Ion beam therapy system and gamma ray detection system therefor - Patent Application 20070122997

Abstract

A gamma ray detection system (10) including a detection module assembly (13, 13a, 13b, 13c) including at least two detection modules (14, 14a, 14s) configured for positron emission tomography (PET) scanning of a target zone (4), each detection module including a plurality of stacked scintillator plates (16), each having a major surface (40a) oriented generally facing the target zone and minor lateral surfaces (40b) defining an edge of the scintillator plate, and a plurality of photon sensors (18) attached to the edge layer photon sensors (18a) configured to detect scintillation events in the scintillator plates from gamma rays incident on the major surfaces. The gamma ray detection system is further configured to function as a Compton camera, wherein at least one scintillator plate other than the scintillator plate closest to the target zone is configured as an absorber scintillator plate for the Compton camera.

Description

Disclosure Contents

[0001] [Technical Field] The present invention relates to an ion beam emission device and a detection system therefor. The ion beam emission device may in particular be a medical device for ion beam therapy, e.g., for proton irradiation of tumors. The detection system is for detecting gamma rays. The detection system can be used for dose and range monitoring during ion beam therapy treatment. The use of the detection system is not limited to radiation therapy for humans or animals, but can also be used as a conventional PET scanner or Compton camera for other applications.

[0002] 〔background〕 Proton or heavy ion beam / ion beam therapy is one of the most precise methods of external radiation therapy. Unlike photon beams, which have a high entrance dose that gradually decreases as they pass through the body, ions penetrate tissue and deposit most of their energy near the end of their trajectory, known as the Bragg peak. References herein to the term "ion" in the general sense should also be understood to encompass negatively or positively charged ions, including protons.

[0003] In today's state-of-the-art ion beam systems for radiation therapy, the dose of radiation is typically delivered by a narrow, typically several millimeters long, ion beam 1 of defined energy, which is directed toward the patient and deflected laterally using fast ramped scanning magnets 2, as shown in Figure 1. The beam's penetration depth is controlled by adjusting the beam's energy, and its intensity and lateral position and size upon reaching the target area are registered by a beam intensity and profile monitor 3. In this way, the tumor is irradiated in 3D. The target area can be divided into iso-energy slices 4, which correspond to the penetration depth of a given set of beam energies. Each iso-energy slice is divided into a series of "spots" with different abscissas, each of which receives a certain number of particles.

[0004] In fact, ion beam therapy typically requires a treatment plan (as shown in Figure 2) before treatment begins. During this treatment planning, a computed tomography (CT) scan is typically performed, possibly in combination with an MRI (magnetic resonance imaging) and / or PET (positron emission tomography) scan of the patient and target tissue. The CT / MRI / PET scan is used to delineate the target volume and define the desired dose distribution. Next, as shown in Figure 1, the way the protons should be delivered is calculated: from which entrance angle the proton beam should enter, what beam energy to use to position the Bragg peak at the desired location, the shape and size of the beam before entering the patient, and the number of protons to be delivered per "spot."

[0005] This process typically occurs days or weeks before the start of actual treatment (indicated by time t0 in Figure 2), and treatment for a single patient can take several weeks, spread across several treatment sessions. During this period, the location and volume of the target tissue can change significantly. To validate the treatment plan, the patient is often imaged immediately before each treatment session, ensuring that the patient's position (relative to the table and imaging device) is correct. Anatomical changes that may affect dose distribution can also be detected. In addition to changes in patient positioning and anatomy, a range of other factors exist that can cause discrepancies between the planned and delivered dose distributions, particularly in beam range, as shown in Figure 2. Imaging artifacts can occur, particularly in patients with metal implants, which is not uncommon in patients who have previously undergone surgery.

[0006] Tissue models generated from images are subject to systematic errors as well as the conversion of CT images to proton ranges. Tissue heterogeneity along the beam path from the skin to the target can introduce significant uncertainty into beam range calculations. Patient setup or immobilization can vary from treatment session to treatment session. Organ movement during irradiation, which may be caused by respiratory motion, heartbeat, peristalsis, or slower drift caused by the patient changing position from upright to supine or prone, can also shift the Bragg peak from its desired position. This is of particular concern in situations where the tumor is close to critical organs, such as the spine, optic nerves, or brainstem. Due to the steeper dose falloff in the Bragg peak region, range deviations in ion therapy are more severe than in photon therapy. Range errors can mean that parts of the tumor do not receive any radiation dose (undershoot) or that normal tissue distal to the beam receives a higher dose (overshoot).

[0007] While passing through tissue, protons / heavy ions undergo nuclear reactions, some of which cause the emission of gamma rays. There are two types of gamma rays that can be detected for treatment monitoring: 1) coincident gamma rays resulting from the production of positron-emitting isotopes; and 2) prompt gamma rays resulting from the excitation of target nuclei. The first type can be detected using positron emission tomography (PET) scans, which are widely used today. They have widespread applications in neurology and oncology because they can monitor glucose metabolism and the uptake of other targeted radiotracers in specific organs and tissues. One particular use case for PET scans is in ion beam therapy, where the beam's penetration depth into the patient can be uncertain due to tissue heterogeneity, necessitating the use of safety margins to spare critical organs from the dose and / or ensure that a sufficient dose is delivered to the entire tumor. PET scans can provide information about the exact location within the patient where the radiation dose was deposited. However, there are many practical and technical challenges to using conventional PET scanners in ion beam therapy treatments.

[0008] An example of a conventional PET scan used in clinical workflow is as follows (Shakirin 2011): - Offline PET, in which the PET scan is performed after irradiation, often with a delay of several minutes while the patient is moved from the irradiation room to a separate room housing the PET scanner. Only isotopes with a lifetime of a few minutes can be detected. Offline PET can be performed with conventional PET scanners, but the relatively long delay for PET acquisition, which depends on the distance between the treatment room and the imaging room, prevents the acquisition of emissions from short-lived radionuclide species. Offline PET can only measure the contribution of long half-lives. Performance is further degraded by biological washout of proton-induced PET activity, which reduces the activity level in the target region and results in a "blurry" image. - In-room PET, where the PET scan is performed immediately after irradiation using a PET scanner installed in the treatment room. While this reduces the delay between irradiation and scanning compared to offline systems, there is still some delay. Furthermore, this approach increases the time the irradiation room is occupied, effectively reducing overall patient throughput. -In-beam PET: A customized PET scanner, either built into the treatment site or directly into the gantry, measures positron annihilation activity during irradiation. Real-time data acquisition allows for more precise dose and range control. PET activity levels in tissues are measured by measuring the long half-life components ( 11 C. 13 N) and short half-life components ( 15 O. 10 C) and minimizes the effects of biological washout. However, integrating a dedicated PET system into the beam delivery system for real-time measurements is expensive and technically demanding due to the geometric constraints of integration into the ion beam device in a therapeutic environment and the intensive computation required for real-time measurements based on the numerous signals output by the gamma-ray detectors. Furthermore, the performance of in-beam PET devices is limited, among other things, by the drowning of coincidence events due to prompt gamma emissions during ion beam delivery and the delay in positron annihilation relative to irradiation (stochastic emission depending on the lifetime of the generated isotopes). Nevertheless, PET scanning technology and image reconstruction methods are mature and proven, and image acquisition can be continued after ion beam delivery, even while the patient is outside the room. Image quality can be improved by extending the image acquisition time, even after the end of irradiation. Similarly, the idle time between portals can also be used for imaging. A further important advantage of PET scanning is its ability to measure the total dose.

[0009] WO 2018 / 081404 A1 discloses a PET scanner scintillation detector with edge detection and the possibility of radially stacking several sensor blocks to achieve inherent depth-of-interaction resolution. The edge detection disclosed therein reduces the number of scintillator elements compared to conventional pixelated scintillator arrangements while simultaneously improving depth-of-interaction measurements. However, with multiple scintillator plates, the number of photon sensors remains high, and the associated signal processing requirements for real-time acquisition are stringent.

[0010] Another known technique for verifying proton range is via measurement of prompt gamma (PG) emissions (Knopf 2013). PG emissions are substantially simultaneous with proton emission, and therefore there is essentially no delay between emission and detection during treatment. Spot-by-spot imaging is also possible, as is imaging close to the particle end-range. PG detection thus allows for the rapid detection of significant range deviations. However, PG detection is a nascent technique, image reconstruction is fairly complex, and image quality cannot be improved by increasing the imaging time. Detector performance and gamma absorption efficiency are key factors in image quality. Furthermore, total dose reconstruction is difficult.

[0011] CN107544086A [1] discloses a scintillator-based hybrid Compton-PET imaging device. The gamma-ray detection elements are face-on ("top-on") or edge-on ("side-on"), as shown in FIG. 39. Compton scattering between radially separated detection modules 50 enables Compton camera imaging, and simultaneous photoelectric absorption enables PET scanning. A detection probe is disclosed that includes multiple scintillation crystal arrays 51 separated by radial gaps. However, the radial gaps are interrupted by either a photon sensor array 52 or a light guide (e.g., optical fiber). Thus, the main surface of each scintillation crystal array (generally facing the imaging volume) is covered with photon sensors. However, CN107544086A does not disclose how or whether the desired Compton scattering angle accuracy is achieved.

[0012] Shimazoe2018 [3] discloses a similar setup in which a 2D surface-coupled scintillator array 50 (GAGG:Ce) is coupled to a 2D array of photon sensors 52.

[0013] Summary of the Invention A general object of the present invention is to provide a safe, reliable and accurate ion beam delivery apparatus.

[0014] A particular object of the invention in the medical field is to provide a safe, reliable and accurate ion beam delivery device for radiation therapy of patients.

[0015] Another object of the present invention is to provide a cost-effective detection system that allows accurate real-time imaging of a volume of interest (also referred to herein as a "target zone") that emits positrons and prompt gamma rays.

[0016] Another object of the present invention is to provide a cost-effective detection system for integration into an ion beam delivery device for safe, reliable and accurate real-time control of ion beam delivery.

[0017] A particular object of the invention in the medical field is to provide a cost-effective detection system for safe, reliable and accurate real-time control of ion beam delivery therapy.

[0018] Disclosed herein, according to one aspect of the present invention, is a gamma ray detection system including a detection module assembly including at least two detection modules configured for positron emission tomography (PET) scanning of a target zone, each detection module including a plurality of stacked monolithic scintillator plates, each having a major surface oriented generally facing the target zone and lateral minor surfaces defining an edge of the scintillator plate, the major surface having a surface area greater than the surface area of the lateral minor surfaces, and a plurality of photon sensors attached to each of the edges configured to detect and determine the position within the plane of the major surface of a scintillation event within the scintillator plate from a gamma ray incident on the major surface. The gamma ray detection system is further configured to function as a Compton camera, wherein at least one scintillator plate other than the scintillator plate closest to the target zone is configured as an absorber scintillator plate for the Compton camera.

[0019] Also disclosed herein is an ion beam therapy system for delivering an ion beam to a zone of tissue, comprising a patient support and an ion beam emitter relatively movable about at least a rotational axis, wherein the gamma ray detection system is configured to perform prompt gamma ray detection and PET scanning during, between, and after the ion beam delivery.

[0020] In an advantageous embodiment, the plurality of photon sensors of at least two radially stacked scintillator plates are connected to a processing circuit configured to multiplex the readout of said plurality of photon sensors.

[0021] In an advantageous embodiment, the photon sensors of at least two azimuthally-axially arranged scintillator plates are connected to processing circuitry configured to multiplex the readout of the photon sensors.

[0022] In an advantageous embodiment, at least one radial gap is provided between at least two of said plurality of stacked scintillator plates or between at least two detection modules.

[0023] In an advantageous embodiment, the radial gap height H relative to the thickness T of one of the scintillator plates may typically be in the range 200>H / T>2, preferably in the range 50>H / T>10.

[0024] In an advantageous embodiment, the plurality of photon sensors includes at least one strip multi-layer photon sensor extending over an edge of the plurality of layers.

[0025] In an advantageous embodiment, a plurality of said strip multilayer photon sensors are attached to each edge side of said plurality of stacked scintillator plates.

[0026] In an advantageous embodiment, the at least one strip multi-layer photon sensor is a double-ended strip detector configured to measure the arrival time of a signal at both ends.

[0027] In advantageous embodiments, the plurality of photon sensors includes at least one individual layer photon sensor on at least one edge of each scintillator plate. In preferred embodiments, there are individual layer photon sensors on two edges, or three or more edges of each scintillator plate, particularly for scintillator plates having four or more edges (e.g., square or hexagonal scintillator plates).

