processes and / or other modules, processes and / or functions associated with the system and / or its associated network. Underlying device technologies can be provided for various component types, such as metal-oxide-semiconductor field-effect transistor (MOSFET) technology such as complementary metal-oxide-semiconductor (CMOS), bipolar technology such as emitter-coupled logic (ECL), polymer technologies (e.g., silicon conjugated polymers and metal conjugated polymer-metal structures), analog and digital hybrid technologies, etc.

[0089] In some embodiments, the memory may include a database and may be, for example, random access memory (RAM), a memory buffer, a hard disk drive, an erasable programmable read-only memory (EPROM), an electrically erasable read-only memory (EEPROM), a read-only memory (ROM), flash memory, etc. The memory may store instructions to cause a processor to execute modules, processes, and / or functions associated with the system, such as one or more treatment plans, imaging data acquired in previous and / or current treatment sessions (e.g., real-time imaging data), bioactivity, physiological and / or anatomical data extracted from the imaging data, updated or adjusted treatment plans, updated or adjusted dose delivery instructions, radiotherapy system instructions (e.g., instructions that may guide the operation of gantry, therapeutic radiation sources, multi-leaf collimators, PET detectors, and / or any other components of the radiotherapy system), and image and / or data processing associated with treatment delivery.

[0090] Some embodiments described herein relate to a computer storage product having a non-transitory computer-readable medium (also referred to as a non-transitory processor-readable medium) having instructions or computer code thereon for performing various computer-implemented operations. A computer-readable medium (or processor-readable medium) is non-transitory in the sense that the computer-readable medium itself does not include transiently propagating signals (e.g., propagating electromagnetic waves carrying information over a transmission medium such as space or cable). The media and computer code (also referred to as code or algorithm) may be designed and constructed for one or more specific purposes. Examples of non-transitory computer-readable media include, but are not limited to: magnetic storage media such as hard disks, floppy disks, and magnetic tapes; optical storage media such as optical discs / digital video discs (CDs / DVDs); optical disc read-only memories (CD-ROMs) and holographic devices; magneto-optical storage media such as optical discs; solid-state storage devices such as solid-state drives (SSDs) and hybrid solid-state drives (SSHDs); carrier signal processing modules; and hardware devices specifically configured to store and execute program code, such as application-specific integrated circuits (ASICs), programmable logic devices (PLDs), read-only memories (ROMs), and random access memories (RAMs). Other embodiments described herein relate to a computer program product that may include, for example, instructions and / or computer code disclosed herein.

[0091] The user interface can serve as a communication interface between an operator or clinician and a treatment planning system. The user interface may include input and output devices (e.g., touchscreens and displays) and is configured to receive input and output data from one or more of a support arm, external magnet, sensors, delivery devices, networks, databases, and servers. Sensor data from one or more sensors may be received by the user interface and output by one or more output devices in the form of visual, auditory, and / or haptic feedback. As another example, a user may receive operator control from an input device (e.g., a joystick, keyboard, touchscreen) and then process it by a processor and memory for use by the user interface to output control signals to one or more support arms, external magnets, intracavitary devices, and delivery devices.

[0092] In some variations, the output device may include a display device, which includes at least one of light-emitting diodes (LEDs), liquid crystal displays (LCDs), electroluminescent displays (ELDs), plasma display panels (PDPs), thin-film transistors (TFTs), organic light-emitting diodes (OLEDs), electronic paper / electronic ink displays, laser displays, and / or holographic displays.

