Transmission imaging in a forward scatter gamma ray based pet scanner in coincidence detection

Abstract

A novel method is disclosed for obtaining transmission scan data in a PET scanner by incorporating one or more stationary gamma ray sources that provide forward scattered gamma photons that can be used as transmission imaging radiation.

Description

[0001] Cross-reference to related applications

[0002] This application claims the benefit of U.S. Provisional Application No. 63 / 018,654, filed May 1, 2020, the disclosure of which is incorporated herein by reference in its entirety. Technical Field

[0003] This disclosure relates to nuclear imaging. More particularly, this disclosure relates to attenuation correction in positron emission tomography (PET). Background Technology

[0004] PET is a nuclear medicine imaging technique that produces three-dimensional images representing the distribution of positron-emitting isotopes within the body. When a radioactive isotope undergoes positron-emitting decay (also known as positive beta decay), it emits an antimatter counterpart of an electron. As the positrons lose energy, they eventually meet and annihilate with an electron, typically producing a pair of annihilation (gamma) photons moving in opposite directions. PET systems determine the line along which annihilation occurs by detecting a pair of gamma photons that are in temporal overlap.

[0005] Attenuation correction in PET imaging is a crucial component for generating artifact-free quantitative data. Most attenuation correction methods are based on measurements of transmission scans acquired before, during, or after the emission scan. The ability to acquire high-quality transmission data is beneficial for achieving accurate attenuation correction.

[0006] Attenuation in PET imaging is due to the absorption of true coincident events within the body or the loss of detection due to their scattering outside the detector's field of view (FOV). In PET imaging, two photons must escape from the patient simultaneously to be detected as a true event. The loss in the detection of true coincident events due to attenuation in PET imaging can range from 50% to 95%.

[0007] Attenuation-induced count loss increases image noise, image artifacts, and image distortion. Significant artifacts that may occur on whole-body PET scans without attenuation correction include: (1) prominent activity at the edges of the body surface due to the relative lack of attenuation at the surface compared to deeper structures; (2) distorted appearance of areas of intense activity (e.g., the bladder) due to variable attenuation levels in different directions originating from these areas; and (3) diffuse, relatively increased activity in tissues with relatively low attenuation (e.g., the lungs). Therefore, attenuation correction is necessary for accurate measurement of PET scan data.

[0008] In PET / CT systems, X-rays from CT scans are used to construct an attenuation map of density differences across the target imaging region. This attenuation map can then be used to correct for the absorption of photons emitted from the decay of fluorinated deoxyglucose in emission scans. However, integrating a PET / CT system in this way requires additional CT scanning hardware to integrate with the PET imaging hardware. Therefore, it would be useful if transmission-type scan data could be obtained using a PET scanner without additional transmission scanning hardware (such as a CT scanner) to construct the attenuation map. This ability to obtain transmission-type scan data for attenuation map generation in a PET scanner system would also be useful in PET / MR systems, since MR systems do not measure everything in that system that causes radiation attenuation in PET scans. For example, PET / MR systems use coils and associated electronics and cables that attenuate annihilated photons, but by design they are ignored so that they do not appear in MR images, and their effect on annihilated photon attenuation cannot be accounted for.

[0009] Several PET scanner designs have been proposed in which attenuation correction is derived from the background radiation emitted from the lutetium-based scintillation crystal (e.g., LSO or LYSO scintillation crystal) within the detector itself. However, the intensity of LSO background radiation is generally too low to be used in patient scans of normal duration. Therefore, an improved PET scanner that can generate transmission imaging data more quickly to produce attenuation μ maps without CT hardware would be beneficial. Summary of the Invention

[0010] A PET scanner system is provided, comprising a gantry; a plurality of PET detector ring assemblies provided within the gantry, wherein each detector ring assembly includes a plurality of PET detectors arranged in a ring structure around a central opening, wherein the plurality of PET detector ring assemblies are coaxially aligned along a longitudinal axis defined by the central opening; a patient tunnel extending through the central opening of the coaxially aligned PET detector ring assemblies, wherein the plurality of PET detector assemblies are coaxially aligned along the length of the patient tunnel, wherein each of the plurality of PET detectors includes a detector and one or more scintillator crystals associated with the detector; and one or more stationary gamma-ray sources provided in each PET detector ring assembly within the gantry.