[0028] In an advantageous embodiment, the individual layer photon sensors of an assembly of sensor plates, each sensor plate comprising a scintillator plate and an associated photon sensor, are interconnected in a cross-wire or resistor network arrangement, and the readout is the sum and / or weighted sum of the signals of the multiple interconnected individual layer photon sensors.

[0029] In an advantageous embodiment, the individual layer photon sensors of a module are multiplexed such that the number of readout signals is a subset of the total number of module photon sensors.

[0030] In an advantageous embodiment, the detection system further comprises a light-reflective or light-absorbing interface layer between or on at least two of said scintillator plates.

[0031] In an advantageous embodiment, the detection system further comprises a low refractive index gap, for example of air, between at least two of said scintillator plates.

[0032] In an advantageous embodiment, the detection system further comprises an electro-optical shutter between the edge of the at least one scintillator plate and the photon sensor.

[0033] In an advantageous embodiment, the electro-optic shutter includes a light spreader material and thickness configured to spread the light from the scintillation events near the edges.

[0034] In an advantageous embodiment, the surface area S of the major surface of the scintillator plate and the thickness T of the scintillator plate are less than 100 mm 2 ≦S≦40000mm 2 , and in the range of 0.5 mm≦T≦30 mm.

[0035] In an advantageous embodiment, the detection module assembly surrounds the target zone and includes at least one gap or orifice for emitting the ion beam.

[0036] In an advantageous embodiment, photon sensors that are optically edge-coupled to one or more sides of the stack of scintillator plates are mounted on a support substrate that includes edge connectors for coupling to a signal processing circuit board, the edge connectors minimizing gaps between adjacent detection modules of the detection module assembly.

[0037] In an advantageous embodiment, the radial gap satisfies the relationship H / (T1+T2)>5, where T1 and T2 are the thicknesses of the two scintillators surrounding the radial gap, and H is the height of the radial gap.

[0038] In an advantageous embodiment, the total thickness of the stacked monolithic scintillator plates in the radial direction is less than 19 mm.

[0039] In an advantageous embodiment, the gamma ray detection system includes two radially stacked scintillator plates, the ratio of the thickness of the radially inner scintillator plate to the total radial scintillator thickness being in the range of 0.2 to 0.6.

[0040] In an advantageous embodiment, the photon sensor bias voltages of the photon sensors of the individual scintillator plates can be independently adjusted or enabled / disabled.

[0041] In an advantageous embodiment, a photon sensor coupled to at least two radially stacked scintillator plates is connected to processing circuitry configured to apply the Compton kinetic law to determine whether two coincident blocking events correspond to a forward or backward scattered Compton scatter followed by absorption.

[0042] In an advantageous embodiment, the processing circuitry is configured to reject events that may be due to primary gamma rays entering the detector from an outward radial direction.

[0043] In an advantageous embodiment, the processing circuitry is configured to utilize the interaction coordinate of the photoelectric absorption as the LOR endpoint of the small angle forward Compton scattered gamma rays resulting from electron-positron annihilation.

[0044] In an advantageous embodiment, the processing circuitry is configured to discard Compton scatter events beyond a configurable primary gamma ray energy dependent scatter angle to improve angular resolution.

[0045] In an advantageous embodiment, the analog signals from adjacent photon sensors are added prior to digitization or other multiplexing circuitry.

[0046] In an advantageous embodiment, a two-stage Compton camera may be implemented via inter-module scattering between adjacent detection modules of a detection module assembly.

[0047] In an advantageous embodiment, a two-stage Compton camera may be implemented via interblock scattering between adjacent sensor plates.

[0048] In an advantageous embodiment, a three-stage Compton camera may be implemented via interblock scattering between adjacent sensor plates.

[0049] Compared to conventional PET scanners, advantageous features of the present invention are substantially higher spatial resolution and low-cost scalability. The latter is crucial for achieving high sensitivity, which is particularly important in the medical field, especially for proton therapy range and dose verification. According to advantageous aspects of the present invention, a combined PET scanner / Compton camera allows for the utilization of benefits from both PET and prompt gamma imaging (PGI). This technology may also be important for other applications, such as whole-body diagnostic PET or combined PET / SPECT scanners, or other nuclear imaging fields. This combination addresses the major limitations of both in-beam PET and PGI by enabling the fusion of these two imaging technologies in a single device.

[0050] Further objects and advantageous features of the present invention will become apparent from the appended claims and the following detailed description of embodiments of the invention taken in conjunction with the accompanying drawings. [Brief explanation of the drawings]

[0051] [Figure 1] 1 is a schematic diagram of a typical ion beam environment. [Figure 2] 1 is a flowchart of a conventional ion beam irradiation preparation plan illustrating the factors and errors that can cause range uncertainty in conventional therapy irradiation treatments. [Figure 3] 1 shows a flowchart of an ion beam irradiation preparation plan illustrating factors and errors that can cause range uncertainty in irradiation treatments, and corrective actions taken in accordance with embodiments of the present invention to reduce range uncertainty and improve dose accuracy. [Figure 4] FIG. 1 is a perspective view of an ion beam therapy system including a gamma ray detection system according to an embodiment of the present invention. [Figure 5a] 10 is a schematic diagram illustrating a variation of a detection module assembly of a gamma ray detection system according to an embodiment of the present invention. [Figure 5b]10 is a schematic diagram illustrating a variation of a detection module assembly of a gamma ray detection system according to an embodiment of the present invention. [Figure 5c] 10 is a schematic diagram illustrating a variation of a detection module assembly of a gamma ray detection system according to an embodiment of the present invention. [Figure 5d] 10 is a schematic diagram illustrating a variation of a detection module assembly of a gamma ray detection system according to an embodiment of the present invention. [Figure 5e] 10 is a schematic diagram illustrating a variation of a detection module assembly of a gamma ray detection system according to an embodiment of the present invention. [Figure 6a] 2 is a perspective schematic view of a detection module of a detection module assembly of a gamma ray detection system according to an embodiment of the present invention; [Figure 6b] FIG. 6b is a view similar to FIG. 6a with some of the photon sensor support substrates and photon sensors removed to reveal an inner portion of the detection module. [Figure 6c] 3 is a detailed schematic cross-sectional view of a portion of a scintillator plate of a detection module according to an embodiment of the present invention. [Figure 6d] 1 is a photograph taken from the radially distal side of a prototype detection module according to an embodiment of the invention having four radially stacked sensor plates. [Figure 7] 1 is a schematic diagram of a detection module, including a sensor plate and a signal processing and control module, of a gamma ray detection system according to an embodiment of the present invention. [Figure 8a] 1 is a simplified schematic diagram of a portion of a detection module assembly of a gamma ray detection system according to an embodiment of the present invention, illustrating positron annihilation detection (corresponding to PET scanner functionality); [Figure 8b] FIG. 8b is a view similar to FIG. 8a, showing the detection of prompt gamma rays (corresponding to the Compton camera function). [Figure 9a] 1A-1C are schematic perspective views illustrating stacked scintillator plates of different shapes for detection modules according to different embodiments of the present invention. [Figure 9b]1A-1C are schematic perspective views illustrating stacked scintillator plates of different shapes for detection modules according to different embodiments of the present invention. [Figure 9c] 1A-1C are schematic perspective views illustrating stacked scintillator plates of different shapes for detection modules according to different embodiments of the present invention. [Figure 10] 1 is a simplified schematic diagram of a detection module of a gamma ray detection system according to an embodiment of the present invention, showing the arrangement of scintillator plates. [Figure 11a] 1 is a simplified schematic side view of a detection module showing the placement of photon sensors according to an embodiment of the present invention; [Figure 11b] 11a is a further view similar to FIG. 11a of a further embodiment of a photon sensor arrangement according to the present invention; [Figure 11c] 11a is a further view similar to FIG. 11a of a further embodiment of a photon sensor arrangement according to the present invention; [Figure 11d] 11a is a further view similar to FIG. 11a of a further embodiment of a photon sensor arrangement according to the present invention; [Figure 11e] 11a is a further view similar to FIG. 11a of a further embodiment of a photon sensor arrangement according to the present invention; [Figure 12a] 1A-1C are simplified schematic diagrams of photon sensor arrangements in detection modules according to various embodiments of the present invention, particularly for functioning as Compton cameras. [Figure 12b] 1A-1C are simplified schematic diagrams of photon sensor arrangements in detection modules according to various embodiments of the present invention, particularly for functioning as Compton cameras. [Figure 12c] 1A-1C are simplified schematic diagrams of photon sensor arrangements in detection modules according to various embodiments of the present invention, particularly for functioning as Compton cameras. [Figure 13a] FIG. 10 is a simplified schematic diagram of a photon sensor arrangement according to yet another embodiment of the present invention, particularly for functioning as a Compton camera. [Figure 13b] FIG. 13b is a view similar to FIG. 13a of yet another embodiment functioning as a Compton camera. [Figure 14a] FIG. 1 is a simplified schematic diagram of a pair of detection modules of a detection module assembly according to an embodiment of the present invention, wherein adjacent modules function as an inter-module Compton camera according to an embodiment of the present invention. [Figure 14b] FIG. 1 is a simplified schematic diagram of a pair of detection modules of a detection module assembly according to an embodiment of the present invention, wherein adjacent modules function as an inter-module Compton camera according to an embodiment of the present invention. [Figure 15a] 1 is a simplified schematic perspective view of a detection module assembly according to an embodiment of the present invention, which may function as a PET scanner and a Compton camera. [Figure 15b] 1 is a simplified schematic perspective view of a detection module assembly according to an embodiment of the present invention, which may function as a PET scanner and a Compton camera. [Figure 16a] FIG. 1 is a simplified schematic perspective view of a detection module assembly according to different embodiments of the present invention, showing a dual head assembly that can function as a Compton camera for prompt gamma ray detection and as a PET scanner for positron annihilation detection. [Figure 16b] FIG. 1 is a simplified schematic perspective view of a detection module assembly according to different embodiments of the present invention, showing a triple head assembly that can function as a Compton camera for prompt gamma ray detection and as a PET scanner for positron annihilation detection. [Figure 17a] 1 illustrates a photon sensor arrangement with photon strip detectors in a detection module according to an embodiment of the present invention. [Figure 17b] 1 shows the arrangement of photon sensors on the four sides of a scintillator plate stack, with both individual and strip photon sensors according to different variants. [Figure 17c] 1 shows the arrangement of photon sensors on the four sides of a scintillator plate stack, with both individual and strip photon sensors according to different variants. [Figure 17d]1 shows the arrangement of photon sensors on the four sides of a scintillator plate stack, with both individual and strip photon sensors according to different variants. [Figure 17e] 1 shows the arrangement of photon sensors on the four sides of a scintillator plate stack, with both individual and strip photon sensors according to different variants. [Figure 18] FIG. 10 is a simplified schematic diagram of a photon sensor arrangement in a detection module of a gamma ray detection system according to yet another embodiment of the present invention comprising a photon strip detector. [Figure 19a] 10 is a perspective view of a different detection module of a gamma ray detection system according to yet another embodiment of the present invention, showing a photon sensor configuration with a crosshair connection arrangement for crosshair readout. [Figure 19b] 10 is a perspective view of a different detection module of a gamma ray detection system according to yet another embodiment of the present invention, showing a photon sensor configuration with a crosshair connection arrangement for crosshair readout. [Figure 19c] FIG. 19c is a simplified circuit diagram of the crosshair connection readout of the embodiment of FIGS. 19a and 19b. [Figure 20a] FIG. 1 is a simplified schematic diagram of a scintillator plate of a detection module including an electro-optical shutter (EOS), showing an open EOS. [Figure 20b] FIG. 1 is a simplified schematic diagram of a scintillator plate of a detection module including an electro-optical shutter (EOS), showing the EOS closed. [Figure 20c] FIG. 1 is a simplified schematic perspective view of an example stack of scintillator plates, with the top and bottom layers having open EOS and three middle layers having closed EOS. [Figure 21a] FIG. 2 is a simplified schematic perspective view of a scintillator plate with individual photon sensors. [Figure 21b] 10 is a plot showing the average number of photons detected by a single photon sensor along the right edge after a scintillation event. [Figure 22]1 is a plot showing the distribution of errors in a conventional PET scanner using photon sensors each having a 1×1 mm light-receiving area. [Figure 23] 9A and 9B are plots showing the ratio of the number of photon sensors per edge (i.e., referred to as "pixels") for a conventional PET scanner and the present invention for scintillator plates shaped as a triangle (FIG. 9C), a square (FIG. 9A), and a hexagon (FIG. 9B) having three, four, and six sides, respectively. [Figure 24] 1 is a plot illustrating the spatial accuracy of an embodiment of the present invention. [Figure 25] 1 is a plot illustrating the spatial accuracy of an embodiment of the present invention. [Figure 26] 1 is a plot illustrating the spatial accuracy of an embodiment of the present invention. [Figure 27] FIG. 2 is a schematic diagram of a detection module assembly of a gamma ray detection system according to an embodiment of the present invention used to triangulate the source location of a tri-gamma event. [Figure 28] FIG. 2 is a schematic diagram of a sensor plate in an azimuthal axial arrangement according to an embodiment of the present invention. [Figure 29] FIG. 10 is a schematic diagram of multiplexed readout of a 2×2 azimuthal axial arrangement of sensor plates according to an embodiment of the present invention. [Figure 30] FIG. 10 is a schematic diagram of how azimuthally positioned sensor plates can be repositioned to allow for a wider field of view. [Figure 31] 1 is a schematic diagram of an advantageous arrangement of an embodiment of the present invention for achieving an extended axial field of view compared to conventional detection modules. [Figure 32] 10 is a schematic illustration of the detection sensitivity along the axial source for different sensor plate arrangements. FIG. [Figure 33] FIG. 1 is a schematic diagram of a ring-shaped assembly of azimuthally arranged sensor plates having two radial gaps and 2+2+2 radial layers, according to an embodiment of the present invention. [Figure 34]10 shows plots of detection probability for different energies, total radial scintillator thicknesses, and scatter / absorber relationships for two radial layer configurations according to embodiments of the present invention. [Figure 35] 10 shows plots of coordinate components of angular accuracy for Compton reconstruction for various combinations of radial gap and scintillator plate thickness, in accordance with an embodiment of the present invention. [Figure 36] Plots showing the probability of absorbing a primary gamma ray directly via photoelectric absorption ("pe"), or via single Compton scattering ("Compton"), or the sum of the two ("pe or Compton") are shown, with the left plot showing the probability of simultaneous absorption. [Figure 37] 10 shows plots illustrating the angular accuracy of Compton scattering angle reconstruction and contributions from energy resolution limitations and spatial resolution limitations, according to an embodiment of the present invention. [Figure 38] 1 shows a variation of a whole body scanner (PET or Compton-PET depending on the configuration) in a rotating dual-head configuration with two radial gaps according to an embodiment of the present invention. [Figure 39] 1 shows a prior art surface-coupled detection module having an array of scintillation crystals and photon sensors. [Figure 40] A 16:4 multiplexing variant (common anode) of photon sensors along the same edge is shown, so that the sum of the currents of adjacent photon sensors can be read out according to an embodiment of the present invention. [Figure 41] We show the impact of energy and spatial uncertainties on Compton scattering angle reconstruction.