[0093] In some variations, the radiotherapy system can communicate with other computing devices via, for example, one or more networks, each of which can be any type of network (e.g., wired network, wireless network). A wireless network can refer to any type of digital network not connected by any type of cable. Examples of wireless communication in a wireless network include, but are not limited to, cellular, radio, satellite, and microwave communications. However, a wireless network can connect to a wired network to interface with the Internet, other carrier voice and data networks, business networks, and personal networks. Wired networks are typically carried over copper twisted-pair, coaxial, and / or fiber optic cables. Many different types of wired networks exist, including Wide Area Networks (WANs), Metropolitan Area Networks (MANs), Local Area Networks (LANs), Internet Regional Networks (IANs), Campus Area Networks (CANs), Global Area Networks (GANs) such as the Internet, and Virtual Private Networks (VPNs). In the following text, "network" refers to any combination of wireless, wired, public, and private data networks typically interconnected via the Internet to provide a unified networking and information access system.

[0094] Cellular communications can encompass technologies such as GSM, PCS, CDMA or GPRS, W-CDMA, EDGE or CDMA2000, LTE, WiMAX, and 5G network standards. Some wireless network deployments combine networks from multiple cellular networks or use a hybrid of cellular, Wi-Fi, and satellite communications. In some embodiments, the systems, apparatus, and methods described herein may include radio frequency receivers, transmitters, and / or optical (e.g., infrared) receivers and transmitters for communicating with one or more devices and / or networks.

Claims

1. A radiotherapy system, comprising: A radiation source configured to direct one or more radiation pulses toward the PET-avid region of interest; Multiple PET detectors, wherein the multiple PET detectors are configured to detect positron annihilation photons; and A controller including a signal processor is configured to gate the PET detector output signal from the plurality of PET detectors to the signal processor for a predetermined time period after each radiation pulse, wherein the predetermined time period is longer than the blanking interval of the plurality of PET detectors.

2. The system according to claim 1, wherein, The controller is configured to pause communication between the PET detector output signal and the signal processor for the duration of each radiation pulse and for the predetermined time period following each radiation pulse.

3. The system according to claim 2, wherein, The controller is configured to suspend communication between the PET detector output signal and the signal processor based on a gating signal.

4. The system according to claim 3, wherein, The gating signal causes the controller to pause communication between the PET detector output signal and the signal processor for 100 μs or longer after each radiation pulse.

5. The system according to claim 4, wherein, The gating signal causes the controller to pause communication between the PET detector output signal and the signal processor for 200 μs or longer after each radiation pulse.

6. The system of claim 1, further comprising a rack, and wherein, The plurality of PET detectors includes two PET detector arrays mounted on the rack and located between each other.

7. The system according to claim 1, wherein, The plurality of PET detectors are configured to measure transit time (TOF) data.

8. The system according to claim 1, wherein, The ratio of the predetermined time period to the duration of each radiation pulse is between approximately 25:1 and approximately 100:

1.

9. The system according to claim 1, wherein, The blanking interval is approximately 50 μs or longer.

10. The system according to claim 1, wherein, The controller is configured to receive photon data output from the plurality of PET detectors, and is further configured to detect overlapping positron annihilation photon pairs by adjusting the photon data output using a gain factor.

11. The system of claim 10, further comprising a current detector configured to measure the bias current of the plurality of PET detectors and to calculate the gain factor based on the bias current measured during the treatment phase.

12. The system according to claim 10, wherein, The controller is further configured to monitor parameters of the radiation source and / or the plurality of PET detectors, and to adjust the gain factor if the parameters meet one or more criteria.

13. The system according to claim 10, wherein, The controller is further configured to monitor parameters of one or more sensors and adjust the gain factor if the parameters meet one or more criteria, wherein the one or more sensors include one or more of a current sensor, a voltage sensor, a temperature sensor, and / or a radiation sensor.

14. The system according to claim 1, wherein, The controller is configured to receive photon data output from the plurality of PET detectors, and is further configured to detect coinciding positron annihilation photon pairs based on a coincidence trigger threshold.

15. The system according to claim 14, wherein, The controller is further configured to monitor parameters of the radiation source and / or the plurality of PET detectors, and to adjust the overlap trigger threshold if the parameters meet one or more criteria.