[0011] A method is also provided for using one or more stationary gamma-ray sources provided in a PET scanner as transmission scanning radiation sources to generate scan data that can be used to generate attenuation maps, wherein the PET scanner includes a plurality of PET detector ring assemblies. The method includes: (a) providing one or more stationary gamma-ray sources in each PET detector ring assembly, wherein the stationary gamma-ray sources are positioned outside the detector ring assembly; (b) identifying forward-scattered gamma photons emitted from the stationary gamma-ray sources, which have been forward-scattered by a scintillator crystal in a first detector block assembly in the PET detector ring assembly, across the field of view (FOV) of the PET scanner, and detected by a scintillator crystal in a second detector block assembly in the PET detector ring assembly via the PET scanner's reconciliation electronics; (c) acquiring list pattern data from a blank transmission scan inactive in the FOV; (d) acquiring list pattern data from a transmission scan having a target body in the FOV; (e) generating an attenuation map by comparing the list pattern data from the blank transmission scan with the list pattern data from the transmission scan having a target body in the FOV; and (f) applying the attenuation map to the list pattern data from the emission scan in step (e) to attenuate the list pattern data from the emission scan.

[0012] A method is also provided for improving the quality of a transmission scan signal from forward-scattered gamma photons by taking time-of-flight (TOF) into account when applying transmission imaging scans using a stationary gamma-ray source, comprising: (a) calculating the TOF of scattered photons arriving at a second scintillator crystal in a second detector block based on the distance between two scintillator crystals, the scattered photons being photons from a gamma-ray source that have undergone Compton scattering in a first scintillator crystal in a first detector block; (b) defining a time window having a width centered on the calculated TOF; (c) measuring the TOF of actual scattered gamma photons originating from the first scintillator crystal using a scanned target object in the FOV of a PET scanner; (d) comparing the measured TOF from (c) with the calculated TOF and identifying the measured TOF within the time window; and (e) identifying scattered gamma photons corresponding to those measured TOFs within the time window as transmission source events originating from Compton scattering in the first scintillator crystal, thereby distinguishing transmission type data from gamma emission annihilation events and random events within the scanned target object. Attached Figure Description

[0013] Figure 1A This is a schematic diagram of a PET system.

[0014] Figure 1BThis is a schematic diagram of a PET detector ring assembly with a stationary gamma ray source according to an embodiment of the present disclosure.

[0015] Figure 1C This is a schematic diagram of a PET detector ring assembly having multiple stationary gamma-ray sources according to an embodiment of the present disclosure.

[0016] Figure 2 This is a schematic diagram of a gamma-ray source located on the back side of a PET detector and provided in an "ON" configuration within a tungsten radiation shield, where gamma rays from the source illuminate the detector's scintillation crystal.

[0017] Figure 3 This is a schematic diagram of a gamma-ray source located on the back side of a PET detector and provided in an "OFF" configuration within a tungsten radiation shield, which is absorbing gamma radiation from the source.

[0018] Figure 4A This is a flowchart of a method according to an embodiment of the present disclosure.

[0019] Figure 4B This is a flowchart of a method according to another embodiment of the present disclosure.

[0020] Figure 5A and 5B These are schematic diagrams of the front and side views of the PET scanner used in the experimental verification, showing the locations of different radioactive sources involved in the measurement.

[0021] Figure 6 These are 2D energy histograms from scans 1, 2, 3, and 4, which were performed during experimental verification.

[0022] Figure 7 This is a grayscale representation of the distribution of single-particle rate in 228 detector blocks of a 6-ring Biograph Vision PET / CT scanner used in experimental validation.

[0023] Figure 8 The net-truesinograms from scans 1, 4, and 5, which were performed in the experimental validation, are shown.