[0052] Referring to the drawings, beginning with FIG. 4 , there is shown an ion beam therapy system 6, in particular for ion beam radiotherapy or proton irradiation of a zone of tissue, according to an embodiment of the present invention. In this embodiment, a patient 5 is positioned on a patient support 7 that is movable relative to an ion beam emitter 8 about at least rotational and translational axes. The patient support 7 may be movable, in particular at least translationally along at least one axis relative to a fixed reference (e.g., the ground), in particular a horizontal axis X, and the ion beam emitter may be rotatable about said horizontal axis X relative to a fixed reference (e.g., the ground). However, the patient support and / or the ion beam emitter may be movable, in translation and / or rotation, along and about multiple axes, up to a fully three-dimensional movement that allows the ion beam emitter to be positioned at any position and angle relative to the patient.

[0053] The ion beam therapy system further includes a gamma ray detection system 10. The gamma ray detection system 10 may, in certain embodiments, be movable relative to the patient support along or about one or more axes. In one embodiment, the gamma ray detection system is movable at least along a translational direction, particularly along an axial direction, and, in a variant, also in an azimuthal rotational direction in coordination with the ion beam emitter 8.

[0054] However, in a variant (not shown), it is also possible to have a gamma ray detection system that is stationary relative to a fixed reference, or that moves only in translation relative to a fixed reference such as the ground.

[0055] In a preferred embodiment, the gamma ray detection system 10 includes a generally ring- or polygon-shaped detection module assembly 13. In one embodiment, the detection module assembly includes an opening 42 to allow the ion beam emitter 8 to transmit ions (e.g., protons) through the opening, with the emission direction of the ion beam emitter 8 being substantially in the same plane as the detection module assembly. This allows for simultaneous and efficient detection of gamma rays emitted from a target zone receiving the ion beam. The detection module assembly 13 may, for example, have a general "C" shape, providing an opening between opposing ends of the C to allow the ion beam emitter 8 to transmit ions through the opening. However, in a variation, a substantially closed ring / polygon shape, e.g., a generally cylindrical detection module assembly, may be provided with an orifice through a portion thereof to allow the ion beam to transmit therethrough (variation not shown).

[0056] The length of the detection module assembly 13 in the direction of the rotation axis X of the ion beam emitter 8 (also referred to herein as the axial direction) may range from about 5 cm to about 200 cm, depending on the variant. For detection configurations with shorter axial lengths, translational movement of the detection module assembly 3, possibly in conjunction with the ion beam emitter, may be performed during ion beam therapy. The detection module assembly may also be translated, according to one embodiment, for scanning of the target zone after ion beam emission or during diagnosis. For detection module assemblies with a length sufficient to extend across the entire target zone, it is possible to have the detection module assembly stationary relative to the patient, so that displacements of the ion beam or ion beam emitter may not be followed by the detection system.

[0057] It is further noted that the motion of the detection module assembly may be parallel or correspond to the motion of the ion beam emitter, or may follow a different motion configured to optimize the detection of prompt gamma rays and positron annihilation gamma rays emitted from the target as a function of the target position, the target environment, and the position and tilt angle of the ion beam emitter 8. The optimal motion of the ion beam emitter and detection system may be obtained, inter alia, from a calibration of the system on a sample tissue.

[0058] A key advantage of the gamma ray detection system 10 used in the ion beam therapy system 6 according to embodiments of the present invention is that detection occurs in real time during proton emission, capturing not only prompt gammas but also positron annihilation gammas. Additionally, positron annihilation gammas emitted for a period of time after proton emission or between successive proton emission pulses during treatment can be detected. This allows proton absorption at the target zone to be continuously monitored, and feedback from the detection system can be used to adjust for precise targeting of the target zone, taking into account movement of the target zone during or after treatment, and avoiding other problems, such as the washout effect, previously discussed with respect to conventional systems.

[0059] Prompt gamma rays emitted from the volume of interest can be detected by a detector functioning as a Compton camera, while positron annihilation gamma rays, which generally have low energy (511 keV), can be detected by a detection module using the PET scanner functional principle; both of these detection methods are integrated into the detection module of a detection assembly according to an embodiment of the present invention, as further described below. Note that PET detection can operate during, between, and after ion beam emission, or alternatively, can be switched on only between and after ion beam emission pulses. During ion beam emission, the rate of prompt gamma emission is very high, which can reduce the accuracy and reliability of measurements of coincident gamma rays from positron emission annihilation. However, for a certain duration after ion beam emission, prompt gamma emission is low and positron annihilation gamma emission continues for a certain period (as is known per se), allowing measurements to be performed during and after ion beam emission.

[0060] 5a and 5b, two different embodiments of a detection module assembly 13 of a gamma ray detection system 10 according to embodiments of the present invention are shown. While the detection module assembly 13 in these figures is shown as a substantially fully closed ring / polygon shape, it will be understood that a portion thereof may be removed to provide a substantially "C" shape with an opening for an ion beam emitter to transmit the ion beam to a target zone. It will also be understood that the detection module assembly 13 may include spatially separated detection modules, for example, in a "dual head" configuration (FIG. 5e) or a "quad head" configuration (FIG. 5f).

[0061] The detection module assembly 13 includes a plurality of detection modules 14. The detection modules 14 may be aligned to form segments in an embodiment such as that shown in Figure 5a, or may be radially staggered in another embodiment such as that shown in Figure 5b, with the radial beams considered from the axis of rotation X. However, various other configurations are possible, whereby the number of modules aligned to form segments or positioned in a substantially circular or polygonal arrangement (as shown) may vary.

[0062] The detection module 14 is configured to function as both a Compton camera 11 and a PET scanner 12, as will be described in further detail herein. However, depending on the application, it is also possible to use a detection module of the present invention that functions only as a Compton camera or only as a PET scanner.

[0063] Each detection module 14 includes multiple stacked scintillator plates 16 and multiple photon sensors 18. The scintillator plates have major surfaces 40a oriented generally facing the target zone or axis X and lateral minor surfaces 40b defining the edges or contours of the scintillator plates. For simplicity, the lateral minor surfaces 40b are also referred to herein as "edges." In an embodiment (not shown), one or more detection modules may be added at the axial ends of the target zone or imaging volume of interest, or at intermediate positions between the axial ends and the radial positions. In an advantageous embodiment, the stacking direction of the scintillator plates in the detection module 14 is perpendicular to the major surfaces. The photon sensors 18 are positioned on the edges of the scintillator plates 16.

[0064] The detection module may, according to an embodiment, comprise a stack of scintillator plates without radial gaps, or, according to another embodiment, a stack of scintillator plates comprising at least one radial gap 17 .

[0065] The radial gaps 17 are particularly useful for the functioning of the Compton camera 11, whereby some scintillator plates act as scatterers and other scintillator plates act as absorbers. Compton kinematic laws, or timing, can be applied to determine the scattering and absorbing layers.

[0066] The interfaces between the scintillator plates may include optically reflective interlayer reflectors 28 to direct light from scintillation events to the scintillator plate edges while allowing gamma rays to pass.

[0067] Instead of or in addition to interlayer reflectors, the scintillator plates may be separated by one or more low-index gaps 31, e.g., of air or of a low-index solid such as a polymer material. The low-index gaps 31 have the effect that the surfaces of the scintillator plates act as internal reflectors, improving transmission of light from scintillation events to the edges of the scintillator plates while allowing gamma rays to pass through the layers. In addition to, or as an alternative to, interlayer reflectors, a light barrier or absorber layer 29 can be inserted between the scintillator plates in combination with the low-index gaps to prevent interlayer light pollution.

[0068] The reflective or absorptive interface layer may constitute a coating on one side of the scintillator plate or on both sides of scintillator plates laminated together.

[0069] The major surface 40a of the scintillator plate is the surface onto which gamma rays generally impinge and may, for example, be substantially perpendicular to the major surface, and the edge 40b, which extends between opposite sides of the scintillator plate, forms the edge of the scintillator plate on which the photon sensor 18 is located. The surface area S and thickness T of the major surface of the scintillator plate may, in preferred embodiments of the present invention, be within the following ranges: 100mm 2 ≦S≦40000mm 2 , and 0.5mm≦T≦30mm; More preferably, 400mm 2 ≦S≦40000mm 2 , and 1mm≦T≦10mm.

[0070] The preferred range seeks to optimize the relationship between the accuracy of the depth of interaction (DOI) measurement (Z direction) and / or the reduction in the number of readout channels on the one hand, and the detection accuracy on the main surface of the scintillation plate (XY plane) of the scintillation location. The optimal range may vary depending on the application.

[0071] Along the edges 40b of the scintillator plate may be provided an edge light spreader material layer 26. The function of the edge light spreader material 26 is to spread the gamma rays so that light from a gamma ray incident on the scintillator very close to one edge 40b is distributed across multiple adjacent photon sensors.

[0072] The edge 40b of the scintillator plate may further comprise a detector-scintillator optical interface 22 that includes an interface material that provides consistent and predictable transmission of photons through the layer to optimize light transmission through the edge to the photon sensors and / or avoid misalignments that may occur due to inconsistent interfaces (e.g., due to air, variable gaps, etc.). The optical interface also serves to spread light from scintillation events that occur near the scintillator edge over multiple photon sensors to improve spatial resolution.

[0073] An electro-optical shutter (EOS) 24 may further be provided along one or more of the scintillator plates, which is electronically operated to be switched on (optically transparent, 24a) or off (absorbing or reflecting, 24b) depending on the operating state of the detection module 14 so that photons pass through the edge to the photon sensor or are blocked from passing to the photon sensor.

[0074] The height H of the radial gap 17 relative to the thickness T of one scintillator plate may typically be in the range 200>H / T>2, preferably 50>H / T>10.