16. The system according to claim 14, wherein, The controller is further configured to monitor parameters of one or more sensors, and to adjust the overlap trigger threshold if the parameters meet one or more criteria, wherein the one or more sensors include one or more of a current sensor, a voltage sensor, a temperature sensor, and / or a radiation sensor.

17. The system according to claim 1, wherein, The predetermined time period is determined at least in part based on one or more characteristics of the plurality of PET detectors, wherein the characteristics of the plurality of PET sensors include one or more of noise level, timing distribution, dark count rate, bias current, temperature and / or ambient radiation level.

18. The system according to claim 1, wherein, The controller further includes a switch configured to selectively transmit the PET detector output signal to the signal processor, wherein the switch is configured to suspend communication between the PET detector output signal and the signal processor for the predetermined time period.

19. The system according to claim 1, wherein, The controller further includes a controller memory for storing the PET detector output signal, and wherein gating the PET detector output signal includes reading the controller memory after the predetermined time period.

20. The system according to claim 1, wherein, The controller further includes a controller memory for storing the PET detector output signal, and wherein gating the PET detector output signal includes not storing the PET detector output signal in the controller memory during the predetermined time period.

21. A radiotherapy system, comprising: A radiation source configured to emit one or more radiation pulses toward the PET-avid region of interest; Multiple time-of-flight (TOF) PET detectors, wherein the multiple TOF PET detectors are configured to acquire positron annihilation event data; as well as A controller is configured to remove radiation pulse noise from acquired positron annihilation event data to generate low-noise positron annihilation event data, and to use the low-noise positron annihilation event data to calculate the location of the positron annihilation event.

22. The system according to claim 21, wherein, The plurality of TOF PET detectors are configured to acquire positron flooding event data of time intervals between the one or more radiation pulses of the radiation source.

23. The system according to claim 21, wherein, The system further includes a current detector configured to measure the bias current of the plurality of TOF PET detectors, and wherein the controller is further configured to generate an interlock if the bias current exceeds a predetermined bias current interlock threshold.

24. The system according to claim 23, wherein, The predetermined bias current interlock threshold ranges from 0.1 mA to 5 mA.

25. The system according to claim 23, wherein, The generated interlock causes the multiple TOF PET detectors to stop acquiring positron annihilation event data.

26. The system according to claim 23, wherein, The generated interlock causes the radiation source to stop emitting radiation pulses.

27. The system according to claim 21, wherein, The controller includes a signal processor, and the controller is further configured to gate the PET detector output signal from the plurality of TOF PET detectors to the signal processor for a predetermined time period after the radiation source emits a radiation pulse.

28. The system according to claim 27, wherein, The predetermined time period is 100 μs or longer.

29. The system according to claim 21, wherein, The controller includes a signal processor, and the controller is further configured to pause the PET detector output signals from the plurality of TOF PET detectors to the signal processor during a predetermined time period, both simultaneously with and after the emission of one or more radiation pulses from the radiation source.

30. The system according to claim 29, wherein, The predetermined time period is 100 μs or longer.

31. The system of claim 21, further comprising a rack, and wherein, The plurality of TOF PET detectors include two TOF PET detector arrays mounted on the rack and located between each other.

32. The system of claim 21, further comprising a sensor, and wherein, The controller is further configured to acquire data values ​​from the sensor and generate an interlock if the acquired data values ​​exceed a predetermined threshold.

33. The system according to claim 32, wherein, The sensor is a temperature sensor, and the controller is configured to generate an interlock if the acquired temperature data value exceeds a predetermined temperature threshold.

34. The system according to claim 32, wherein, The sensor is a radiation sensor, and the controller is configured to generate an interlock if the acquired radiation data value exceeds a predetermined radiation threshold.

35. The system according to claim 32, wherein, The sensor is a voltage sensor, and the controller is configured to generate an interlock if the acquired voltage data value exceeds a predetermined voltage threshold.

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