[0024] Figure 9 The diagram S shows the sine wave of the reorganized compartments. cs1-blank / S bg-blank A plot of the ratio.

[0025] Figure 10 This is a diagram of the estimated 38-source sine curve. Detailed Implementation

[0026] The description of the exemplary embodiments is intended to be read in conjunction with the accompanying drawings, which are to be considered an integral part of the entire written description.

[0027] PET emission data collected in conventional PET scans includes all information about the physical effects experienced by emitted photons before detection in the detector. To reconstruct the collected emission data, corrections are applied to reconstruct the true measured emission events. Systematic corrections such as normalization, randomization, and dead-time correction are object-independent and depend primarily on the systematic and count rates of the activity illuminating the PET detectors inside and outside the FOV. Other corrections are object-dependent, such as attenuation correction and scattering estimation, and require attenuation information about the object material within the scanner's FOV.

[0028] [General Operation of a PET Scanner]

[0029] Figure 1A An example of a PET scanner system 200 in which the inventive concepts disclosed herein can be implemented is shown. A human subject 4 for PET scanning is shown positioned within a gantry 210 of the PET scanner system 200. The gantry 210 includes a plurality of PET detector rings 100. Each detector ring includes a plurality of scintillation crystals 216 and associated detectors 213. An open space at the center of the detector rings 100 defines the patient tunnel T of the gantry 210 and the field of view (FOV) of the PET scanner. During a PET scan, positron-emitting radioisotope 6 is introduced into the human subject 4 via metabolically active molecules. These metabolically active molecules are then carried to the organ of interest via the bloodstream of the human subject.

[0030] When a positron emitted from a radioactive isotope currently within a human subject encounters an electron, the two are annihilated, producing two gamma photons 7 traveling in approximately opposite directions. The annihilation event is identified by the interaction of the gamma photon with the detector scintillation crystal 216 by two opposing detectors, through the time coincidence between the detection of the two gamma photons; that is, the gamma photon emission is actually detected simultaneously by each detector 213. When the two opposing gamma photons collide with their corresponding opposing detectors to produce a time coincidence event, these photons identify the annihilation event along the response line (LOR) where it occurs.

[0031] Images (nuclear medicine images) of metabolic activity in human subject 4 are reconstructed through computer analysis. The PET scanner system 200 includes a system controller 290 connected to and in communication with the detector ring 100. The PET scanner system 200 further includes a data processing unit (event detection unit) 220 that identifies and evaluates coincidence events generated by gamma ray pairs and forwards this information to an image processing unit (computation unit) 230. The detector pair associated with each LOR generates numerous coincidence events during a measurement session (i.e., a data acquisition scan). The PET scanner system 200 further includes at least one machine-readable storage medium 250 encoded with computer program code, which, when executed by the system controller 290, performs various operational functions of the PET scanner system 200.

[0032] [Improvements provided in this disclosure]

[0033] refer to Figure 1B According to this disclosure, a novel configuration for a PET scanner system is provided that allows for the simultaneous generation of transmission scan data during a PET scan session. To provide the PET scanner with a means of generating simultaneous transmission scan data during a PET scan session, one or more stationary gamma-ray sources 130 are incorporated into each PET detector ring assembly 100 within a gantry 210. Each PET detector ring assembly 100 includes a plurality of detector electronics assemblies 110 arranged in a ring configuration around the patient tunnel T. Each detector electronics assembly 110 includes, among other things, a detector 213 (photodetector) and one or more associated scintillator crystals 216. Figure 1B This is illustrative, and for the sake of simplicity, detector 213 and the associated one or more scintillator crystals 216 are referred to as a group as PET detector block 214. This designation will be used throughout the remainder of this disclosure.

[0034] A stationary gamma-ray source 130 is located within the PET gantry 210 and is positioned close to but behind the PET detector block 214, away from the patient tunnel T, such that some of the gamma photons from the gamma-ray source are forward-scattered by a first scintillator crystal associated with the first PET detector block 214a, and then across the FOV of the PET scanner to a second scintillator crystal associated with a second PET detector block 214b located on the opposite side of the FOV of the PET scanner. For practical reasons, the gamma-ray source 130 is preferably positioned within the PET scanner gantry 210 to shield and contain the gamma rays within the PET scanner.