[0075] In a variant, different sensor plates may have different scintillator plate thicknesses. For example, the radially inner scintillator plates, which act primarily as Compton scatterers, may be thinner to reduce the probability of absorption or rescattering of Compton-scattered gamma rays at the radially inner layers. The radially outer scintillator plates may be thicker to increase the probability of total absorption. The scintillator plate thickness may vary as a function of radial position or successive radial positions.

[0076] The height H of the radial gap 17 relative to the thicknesses T1 and T2 of the two scintillator plates radially surrounding the radial gap 17 may typically be in the range 100 > H / (T1 + T2) > 1, and preferably in the range 25 > H / (T1 + T2) > 5.

[0077] It should be noted that the radial direction referred to herein corresponds to the direction Z shown in the drawings showing the detection module.

[0078] The photon sensors 18 disposed along the edge 40b of the scintillator plate 16 may be mounted on a photon sensor support substrate 20, which may be in the form of, for example, a circuit board having circuit traces for interconnecting the photon sensors to the signal processing and control system 30 of the detection module 14. The support substrate 20 may be a flexible or rigid-flex circuit. The flexible circuit may cover one or more edges of the module and be folded around and optically coupled to the edge of the radial stack of scintillator plates. To minimize dead space between the detection modules, it is advantageous to make the photon sensor support substrate thin.

[0079] The support substrate may include protruding guide elements to facilitate alignment of the scintillator edge with the photon sensor.

[0080] The signal processing and control system 30 of the detection module 14 can include, for example, a circuit board 32 and electronic components 34 mounted thereon, including, for example, analog components for signal filtering, signal shaping, multiplexing, and combining of individual photon sensor signals, as well as photon sensor bias voltage components, a microprocessor, and memory for processing and control of the detection module. The circuit board 32 is mounted at the outermost radial end of the module and can include connectors 36a, 36b for connecting the circuit board 32 to the photon sensor support substrate 20 and to the electronic control system of the gamma ray detection system 10 for image reconstruction, as best shown in FIG. 7 in combination with FIGS. 6a and 6b. The support substrate 20 can be configured as a silicon photomultiplier array substrate with edge connectors 36b that advantageously minimize gaps between adjacent detection modules 14 of the detection module assembly 13.

[0081] In a variant, some analog components, such as signal shaping, filtering, or signal multiplexing components, are mounted directly on the photon sensor support substrate in close proximity to the photon sensors.

[0082] Signal processing components include components for analog signal digitization, such as triggering, time stamping, and energy measurement (e.g., charge integration or time-over-threshold). Additional processing components may be low-level event processing for event validation or event rejection using pre-determined or configurable rules, for example, based on Compton kinematics, photon sensor threshold, energy threshold, number of simultaneously triggered sensor plates or photon sensors, or other applicable rules determined from previous detector calibration.

[0083] The analog and digital signal processing components may be distributed so as to be connected to two or more radial stacks of sensor plates, i.e., one or more radial stacks of sensor plates can "share" the analog and digital signal processing components.

[0084] The photon sensors 18 may include individual layer photon sensors 18a and / or strip multilayer photon sensors 18b. In certain embodiments, the photon sensors 18 may include both strip multilayer photon sensors 18b extending radially (in the Z direction) across the edges of the stacked scintillator plates 16 and individual layer photon sensors 18a positioned on individual scintillator plates. The detection module 14 may include multiple strip multilayer photon sensors 18b on each side of the module, as shown in Figures 17a-17e, for example, and may further include a column of individual layer photon sensors 18a on each side of the module, or on only some of the sides, or in some variations, on only one side. The individual layer photon sensor 18a allows for determination of the layer or layers in which incident gamma rays are absorbed, while the multi-layer strip photon sensor 18b (possibly in combination with an illuminated individual layer photon sensor) allows for determination of the incident position of the absorbed gamma rays in a plane perpendicular to the radial direction (i.e., a plane parallel to the major surface 40a of the scintillator plate 16).

[0085] A key advantage of using strip multilayer photon sensors 18b is that they reduce the number of channels that need to be processed by the signal processing and control electronics for a given number of stacked scintillator plates without reducing measurement accuracy. Thus, data processing requirements are significantly reduced, as are the associated costs of the equipment, or alternatively, higher accuracy in interaction depth measurements can be obtained by having a larger number of stacked scintillator plates for a given number of readout channels.

[0086] In a variant, as best shown in Figure 18, instead of providing individual layer photon sensors 18a to determine the interaction depth, it is possible to provide only strip multi-layer photon sensors 18b extending across the stack of scintillator plates, but these strip multi-layer photon sensors are configured to measure the time difference between the ends of the strip multi-layer photon sensors, which is related to the position of the illumination along the strip, from which the layer in which the scintillation event occurred can be inferred.

[0087] In yet another embodiment for reducing the number of readout channels, as best shown in Figures 19a-19c, individual layer photon sensors 18a arranged in a column may be interconnected in a cross-connect configuration, a resistor network, or generally in a multiplexed manner, thus reducing the number of channels. Note that the diodes in the example of Figure 19c represent SiPMs (silicon photomultipliers). Multiplexed readout allows for measurements of the location of scintillation within the scintillator plate to be determined by the intersection between rows and columns of individual layer photon sensors, while reducing the number of electronically processed channels. An electro-optical shutter, shown schematically in Figures 20a-20c, allows for optical blocking of a number of scintillator plates during very high rates of prompt gamma emission to prevent photon sensor signals from multiple simultaneously triggered sensor plates from superimposing in the multiplexed readout and corrupting the information from each triggered sensor plate.

[0088] Using digital silicon photomultipliers as photon detectors, it is possible to enable / disable individual cells of the strip photon detector to mask (ignore) light emitted from selected scintillator layers, which is an alternative way to achieve the same functionality as an optical shutter.

[0089] Note that in a Compton camera configuration, the detection module 14 may include multiple scatter sensor plates 14s with a greater number and surface area than the absorber modules 14s, as illustrated in Figures 15a and 15b. In such a configuration, the positron emission tomography scanner function is performed by the larger number of modules 14s radially closer to the target zone, while the absorber modules 14a located radially away from the target zone function as absorber modules for the Compton camera function. Such an arrangement also allows both the Compton camera and the PET scanner to function simultaneously, further reducing the number of readout channels for signal processing.

[0090] Thus, the functionality of the PET scanner of the present invention can be realized using a stack of monolithic scintillator crystals. The scintillation light propagates from the interaction point toward the side, where it is detected by multiple photon detectors. The photon sensors can be, for example, SiPMs (analog or digital) or other types of detectors known per se. To improve spatial resolution of events located near the side of the scintillator, an optical (non-scintillating) "spreader" material 26 can be inserted between the scintillator and the photon sensor, as previously described. This spreads the light emitted by gamma interactions across multiple pixels, even when the interaction occurs near the side of the scintillator. Examples of spreader materials include glass, silicone rubber, etc., and can vary in thickness, allowing different spreader shapes to be implemented to optimize light yield on the photon sensor. Instead of or in addition to the spreader material, a thin interface optical layer 22, including, for example, grease, adhesive, or meltmount, can be provided between the scintillator plate edge 40b and the photon sensor, as previously described.

[0091] The propagation of light from the interaction point to the photon sensor can be achieved by total internal reflection. This can be achieved by inserting a material with a low refractive index, such as air, between the scintillator plates. The use of air has the advantage of not having any manufacturing constraints and is not subject to degradation over time, use, or radiation. Another feature is the lamination of the scintillator plates with a highly reflective material or film between layers (e.g., ESR = Enhanced Specular Reflector). Care must be taken to ensure that the reflectivity remains sufficiently stable while the device is in operation and / or between expected device calibrations.

[0092] In embodiments of the present invention, the PET scanner functionality has inherent depth-of-interaction capabilities because signals from individual layers can be resolved. Depth resolution is primarily determined by the thickness of the scintillator plates. The thinner the plates, the better the DOI (depth of interaction) resolution. However, as the number of plates 16 increases, the number of photon sensors 18 required also increases. To alleviate this problem, embodiments of the present invention include the use of elongated photon sensors that span multiple scintillator edges, i.e., the multi-layer strip photon sensors mentioned above. Thus, one single channel can measure light from multiple scintillator plates. To resolve which scintillator plate an interaction occurred on, single-plate pixels are included in at least one photon sensor array per side, i.e., the individual layer photon sensors mentioned above.

[0093] The number of layers covered by the strip detector can be tailored to the expected event rate range: for low count rate applications, the possibility of multiple gamma interactions occurring "simultaneously" in several layers (e.g., during the coincidence window or response time of the photon sensor) is negligible. For SiPMs, the practical dead time between events is typically on the order of a few hundred nanoseconds.

[0094] Compton camera imaging typically uses a first "scattering" layer, where X-rays / gamma rays interact via Compton scattering, depositing a fraction of their initial energy, E1. The X-rays / gamma rays are emitted at a slightly different angle Θ from their initial direction, and this change in angle is related to the deposited energy. These photons are then absorbed by a second scintillator plate, the "absorber."

[0095] By calculating the energy deposition E1+E2 in the two scintillator plates and the interaction coordinates in the two layers, information about the initial position of the initial radiation can be inferred.

number

[0096] The coordinates of the interactions in the two layers must be measured, from which the angle can be determined. Unlike PET reconstruction, where the LOR can be drawn between the interaction coordinates of coincident scintillation events, Compton imaging produces a "cone" that emanate from the interaction point in the scattering layer and whose direction and opening angle are given by the energy and coordinate information from the two separate layers.

[0097] When a detector functions as a Compton camera, its angular accuracy is determined primarily by two factors: (1) the energy accuracy with which E1 and E2 can be determined, and (2) the spatial coordinate accuracy that defines the line between the absorption and scattering events, from which the Compton cone of angle Θ is generated for image reconstruction, as shown in Figure 41. A detailed study comparing the magnitudes of these two components has been conducted and is summarized in Figure 35 (coordinate components only) and Figure 37 (coordinate and energy components) under the following assumptions: The energy accuracy is roughly proportional to the square root of the deposited energy, and for example, in the case of LYSO, it is assumed to be about 10% at 511 keV. The spatial coordinate accuracy Δx is ±0.7 mm in both the azimuthal and axial directions (XY). These two components are sufficiently independent that they can be added in quadrature. Energy precision is difficult to overcome because it is a fundamental limitation inherent in the scintillation crystal material and limited by the energy resolution of the photon sensor. Therefore, care can be taken to ensure that the contribution from spatial coordinate precision is at least significantly lower than the energy component. For H / T ≥ 10, this is well achieved at most energies of interest, e.g., for prompt gamma imaging in proton therapy.

[0098] Figure 37 shows the angular accuracy as a function of scattering angle for energies from 0.511 to 7.0 MeV at H / T = 10. It can be seen that for scattering angles above about 40°, the overall angular accuracy (solid line with markers) is dominated by the contribution from energy resolution (solid line), and the contribution from coordinate accuracy (dashed line) is less important.

[0099] For large-angle scattering, especially backscattering events (Θ>90°), the angular accuracy deteriorates significantly. Therefore, to improve image quality, it is beneficial to implement rejection of events whose reconstructed scattering angle exceeds a configurable value. Different upper thresholds may be used for different energies. The thresholds may be implemented as energy discrimination thresholds in the absorber or scattering scintillator plate, or as actual angle thresholds applied after scattering angle reconstruction.

[0100] To accurately determine the abscissa of Compton scattering or photoelectric absorption, a certain minimum energy deposition is required to obtain a sufficiently accurate reading from the scintillator photon sensor. For forward scattering events, the energy deposition in the scattering layer may not reach that threshold, resulting in an uncertainty in the abscissa. In PET scanning mode, this means that the coordinate on at least one side of the line-of-response (LOR) is uncertain. In this case, it is advantageous to use the coordinate of the absorption event instead, since it carries most of the original energy. If scattering and absorption occur in sufficiently radially adjacent scintillator plates, it is possible to use the coordinate of the absorption event instead as the end point of the LOR (assuming the scattering angle is small and the drift length between Compton scattering and absorption is short).

[0101] 6b, 10, 12a-12c, and 14b show examples of detection modules 14 in which one or more scintillator plates 16 facing the object (target) are separated from one or more other scintillator plates 16 by radial gaps 17. The plates closest to the target form the "scattering layer." Introducing radial gaps for the other layers improves angular resolution. The length H of the radial gaps can be adjusted to optimize the conflicting constraints of maximizing angular resolution while keeping the PET-layers as close as possible to each other (for compactness and imaging reconstruction accuracy) and to the target.

[0102] Three-stage Compton camera In a variation of the invention, at least one module may be further configured to function as a three-stage Compton camera, which requires at least two radial gaps (i.e., at least three radially separated sensor plates).

[0103] Cost Issues, Photon Sensors In a conventional PET scanner (Figure 39) that includes a major surface coupled face-on detection module, the total area of the photon sensors in the scintillator block or assembly of individual scintillator rods or pixels is L 2 where L is the length of the scintillator unit perpendicular to the radial direction. For a square edge-coupled detector, the total area of the photon sensor is 4LT, where T is the radial scintillator thickness. To ensure that the area of the edge-coupled photon sensor is smaller than that of the main surface-coupled photon sensor, it is required that L>4T.

[0104] For a conventional radial scintillator thickness of about 20 mm, the side L needs to be at least 80 mm to achieve a comparable or smaller photon sensor area.

[0105] Counting rate problem Particularly for range verification applications, the rate of emitted prompt gammas can be very high. As an example: with a high therapeutic proton rate of 1.2E10 protons / s, the rate of prompt gammas is approximately 1E9 / s (Rohling 2017). At a radial distance of 30 cm, this corresponds to approximately 0.1 gammas / cm² / µs, or approximately 2–2.5 gammas / µs for a 5 cm × 5 cm square detector module. Care must be taken to ensure that the detector is not saturated or blinded by the high instantaneous rate of gamma rays. To improve the ability to distinguish between Compton scattering of gamma rays from one layer to another and simultaneous detection of two independent gamma rays in two layers, events whose deposited energy does not coincide with the desired prompt gamma peak can be rejected.

[0106] Embodiments of the present invention include different configurations that allow the detection module to be tuned for the expected count rate.

[0107] Configuration 1 - Independent layer with pixel detector In a first configuration, shown in Figures 11b and 11c, each scintillator plate has a separate photon sensor along its edge, allowing the lateral position and energy of gamma rays photoelectrically absorbed in one layer, or Compton scattered in one layer and absorbed in another, to be measured. A second Compton scattering in a second layer is of course also possible, as are other interactions, but for simplicity, we will focus here on using the invention as a combined single-scatter Compton camera / PET scanner.

[0108] Configuration 2 - Strip detector across multiple layers In the second configuration (Figure 11a), some of the photon sensor "pixels" are replaced with "strips" that span multiple layers. This has the advantage of reducing the total number of channels and readout complexity. To be able to identify which layer a scintillation event occurred in, each layer has at least one pixel 18a that can detect scintillation light from that layer only. This is a feasible solution if the detector functions as a pure PET scanner. However, Compton camera functionality is more difficult to achieve, since the signal readout from one strip is essentially the sum of two scintillation events in two layers.

[0109] Configuration 3 - Shared detector strips + isolated Compton layer In the third configuration (Figures 11d and 11e), one of the layers (16s in Figure 11d, 16a in Figure 11e) is optically, electronically, or electro-optically isolated from the other layers. For Compton camera functionality, this isolated layer is used either as an absorber (16a in Figure 11e) or a scattering (16s in Figure 11d) part, depending on the predominant direction of the incoming high-energy photons. This layer, in combination with a stack of scintillation plates, is used in a photon sensor configuration as described in configuration 2.

[0110] Configuration 4 - Split Block In the fourth configuration (Figures 12a-12c), the detection module stack of scintillator plates is spatially separated into two sections or blocks 15a, 15s. Each block is arranged similarly to configuration 2. Both blocks can be used as PET detectors. In the case of a Compton camera, one block 15s would be used as the "scatter section" and the other block 15a would be used as the "absorber section."

[0111] Configuration 5 - Compton Pixel In a fifth configuration (FIG. 13b), intended primarily for high-prompt gammas, a single photon sensor 18p (pixel) optically coupled to a scintillation crystal 16p is used as a Compton absorber (or Compton scatterer) in addition to any of the preceding configurations. This pixel should be small enough to achieve the desired spatial / angular resolution. This configuration has the advantage of requiring only a few additional channels to achieve Compton functionality. In a variant, multiple individual pixels can be added as needed. It will also be understood that this configuration can be "inverted," i.e., the single pixel 18p, 16p functions as the scattering layer and the scintillator stack functions as the absorber. In either case, the readout chain can be configured such that when the present invention is operating in Compton mode, only events that trigger the "Compton pixel" are further processed, while all other events are discarded.

[0112] Configuration 6 - Compton Pixel + Compton Layer Configuration 5 can suffer from count rate saturation if the proportion of prompt gammas is very high. To address this, configuration 3 can be modified so that the electronically insulating Compton layer is coupled to a scintillator that can be thinner than the PET scintillator. Reducing the thickness reduces the probability of interaction and the overall count rate. Furthermore, the PET scintillator absorbs some of the prompt gammas, which can further reduce the overall count rate (Figure 13a).

[0113] Inter-module Compton camera Compton camera functionality can also be achieved through inter-module scattering. One detector module 14a acts as a scatterer, and another detector module 14b, e.g., an adjacent module, acts as an absorber (Fig. 14a, 14b). In this configuration, a separate layer for identifying Compton scattering events is not required, and hardware modifications relative to a pure PET scanner are minimal. Compton scattering events can be identified, for example, through total energy discrimination and inter-module coincidence timing. At gamma energies of a few MeV, primarily of interest for proton range verification, forward scattering dominates, and it would be advantageous to introduce a spatial gap 17 between two or several layers 16 to improve angular resolution and increase the probability of inter-module Compton scattering.

[0114] Conventional PET scanners require a circular assembly to ensure that the majority of gamma rays enter the scintillation crystal element approximately perpendicular to the crystal face facing the source. However, off-center emissions can result in parallax errors. The depth of interaction capability of the present invention alleviates this problem. Furthermore, the inventors have realized that the depth of interaction capability can be utilized for non-circular assemblies (e.g., by having a hexagonal assembly as shown in FIG. 5a) to increase the probability of forward scattering across adjacent modules. The present invention also allows for the sensor plate to be positioned close to the patient or in close proximity to the object being scanned.

[0115] In another embodiment, the modules may be arranged in a radially staggered pattern, as shown in Figure 5b, which also serves to increase the probability of inter-module Compton scattering. Such an arrangement reduces the need to introduce radial gaps between the scintillator plates within each module.

[0116] In another embodiment, the detection modules 14s, 14a are themselves arranged in radial groups, as shown in Figure 15a. The inner group 14s closest to the source (target) would function as a PET module and a scatter module. The outer group 14a would function as an absorber module for Compton imaging. Figure 15a shows an example of a 9:1 module arrangement, in which one single absorber detection module 14a is radially offset and centered over the 3x3 PET / scatter detection modules 14s. It will be understood that other arrangements and ratios between scatter / absorber modules are of course possible, such as 1:1, 9:4 (as shown in Figure 15b), and others.

[0117] The solution of using substantially identical modules as both scatterer and absorber, optionally separated in the radial direction and optionally with different scintillator thicknesses, has the advantage of simplifying manufacturing and readout electronics.

[0118] Sensor Plate Deactivation In a crosshair readout (Figures 19a-19c) or in a photon sensor strip 18b shared between layers 16 (Figures 17a-17e, 18), it is possible to block scintillation light from one or several layers 16 to ensure that the column signal originates from only one layer. This can be achieved by mechanical shutters surrounding the edges of each layer. Another option is to use an electro-optical shutter 24 (Figures 20a-20c), such as a polarizing liquid crystal or transflector, switchable between a transmissive state 24a and an absorptive / reflective state 24b using a drive voltage. In the transmissive state, the electro-optical shutter can also act as an optical spreader between the scintillator 16 and the photon sensor 18.

[0119] In an advantageous embodiment, an alternative means of achieving similar functionality is to selectively enable or disable photon sensor bias voltages for groups or individual photon sensors, for example grouped by sensor plate, as illustrated via a bias switch network in Figure 7. Alternatively or additionally, the bias voltage may be adjustable, and the gain of an individual photon sensor, a group of photon sensors, a group of photon sensors optically coupled to a scintillator plate, or a group of sensor plates may be adjustable according to the expected primary gamma ray energy of interest.

[0120] Another alternative is to use digital silicon photomultiplier tubes as photon sensors, whereby individual microcells of the strip photon sensor can be enabled / disabled to mask (ignore) light originating from selected scintillator plates.

[0121] Read Chain As an alternative to detector strips shared between layers, individual pixels along the side of each scintillator block and multiplexed readout chains can be used instead.

[0122] An example of a first row / column readout (crosshairs) for a PET scanner module is shown in Figure 19a. Here, there are five square scintillator plates and five photon sensor columns per side. A total of 100 photon sensors can be read out via 25 channels. Channels Y1-Y5, corresponding to the set of layers 1-5, provide information on which layer a scintillation event occurred in. The readout electronics can be further simplified by using these channels, for example, as signal amplitude threshold triggers to assess whether a scintillation event occurred in each layer. The PET scan data processing circuitry then further processes only events involving exactly one layer, providing column sum signals X1, ..., X N corresponds to light emission within only one layer.

[0123] An example of a second row / column readout (crosshairs) for a combined PET scanner and Compton camera module is shown in Figures 19b and 19c. Here, there are four square scintillator plates and five photon sensor columns per side. A total of 80 photon sensors can be read out via 24 channels. Channels Y1-Y4 in Figure 19b provide information on which layer a scintillation event occurred in and the energy (total amount of light). Energy rejection / filtering can be achieved by a discriminator. In the embodiment of Figure 19b, one layer 16a is spatially separated from the other layers to act as a Compton absorber layer. Channels Z1-Z of the photon sensors in this layer 20 are read out separately. The Compton camera data processing circuitry processes only events involving exactly two layers, one of which is an absorber layer. In this example, 44 channels are required.

[0124] Other multiplexing schemes, such as symmetric charge division, can also be used.

[0125] Another possibility is to use aggregated quantity information from the edges of the scintillator plate. One example is to use the centroid and sum (of charge) of each edge. If reducing the number of readout channels and / or digitizers is a priority, this can be implemented, for example, by a resistor network / ASIC before digitization as a sum and weighted sum of pixels along each edge. The output from edge i, based on a row of N photon sensors, is then reduced to two quantities per scintillator edge. {X1,…X N} i →S tot,i ,S weighted,i

number

number

[0126] The original coordinates of the scintillation event are then calculated as cog i The total energy of the event can be reconstructed based on measurements of the centroid edges using a calibration table consisting of measurements of S across all edges. tot,i is given by the sum of

[0127] Note that alternatively or additionally, aggregations other than centroids can be used, such as the index of the strip with the most counts, a truncated centroid (discarding strips with few counts), full width at half maximum, skew, or more complex functions.

[0128] The technique of using aggregated edge quantities for event reconstruction can be implemented in an analog manner (before digitization), for example using a resistive charge division circuit (CDC), or after digitization with the aim of speeding up the image reconstruction method by reducing the dimensionality of the data set per scintillation event.

[0129] Layer identification based on time difference between double-ended strip detectors Another method of layer identification is to read out the strip photon sensor 18 at both ends 41a, 41b, as shown in Figure 18. Such dual readout of a strip silicon photomultiplier is known per se and can be achieved, for example, by detecting that the propagation speed of the signal across the strip is approximately TIFF0007725074000004.tif730, as reported in Doroud2017

[11] (which uses differential readout for noise suppression). The difference in arrival times of the pulses at both ends, upper (subscript u) and lower (subscript l), is:

number

[0130] System assembly Figure 16a shows a dual-head assembly in a proton therapy environment, and Figure 16b shows a triple-head assembly. In a dual-head assembly, an ion beam 1 enters the target in the y-direction, and two detector assemblies 13a, 13b are positioned symmetrically in the ±z-direction around the target 4. These two assemblies 13a, 13b intercept a portion of the gamma rays emitted along the proton beam path, namely, prompt gamma rays 21b and positron-electron annihilation gamma rays 21a. In a triple-head assembly, an additional detector assembly 13c is positioned around the target substantially facing the proton beam emitter.

[0131] Layer Identification Pixel-Configuration In a configuration where all but one column of photon sensors is shared across two or more layers 16, a sufficient amount of light should reach the pixels 18a that are used to identify which layer a scintillation event occurred in. As shown in Figures 17a-17e, these pixels 18a can be located in the corners, in the center of an edge, or anywhere in between.

[0132] Figure 21b shows a contour map of the number of photons arriving at a pixel as a function of the abscissa of the scintillation event, using fairly conservative assumptions regarding the crystal light yield (30,000 photons / MeV, gamma energy = 511 keV) and photon sensor parameters (photon detection efficiency = 0.2, dark count rate = 130,000 Hz / mm², excess noise function = 0.2, signal integration time = 250 ns). These figures assume five photon sensors (20 in total) on each edge, as shown in Figure 21a. The plot shows the light received for a single 10 mm-wide photon sensor located at the right edge (x = 25 mm) from bottom to top (y = -25 to y = +25 mm). For photon sensors located at other edges, a rotationally symmetric situation exists. Photon sensors located near the bottom corner receive almost no light for events occurring near the right edge (x ~ 25 mm, y > -15 mm) and in the upper right quadrant. A photon sensor located in the center of the edge receives approximately 50 or more photons for every event that occurs to the left (x < 0), but almost no photons for events that occur around the top-left corner. Note that the 50+ photons is the average value for the simulated events. Events with more or fewer detected photons will occur. However, even accounting for statistical fluctuations, this level is sufficient to provide a signal well above the noise floor and estimated dark count rate.

[0133] Therefore, to reliably determine which layer an event occurred in, it would be sufficient to position two layer-identifying photon sensors at the centers of two opposing edges, as shown in Figure 17e. By considering only the sum of the two pixels, or the pixel with the largest signal amplitude or integral, the layer of the scintillation event can be reliably identified. Five detector rows on two opposing edges of a four-layer square detection module, as shown in Figure 17e, require only 26 channels in this case.

[0134] Detector geometry and pixel configuration The geometrical outline of the monolithic detection module affects its performance in terms of shape, area, and photon sensor configuration. Therefore, different polygonal shapes on the front and back surfaces and different numbers of photon sensors per edge were simulated. Each simulation scenario simulated a large number of events, each representing the isotropic photon emission from a scintillation event at a randomly selected point (x0, y0, z0) in the scintillator bulk, ray tracing within the crystal and spreader material, and the response (number of detected photons) of an individual photon sensor. As previously mentioned, conservative manufacturer figures, such as dark count rate and excess noise, were used to estimate the photon sensor response, taking into account statistical variations.

[0135] The photon sensor responses from a large number of events were used to estimate the detector mean and standard deviation as a function of the abscissas x0 and y0 within the scintillator. This was the training or calibration set. Depth z0 was not part of the calibration set. Another set of events was then used as the evaluation set to investigate how well the backtracking algorithm could predict the original abscissas x0, y0 of the scintillation events based on the training set. The predicted, or fitted, coordinates were calculated as x fit , y fit The lateral error ε was then calculated as the Euclidean distance:

number

[0136] The mean error itself is not a useful indicator of whether a monolithic edge detector is a better choice than a conventional PET scanner configuration. To reduce the mean error of a conventional PET scanner, one simply needs to reduce the size of the scintillator crystals and photon sensors and increase the number of scintillators and detectors.

[0137] As a comparison index, we chose to compare the number of photon sensors of the present invention with the number of photon sensors of a "conventional PET scanner" with a similar total lateral area and similar spatial resolution (same mean error). A conventional PET scanner is defined as follows: · Individual scintillation crystals (rods) that are square and face-coupled with individual photon sensors.

[0138] To estimate the average error of the conventional configuration, simulations were performed, resulting in the error distribution shown in Figure 22. Because a single pixel crystal does not provide information about where on the single crystal a scintillation event occurred, the abscissa of each event was assigned to the center of the crystal. The largest error will occur for events occurring at the corners of the crystal. The average error is approximately equal to:

number

number

[0139] A typical example of a conventional PET scanner configuration might be a scintillation crystal with dimensions of 3.1 mm x 3.1 mm x 20 mm (3.1 mm in X and Y, 20 mm in Z) coupled to a SiPM array with SiPM pixels having dimensions of 3.3 mm x 3.3 mm. The fill factor of such a configuration, not considering gaps between modules, is approximately 88%.

[0140] A monolithic detector was evaluated. The detector was 2500 mm 2and is surrounded by a 4 mm non-scintillat- ing frame-like gap to allow space for the optical spreader and photon sensor. The fill factor of such a detector is: Triangle: 72% (Figure 9c) Square: 74% (Figure 9a) Hexagon: 76% (Figure 9b) Taking into account the non-scintillation gaps between modules, overall, the fill factor of a conventional PET scanner is believed to be similar to that of embodiments of the present invention.

[0141] Figure 24 shows an example of true versus reconstructed interaction positions for a 50x50mm scintillator plate generated from a simulation. Figure 25 shows the average lateral (azimuthal) spatial reconstruction error across the scintillator plate, and Figure 26 is a histogram of the average lateral reconstructed position error. In this example, an average lateral error of 0.93mm was achieved.

[0142] Embodiments of the present invention allow for an increase in the effective thickness of scintillation material by adding more layers without significant degradation in image quality, so that a somewhat lower fill factor compared to conventional PET scanners can be compensated for by the increased coincidence probability due to more scintillation material.

[0143] The optimization index for embodiments of the present invention can also be defined as the ratio of the number of channels in a conventional PET scanner to the number of channels in a single layer of the device.

number

[0144] Figure 23 shows the ratio R of the number of photon sensors per edge for triangular, square, and hexagonal scintillator plates. Square detectors may be the easiest to fabricate, resulting in an optimal number of pixels per side of 5, but similar gains can be achieved with 4 or 6 photon sensors per side. This ratio R is lower when considering a conventional PET scanner in which an M × M detector photon sensor array is coupled to an N × N crystal array, where N > M, for example, a monolithic light guide is inserted between the crystals and the photon sensors. A 4 × 4 detector array coupled to a 5 × 5 crystal array reduces R by 36%. Obviously, this ratio R also decreases with increasing the number of layers. However, this metric does not take into account the gain in DOI. R is valid when DOI information is not required (e.g., the thickness of the scintillator plate is the same as the height of each individual crystal in a conventional PET scanner).

[0145] Typically, for triangular, square, or hexagonal configurations, a 30-40x reduction in channel count can be achieved compared to a conventional PET scanner with one-to-one coupling. This is a significant improvement, significantly reducing the cost of the PET scanner. Alternatively, the gain in reduced channel count can be utilized to instead increase the axial field of view (FOV) of the PET scanner. This is particularly advantageous for whole-body PET scanner applications, where the axial FOV may span the entire patient.

[0146] Optimization of scintillator plate thickness for Compton cameras. The probability of a valid Compton scattering event (Compton scattering in one scintillator plate, photoelectric absorption in another) generally depends on the thickness of the scintillator plates. A detailed study of this probability with different primary gamma-ray energies and total radial scintillator thicknesses was performed for a two-layer configuration with varying scatterer and absorber scintillator thicknesses, as summarized in Figure 34. At low energies (0.511 keV), the optimal ratio is in the range of 30% to 50%, depending on the total scintillator thickness (6 to 20 mm). However, at higher energies, the optimal ratio is closer to 50%, i.e., the scatterer and absorber thicknesses are equal. Also, depending on the variant and the desired energy, a scattering layer thickness of 20% to 60% of the total thickness is preferred.

[0147] Calibration-Reference Table Calibration of the detection system may advantageously include the following steps: Known abscissa (x cal , y cal The detector is illuminated with a collimated source at the interaction depth z cal You don't need to know exactly. Record a sufficient number of events per detector layer and calibration position The mean and standard deviation of the response of each photon sensor i at a given location, i.e. μ i (x, y) and σ i Calculate (x, y). A look-up table can be generated from a fine-grid interpolation of the mean and standard deviation of each photon sensor at the calibration positions, and optionally at any intermediate positions that are not part of the calibration procedure.

[0148] μ of individual photon sensors i and σ i Alternatively / in addition, it is also possible to calculate any aggregate quantity (such as the centroid).

[0149] Reconstruction of events To reconstruct the interaction coordinates of one event, the following method can be used: Digitization of photon sensor responses and / or aggregates Find the position in the calibration table that best matches the response (various methods known per se can be used to do this efficiently).

[0150] Event Rejection In both PET scan and Compton camera modes, it is advantageous to reject events based on energy deposition. Additionally, it is also advantageous to reject double-scattered or multiple-scattered events in the same scintillator. One method is to compare the event signature (signal per readout channel) with the closest match in a lookup table. The "closest match" refers, for example, to the standard deviation-normalized Euclidean distance (sum of squared differences divided by the standard deviation of the pixel response or tally) between the event and the closest reference match. If this difference is greater than a configurable threshold, the event can be rejected.

[0151] PET scan mode vs. Compton camera mode When the detector assembly operates in PET scan mode, only events triggered on the two sensor plates opposite the source (i.e., along the LOR) are considered. In proton therapy situations, due to the potentially high trigger rate from prompt gamma rays during proton delivery, PET scan mode can optionally be completely disabled while the ion beam is being delivered. Because the intended imaging volume (defined by the treatment volume and proton beam path) is well defined (within uncertainty limits), valid coincidence groups for the detection module can be defined to discard any coincidence events outside the intended imaging volume. Additionally, one or more energy windows can be defined to reject events that do not correspond to the desired gamma energy.

[0152] In Compton camera mode, for example, only events triggered by two layers (intra-module Compton cameras) within one module or by one layer (intra-module Compton cameras) within two nearby modules can be accepted. To accept only events corresponding to known prompt gamma emission peaks, multiple energy rejection windows (where the energy is the sum of the signals from the two triggered layers) can be defined. To further limit the data rate, if one is primarily interested in the proton beam range within the target, one can enable only the module with the highest image reconstruction resolution in the direction of the proton beam.

[0153] Multiplexed readout of sensor plates aligned along the azimuth axis A major drawback of previously disclosed arrangements in which radially stacked blocks are read out in a multiplexed manner or via strip sensors is their inefficiency in detecting forward Compton-scattered gamma rays. Forward-scattered gamma rays can interact with a single sensor plate (where no scattered gamma rays are detected) or with two roughly radially aligned sensor plates. In most cases, it will be impossible to reconstruct the two interaction locations separately. Only if the Compton scattering is such that the two interactions occur in two separate modules ("inter-module scattering," low probability), or if additional scattering / absorption sensor plates with independent photon sensors or Compton pixels are added (increasing complexity), will it be possible to reconstruct the two interaction locations. In particular, at energies of interest for prompt gamma detection (up to 7 MeV), small-angle forward scattering dominates.

[0154] To overcome the inherent drawbacks of the combined readout of a radially stacked multilayer configuration, a novel multiplexed readout scheme is introduced. Rather than combining signals from compactly arranged, radially stacked sensor plates into a common readout channel (with the drawbacks previously mentioned), the inventors have realized that arranging the common readout sensor plates in a radial plane rather than in a radial stack provides significant functional and performance advantages. By arranging the common readout scintillator slab azimuthal rather than radially, several drawbacks can be overcome. A diagram of azimuthal arranged sensor plates is shown in FIG. 28.

[0155] The term radially-azimuthally-axially herein refers to the general cylindrical arrangement of scintillator elements in a PET scanner. Other arrangements are possible where the cylindrical coordinate system terminology is not directly applicable, such as spherical, "box," dual-head ( FIG. 5 e), quad-head ( FIG. 5 f), or helmet-type (e.g., for brain-specific PET scanners). However, all of these have in common that the scintillator elements are arranged in some way around the volume of interest.

[0156] One example of a multiplexed readout configuration is shown in FIG. 29 for a 2×2 sensor plate in an azimuthal axial configuration. Each sensor plate has multiple photon sensors, e.g., eight photon sensors 18. In some configurations, multiplexing circuitry 33 may perform an analog sum of the connected photon sensors 18. This allows for the sums S1–S4 of each sensor plate to be read out individually, as well as the sums E1–E8 of all photon sensors at a particular location on the sensor plate (e.g., “top right,” “top left,” etc.). The sum circuits S1–S4 allow identification of which scintillator plate 18 a scintillation event occurred on, and E1–E8 can provide spatial information about the scintillation event.

[0157] By multiplexing the photon sensor signals from azimuthal blocks, all of the benefits of reducing the number of readout channels are maintained. Advantageously, this scheme also allows for independent readout of radially separated layers. The azimuthal multiplexing arrangement is particularly well-suited for resolving Compton-scattered 511 keV gamma rays from electron-positron annihilation. If the layers are sufficiently thin, it is unlikely that a single 511 keV gamma ray will be both Compton-scattered and absorbed by the same sensor plate. This advantage is particularly important because it allows for a larger fraction of the incoming 511 keV gamma rays to be accepted. Conventional PET scanners typically only accept events with a deposited energy of 511 keV (within the instrument's energy acceptance window) to reject Compton-scattered gamma rays, and only accept events in which the 511 keV gamma ray is directly photoelectrically absorbed. One reason for this is that conventional pixelated PET scanners lack interaction depth resolution, making it impossible to unambiguously determine the initial interaction location. However, the present invention allows for the determination of both scattering and absorption coordinates. Depending on the gamma-ray energy of interest, Compton kinematic rules can be applied to resolve which scintillator plate scattering occurred and which plate absorption occurred. The time sequence can also be determined via time stamps. If the time sequence cannot be determined unambiguously, probability-weighted LOR (PET) or Compton cone (Compton camera imaging) can be used for image reconstruction. Prior knowledge of the spatial origin of the primary gamma rays can also be used to accept or reject the LOR or Compton cone.

[0158] In general, the ability to resolve Compton scattered gamma rays in PET scanning operation significantly improves the overall sensitivity of the detector, i.e., the proportion of 511 keV gamma rays that are simultaneously detected.

[0159] As an example, consider a configuration with a total radial scintillator thickness of 20 mm (e.g., LYSO) divided into five azimuthal stacked modules of 4 mm thickness. A 511 keV gamma ray impinging on this configuration has an approximately 80% probability of interacting in some way. However, it only has a 26% probability of being directly photoabsorbed. Thus, the probability of two simultaneously occurring antiparallel gamma rays being simultaneously photoabsorbed is only 0.26 * 0.26 = 6.7%. This fundamentally limits the sensitivity of conventional PET scanners.

[0160] For comparison with our azimuthal-axially multiplexed sensor plate, consider, for example, an event where E1 > 100 keV and E2 > 100 keV can be clearly resolved. In addition to the direct photoelectric absorption event, the probability that a 511 keV gamma ray is first Compton scattered and then photoelectrically absorbed in a radially different layer is 11.4%. Therefore, the probability of detecting a 511 keV gamma ray via direct photoelectric absorption or as a two-step Compton event (scattering + absorption) is 26% + 11.4% = 37.4%. The probability of coincidence detection is 0.374^2 = 14%.

[0161] Thus, embodiments of the present invention can have an overall sensitivity, or effective coincidence detection rate, that is approximately twice as high as conventional detectors. Half the number of emitted gamma rays is required to form an image of equivalent quality. Alternatively, in the context of radiopharmaceuticals, the amount of injected tracer isotope can be significantly reduced to minimize patient radiation dose.

[0162] Hybrid signal readout of edge-adjacent photon sensors Commonly available photon sensors, such as silicon photomultiplier tubes, are typically square in shape, e.g., 1 x 1, 3 x 3, 4 x 4, or 6 x 6 mm, and the exact dimensions of the scintillator plates of the present invention can be adapted to this. In a variant, the desired scintillator plate dimensions may be, for example, 48 x 48 x 3 mm (3 mm being the radial thickness). To match the radial thickness, a 3 x 3 mm photon sensor would be suitable. In a variant of the present invention, the number of photon sensors per edge should be at least 2, at least 3, at least 4, at least 5, at least 6, at least 7, or at least 8. However, a 3 x 3 mm photon sensor would correspond to 16 channels per edge. In an advantageous embodiment of the present invention, signals from adjacent photon sensors can be electronically summed, and groups of adjacent photon sensors are read out together or connected together to a multiplexing circuit before digitization. Figure 40 shows an example of a 4:1 reduction in the number of channels per edge using common anode current summation. Other signal summing techniques are known per se and address issues such as increasing the effective sensor capacitance and may be used in embodiments of the present invention.

[0163] Axial field of view expansion In an advantageous embodiment, at least two azimuthally opposed sensor plate arrangements can rotate around the patient or scan subject to acquire PET and / or SPECT images of the entire body or part of the body, as shown in Figure 38. The detectors may be combined with CT imaging equipment. The detectors may be combined with MRI. The detectors can acquire dynamic images ("4D" images with volumetric and temporal information).

[0164] A major limitation of many conventional PET scanners is their limited axial field of view. Increasing the FOV generally requires adding more photon sensors, readout channels, and scintillator material, resulting in a linear increase in manufacturing costs. While cost can be partially reduced by reducing the radial scintillator thickness, this does not reduce the scintillator area that must be covered by photon sensors. As an example, consider a source with an axial extent similar to that of a conventional PET scanner, as conceptually shown in Figure 31. Simultaneous detection of annihilation gamma rays emitted near both ends of the source requires axial movement of the scanner or the scanned object, lengthening imaging time and making imaging of dynamic processes along the source difficult. However, the axial FOV of our invention can be increased with minimal or no additional cost by simply repositioning the sensor plate axially, as shown in Figure 30. While the total radial scintillator thickness decreases, resulting in a reduced coincidence probability, this is compensated for by the increased FOV. Furthermore, the azimuthal multiplexing configuration significantly improves overall sensitivity through its ability to accommodate Compton scattering events. A comparison of axial sensitivity is shown in Figure 32.

[0165] As shown in Figure 36, changing the sensor plate configuration from one radial layer to two or more radial layers can actually increase the probability of coincidence detection, even if the total scintillator thickness is reduced. For example, the probability of effective simultaneous absorption for a single-layer configuration with a radial scintillator thickness of 20 mm is similar to the probability of effective simultaneous absorption for a two-layer configuration with a total thickness of 15 mm, or a four- to six-layer configuration with a total thickness of 10 mm. Therefore, it is an advantage of the present invention that the total scintillator volume and the total area required to be covered by the photon sensor can be significantly reduced, thereby reducing overall cost and / or enabling an increase in the axial FOV.

[0166] The radial scintillator thickness, distributed over at least two scintillator plates, can be, for example, less than 40 mm, less than 30 mm, less than 20 mm, less than 15 mm, or less than 10 mm.

[0167] In an area-coupled detector module, a radial scintillator thickness of about 20 mm is typically considered optimal for cost-effective PET scanning, considering the probability of coincidence detection. In an advantageous embodiment of the present invention, the radial scintillator thickness can be reduced, for example, to less than 19 mm, since Compton scattered gamma rays can also be accepted as valid events.

[0168] In advantageous embodiments, the azimuthally multiplexed configuration may also be adapted to include more than one radial gap, as shown in Figure 33, which shows a configuration with 2+2+2 radially stacked scintillator blocks multiplexed in a 3x3 fashion. Such a configuration may function as a PET scanner, a two-stage Compton camera, and a three-stage Compton camera. Radial air gaps between sensor plates may also improve heat dissipation.

[0169] The multiplexed sensor plates can be arranged in an azimuthal axial array in a 1x2, 1x3, 1x4, 2x2, 2x3, 2x4, 3x3, 3x4, 4x4, 4x5, or 5x5 fashion, for example.

[0170] Function combinations and uses The features of different embodiments are interchangeable between the embodiments and can be combined in various ways, unless otherwise indicated. Although numerous specific details are set forth in the following description to provide a more thorough understanding of the present invention, it will be apparent to those skilled in the art that the present invention can be practiced without these specific details. Because the basic and conventional techniques in electronics, sensor systems, image analysis, signal processing, data communication systems, image acquisition systems, and other components for implementing the present invention are believed to be readily apparent to those skilled in the art, further explanation and details are omitted in this description for the sake of brevity.

[0171] Detection systems according to embodiments of the present invention can also be used for other types of nuclear imaging, for example, imaging of 3-gamma emissions, or one positron-emitting isotope (which produces a coincident, opposing 511 keV gamma which, when detected, gives the LOR 27 along which the source is located), and another gamma (which produces the Compton cone 25 along which the source is located). Combining information from the LOR and Compton cone allows possible isotope locations to be triangulated to a very high degree of accuracy, particularly in the context of ion beam therapy where the volume of interest is known, as shown in FIG.

[0172] WO 2018 / 081404 A1 discloses a radially stacked edge-detection detector with an individual photon sensor for each layer. A single layer significantly reduces the number of channels compared to a one-to-one primary surface-coupled PET scanner. However, the number of channels increases linearly with the number of layers. According to one aspect of the present invention, a combination of strip photon sensors across multiple layers and pixel photon sensors coupled to a single layer significantly reduces the number of channels even in a multi-layer detector, while still allowing the pixel detector to identify which layer a scintillation event occurred in.

[0173] According to another aspect of the invention, a double-ended strip photon sensor can be used, and the time difference between the two ends can be used to determine at which layer a scintillation event occurred. In this case, the total number of readouts is independent of the number of layers in the module. This solution has the advantage that layer-specific pixels are not required. The interaction depth resolution is limited only by the timing accuracy and the number of scintillator plates.

[0174] According to another aspect of the present invention, a detection system uses adjacent modules as scatterer / absorber modules of a Compton camera ("inter-module Compton camera"). Due to the interaction depth resolution, this configuration is more robust to parallax errors caused by gamma rays entering the detector at an angle. Therefore, a non-circular arrangement of detection modules around the scanned object, such as a hexagon, octagon, or other polygon, increases the probability of detecting forward-scattered gamma rays in two different modules compared to a conventional circular arrangement.

[0175] Another aspect of the present invention provides an electro-optical shutter for temporarily blocking optical signals from selected layers. This is particularly advantageous when stacked detectors are readout with strip detectors spanning multiple layers. The electro-optical shutter can be used, for example, as a means to functionally convert some of the layers radially closest to the imaging target into "gamma filters." This is a useful feature in situations where the instantaneous rate of prompt gammas is so high that optically activating all layers would saturate the detector. By temporarily blocking light from the layers closest to the target, these proximal layers serve to absorb some of the prompt gammas without blinding the detector, thereby reducing the overall count rate of the detector.

[0176] According to another aspect of the invention, an arrangement of conceptually identical modules (each consisting of a scintillator plate, photon sensor, and readout electronics) can be arranged to function as a combined PET scanner and Compton camera, with radially offset groups of modules functioning as absorbers in a two-stage Compton camera. This configuration is advantageous from a manufacturing and cost standpoint, and is also easily customizable.

[0177] According to another aspect of the invention, an azimuthal axial arrangement of sensor plates with multiplexed readout is provided, which is particularly advantageous for detecting Compton scattered gamma rays for both PET scanning and Compton camera functionality over a large solid angle around the volume of interest, and also reduces parallax errors.

[0178] [Prior art documents] [1]CN107544086 A [2] WO2018 / 081404 A1 [3] K. Shimazoe et.al, 2020, Nuclear Inst. and Methods in Physics Research, A: https: / / doi.org / 10.1016 / j.nima.2018.10.177 [4]EP1617237 A1 [5]US2018 / 172847 A1 [6]US2005 / 116173 A1 [7]Georgy Shakirin et al 2011 Phys. Med. Biol. 56 1281 (Shakirin2011): https: / / doi.org / 10.1088 / 0031-9155 / 56 / 5 / 004 [8]Rohling et al, 2017, Phys. Med. Biol., at press (Rohling2017): https: / / doi.org / 10.1088 / 1361-6560 / aa6068 [9]Antje-Christin Knopf and Antony Lomax 2013 Phys. Med. Biol. 58 R131 (Knopf2013)

[10] Jan et al, 2017, Med. Phys. 44 (12), December 2017 (Jan2017) (https: / / doi.org / 10.1002 / mp.12626)

[11] K. Doroud, MCS Williams, K. Yamamoto (Doroud2017) The Strip Silicon Photo-Multiplier:An innovation for enhanced time and position measurement, Nuclear Instruments and Methods in Physics Research Section A: Accelerators, Spectrometers, Detectors and Associated Equipment, Volume 853, 2017, Pages 1-8, ISSN 0168-9002,

[0179] [List of features referenced in the drawings] patient 5 Target zone (e.g., tumor) 4 Ion beam therapy system 6 (e.g., proton therapy system) patient support 7 Ion Beam Emitter 8 Ion Beam 1 Scan Magnet 2 Beam Intensity and Profile Monitor 3 Gamma ray detection system 10 Compton Camera 11 PET scanner 12 Detection module assembly 13, 13a, 13b, 13c Opening 42 Detection modules 14, 14a, 14s Scattered parts / block 15s Absorbent part / Block 15a Sensor plates 16 and 18 Scintillator Plate 16 scattering layer 16s Absorber layer 16a Main surface 40a Lateral minor surface 40b (also called "edge") Scintillator rod 16p Radial gap 17 Photon Sensor 18 Individual layer photon sensors 18a (also called "photon sensor pixels" or simply "pixels") Photon sensor 18p coupled to scintillator rod 16p Crosshair connection arrangement 18c Strip multilayer photon sensor 18b (also called a "photon sensor strip" or simply a "strip detector") Photon sensor support (substrate) 20 Detector-Scintillator Optical Interface 22 Electro-Optical Shutter (EOS) 24 Edge Light Spreader 26 Interlayer reflector 28 Light Partial Barrier / Absorber 29 Low index gap 31 Signal Processing and Control Systems 30 Circuit board 32 Multiplexing circuit 33 Electronic components 34 (e.g., microprocessors, memories, FPGAs, etc.) Connectors 36a, 36b Gamma Ray 21 Positron gamma ray 21a Prompt Gamma Ray 21b Source 23 Compton Corn 25 Line of Response (LOR) 27 Compton Cone-LOR Intersection 27b Desired volume (target zone) 27c Scintillation Ray 53 Main surface coupled detection module 50 Scintillator Array 51 Photon Sensor Array 52

[0180] [Embodiment] (1) A gamma ray detection system (10) including a detection module assembly (13, 13a, 13b, 13c) including at least two detection modules (14, 14a, 14s) configured for positron emission tomography (PET) scanning of a target zone (4), each detection module comprising a plurality of stacked monolithic scintillator plates (16), each having a major surface (40a) oriented generally facing the target zone and minor lateral surfaces (40b) defining edges of the scintillator plate, the major surfaces having a surface area greater than the surface area of the minor lateral surfaces. a monolithic scintillator plate (16) mounted on each of the edges and a plurality of photon sensors (18) mounted against each of the edges and configured to detect and determine the position within the plane of the major surface of a scintillation event in the scintillator plate from gamma rays incident on the major surface, wherein the gamma ray detection system is further configured to function as a Compton camera, and at least one scintillator plate that is not the scintillator plate closest to the target zone is configured as an absorber scintillator plate for the Compton camera. (2) A gamma ray detection system as described in embodiment 1, wherein the plurality of photon sensors of at least two radially stacked scintillator plates are connected to a processing circuit configured to multiplex the readout of the plurality of photon sensors. (3) A gamma ray detection system as described in embodiment 1, wherein the plurality of photon sensors of the scintillator plates arranged in at least two azimuthal axis directions are connected to a processing circuit configured to multiplex the readout of the plurality of photon sensors. (4) A gamma ray detection system as described in embodiment 1, comprising at least one radial gap (17) between at least two of the plurality of stacked scintillator plates or between at least two detection modules. (5) A gamma ray detection system as described in embodiment 4, wherein the height H of the radial gap relative to the thickness T of one of the plurality of scintillator plates may typically be in the range of 200 > H / T > 2, preferably in the range of 50 > H / T > 10.

[0181] (6) A gamma ray detection system as described in embodiment 1, wherein the plurality of photon sensors includes at least one strip multi-layer photon sensor (18b) extending over an edge of the plurality of layers. (7) A gamma ray detection system as described in embodiment 6, comprising a plurality of the strip multilayer photon sensors on each edge side of the plurality of stacked scintillator plates. (8) A gamma ray detection system as described in embodiment 6, wherein the at least one strip multilayer photon sensor is a double-ended strip detector configured to measure the arrival time of a signal at both ends. (9) A gamma ray detection system as described in embodiment 1, wherein the plurality of photon sensors includes at least one individual layer photon sensor (18a) on at least one edge of each scintillator plate, preferably on at least two edges of each scintillator plate. (10) A gamma ray detection system as described in embodiment 9, wherein the individual layer photon sensors in rows and / or columns are interconnected in a cross-connection arrangement (18c), and the readout is the sum and / or weighted sum of the signals of multiple interconnected individual layer photon sensors.

[0182] (11) The gamma ray detection system of claim 1, further comprising a light-reflective (28) or light-absorbing (29) interface layer between at least two of said scintillator plates. (12) A gamma ray detection system as described in embodiment 1, further comprising a low refractive index gap, e.g., air, between at least two of the scintillator plates. (13) The gamma ray detection system of embodiment 1, further comprising an electro-optical shutter (24) between the edge of at least one scintillator plate and the photon sensor. (14) A gamma ray detection system as described in embodiment 13, wherein the electro-optical shutter includes a light spreader material and thickness configured to spread light from a scintillation event near the edge. (15) The surface area S of the main surface of the scintillator plate and the thickness T of the scintillator plate are 100 mm 2 ≦S≦40000mm 2 and 0.5 mm≦T≦30 mm.

[0183] (16) A gamma ray detection system as described in embodiment 1, wherein the detection module assembly surrounds a target zone and includes a gap (42) or orifice for emitting an ion beam. (17) A gamma ray detection system as described in embodiment 1, wherein the radial gap satisfies the relationship H / (T1+T2)>5, where T1 and T2 are the thicknesses of the two scintillators surrounding the radial gap, and H is the height of the radial gap. (18) A gamma ray detection system as described in embodiment 1, wherein the total thickness of the plurality of stacked monolithic scintillator plates in the radial direction is less than 19 mm. (19) A gamma ray detection system as described in embodiment 1, comprising two radially stacked scintillator plates, the ratio of the thickness of the radially inner scintillator plate to the total thickness of the scintillators in the radial direction being in the range of 0.2 to 0.6. (20) A gamma ray detection system as described in embodiment 1, wherein the photon sensor bias voltages of the photon sensors of each scintillator plate can be independently adjusted or enabled / disabled.

[0184] (21) A gamma ray detection system as described in embodiment 1, wherein the photon sensor coupled to at least two radially stacked scintillator plates is connected to a processing circuit configured to apply the Compton kinetic law to determine whether two coincident blocking events correspond to forward or backward scattered Compton scattering followed by absorption. (22) A gamma ray detection system as described in embodiment 21, wherein the processing circuit is configured to reject events that are likely to be caused by primary gamma rays incident on the detector from a radially outward direction. (23) A gamma ray detection system as described in embodiment 21, wherein the processing circuit is configured to use the interaction coordinate of the photoelectric absorption as the LOR endpoint of the small-angle forward Compton scattered gamma ray resulting from electron-positron annihilation. (24) A gamma ray detection system described in any of embodiments 21 to 23, wherein the processing circuit is configured to discard Compton scattering events beyond a configurable primary gamma ray energy-dependent scattering angle to improve angular resolution. (25) A gamma ray detection system as described in embodiment 1, wherein analog signals from adjacent photon sensors are added prior to digitization or other multiplexing circuitry.

[0185] (26) An ion beam therapy system (6) for irradiating a tissue zone with an ion beam, comprising: a patient support (7) and an ion beam emitter (8) that are relatively movable about at least a rotation axis; and a gamma ray detection system according to any one of embodiments 1 to 25 configured to perform prompt gamma ray detection and PET scanning during, between, and after ion beam irradiation.

Claims

1. 1. A gamma ray detection system (10) comprising a detection module assembly (13, 13a, 13b, 13c) including at least two detection modules (14, 14a, 14s) configured for a positron emission tomography (PET) scan of a target zone (4), each detection module comprising: a signal processing and control system (30) including a circuit board (32) and electronic components (34) mounted thereon; a plurality of stacked monolithic scintillator plates (16), each having a major surface (40a) oriented generally facing the target zone and minor lateral surfaces (40b) defining edges of the scintillator plate, the major surfaces having a surface area greater than the surface area of the minor lateral surfaces; and a plurality of photon sensors (18) attached to each of the edges configured to detect and determine the position within the plane of the major surfaces of scintillation events within the scintillator plate from gamma rays incident on the major surfaces, The gamma ray detection system is further configured to function as a Compton camera, wherein at least one scintillator plate other than the scintillator plate closest to the target zone is configured as an absorber scintillator plate for the Compton camera, the circuit board (32) is attached to the outermost radial end of the detection module, the plurality of photon sensors (18) are provided on a photon sensor support substrate (20), and the photon sensor support substrate includes an edge connector (36b) arranged on a side along the thickness direction of the photon sensor support substrate that is connected to a connector of the circuit board (32) of the signal processing control system (30).

2. 10. The gamma ray detection system of claim 1, wherein the photon sensor support substrate (20) is configured as a silicon photomultiplier array substrate.

3. 10. The gamma ray detection system of claim 1, wherein the plurality of photon sensors of the at least two radially stacked scintillator plates are connected to processing circuitry configured to multiplex a readout of the plurality of photon sensors.

4. 2. The gamma ray detection system of claim 1, wherein a plurality of photon sensors of at least two azimuthal aligned scintillator plates are connected to processing circuitry configured to multiplex the readout of the plurality of photon sensors.

5. 2. The gamma ray detection system of claim 1, comprising at least one radial gap (17) between at least two of the plurality of stacked scintillator plates or between the at least two detection modules.

6. 6. The gamma ray detection system of claim 5, wherein the height H of the radial gap relative to the thickness T of one of the scintillator plates is in the range 200 > H / T > 2.

7. 2. The gamma ray detection system of claim 1, wherein the plurality of photon sensors includes at least one strip multi-layer photon sensor (18b) extending over an edge of the plurality of layers and individual layer photon sensors (18a) positioned on each of the scintillator plates.

8. 8. The gamma ray detection system of claim 7, comprising a plurality of said strip multilayer photon sensors on each edge side of a plurality of said stacked scintillator plates.

9. 8. The gamma ray detection system of claim 7, wherein the at least one strip multi-layer photon sensor is a double-ended strip detector configured to measure the arrival time of a signal at both ends.

10. 2. The gamma ray detection system of claim 1, wherein the plurality of photon sensors includes at least one individual layer photon sensor (18a) on at least one edge of each scintillator plate or on at least two edges of each scintillator plate.

11. 11. The gamma ray detection system of claim 10, wherein the individual layer photon sensors of the rows and / or columns are interconnected in a cross-connection arrangement (18c), and the readout is the sum and / or weighted sum of the signals of a plurality of interconnected individual layer photon sensors.

12. 10. The gamma ray detection system of claim 1, further comprising a light-reflective (28) or light-absorbing (29) interface layer between at least two of said scintillator plates.

13. The gamma ray detection system of claim 1 further comprising a low refractive index gap between at least two of said scintillator plates.

14. 10. The gamma ray detection system of claim 1, further comprising an electro-optical shutter (24) between said edge of at least one scintillator plate and said photon sensor.

15. 15. The gamma ray detection system of claim 14, wherein the electro-optic shutter includes a light spreader material and thickness configured to spread light from scintillation events near the edge.

16. The surface area S of the main surface of the scintillator plate and the thickness T of the scintillator plate are 100 mm ２ ≦S≦40,000 mm ２ 10. The gamma ray detection system of claim 1, wherein T is in the range of 0.5 mm≦T≦30 mm.

17. 2. The gamma ray detection system of claim 1, wherein the detection module assembly surrounds the target zone and includes a gap (42) or orifice for emitting an ion beam.

18. 6. The gamma ray detection system of claim 5, wherein the radial gap satisfies the relationship H / (T1+T2)>5, where T1 and T2 are thicknesses of the two scintillator plates surrounding the radial gap, and H is the height of the radial gap.

19. 2. The gamma ray detection system of claim 1, wherein the total radial thickness of the plurality of stacked monolithic scintillator plates is less than 19 mm.

20. 10. The gamma ray detection system of claim 1, comprising two radially stacked scintillator plates, wherein the ratio of the thickness of the radially inner scintillator plate to the total radial scintillator thickness is in the range of 0.2 to 0.

6.

21. 10. The gamma ray detection system of claim 1, wherein photon sensor bias voltages of the photon sensors of the individual scintillator plates can be independently adjusted or enabled / disabled.

22. 10. The gamma ray detection system of claim 1, wherein the photon sensors coupled to the at least two radially stacked scintillator plates are connected to processing circuitry configured to apply the Compton kinetic law to determine whether two coincident blocking events correspond to a forward or backscattered Compton scatter followed by absorption.

23. 23. The gamma ray detection system of claim 22, wherein the processing circuitry is configured to reject events that are attributed to primary gamma rays incident on the gamma ray detection system from a radially outward direction.

24. 23. The gamma ray detection system of claim 22, wherein the processing circuitry is configured to utilize an interaction coordinate of photoelectric absorption as an LOR endpoint for small angle forward Compton scattered gamma rays resulting from electron-positron annihilation.

25. 25. A gamma ray detection system as described in any one of claims 22 to 24, wherein the processing circuitry is configured to discard Compton scatter events beyond a configurable primary gamma ray energy dependent scatter angle to improve angular resolution.

26. 10. The gamma ray detection system of claim 1, wherein analog signals from adjacent photon sensors are added together prior to digitization or other multiplexing circuitry.

27. 27. An ion beam therapy system (6) for irradiating a tissue zone with an ion beam, comprising: a patient support (7) and an ion beam emitter (8) that are relatively movable about at least a rotation axis; and a gamma ray detection system according to any one of claims 1 to 26 configured to perform prompt gamma ray detection and PET scanning during, between, and after ion beam irradiation.

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