[0035] exist Figure 1BIn the diagram, arrow 11 represents an exemplary gamma photon forward-scattered by the first scintillator crystal of the first detector block 214a, and arrow 12 represents the trajectory of a scattered gamma photon that traverses the FOV of the PET scanner and reaches the second scintillator crystal of the second detector block 214b. The scattered gamma photon is identified by 12. Because the forward-scattered gamma photon 12 traverses the FOV, a sufficient number of such forward-scattered gamma photons 12 can be used for transmission imaging in the PET scanner by identifying a sufficient number of them via the PET scanner's reconciliation electronics.

[0036] In some preferred embodiments, one or more stationary gamma-ray sources 130 in each detector ring assembly 100 are positioned behind the detector block 214. This configuration allows more gamma photons emitted from the gamma-ray source 130 to be forward-scattered by scintillator crystals in one or more nearby detector blocks 214, thereby increasing the amount of radiation used for transmission imaging.

[0037] refer to Figure 1C By providing multiple stationary gamma-ray sources 130 in each PET detector ring assembly 100, it is more practical to provide and identify a sufficient number of forward-scattered gamma photons for transmission imaging. All gamma-ray sources 130 are located behind the PET detector block 214, as shown below. Figure 1B As illustrated in the diagram. "Located behind the PET detector block 214" refers to any location on the side of the detector block 214 facing away from the patient tunnel T. The gamma source 130 can be positioned anywhere within the gantry, as long as it is behind the PET detector block 214. However, in some preferred embodiments, each stationary gamma source 130 is positioned aligned between a pair of detector electronics assemblies 100, i.e., between two PET detector block assemblies 214, as shown in the diagram. Figure 1B and 1C As shown, this is to maximize the detector's detection of forward-scattered photons. If the gamma-ray source is directly behind the detector, this corresponds to zero-degree scattering, which cannot be detected because in this case, the scattering process leaves almost no energy in the detector where it is scattering. The scattering process is undetectable until the energy left during scattering is higher than a threshold of the detector electronics (the specific threshold varies depending on the PET scanner manufacturer, but an example is approximately 150 keV), so the most favorable scattering angle is approximately 40 degrees or greater. This can be achieved by positioning the stationary gamma-ray source 130 behind the PET detector block 214 and between a pair of detector electronics assemblies 110. When the gamma-ray source is positioned behind the PET detector block 214 and between a pair of detector electronics assemblies 110, both detector blocks 214 on either side are well-positioned for scattering.

[0038] All stationary gamma-ray sources 130 will emit some gamma photons that are forward-scattered through the FOV, thus providing more radiation necessary for transmission imaging. Figure 1C This is merely an illustrative illustration, and the number of detector electronics, detector block 214, and gamma source 130 illustrated does not represent the actual number in an actual PET scanner system employing embodiments of this disclosure.

[0039] Using this configuration, scattering events in the first scintillator crystal of the first PET detector block 214a provide a start signal for the re-coincident electron device, and the detection of scattered gamma photons 12 by the second scintillator crystal of the second PET detector block 214b provides a stop signal for the re-coincident electron device to detect re-coincidentity. Photons from the gamma-ray source undergo Compton scattering in the scintillator crystal of the first PET detector block 214a. Some of the photon's energy is transferred to an electron in the scintillator, and this electron immediately stops due to its interaction with the scintillator, resulting in a flash in one of the scintillator crystals of the first PET detector block 214a, which leads to the start signal for the re-coincident electron device. Forward-scattered gamma photons 12 with reduced energy travel across the FOV and are detected by the second detector block 214b, thus providing a stop signal for the re-coincident electron device. This coincidence detection identifies two scintillation crystals in the PET detector ring assembly 100, and because the two points in space define the straight path of the scattered photons across that space, it provides a transmission sine curve for counting the gamma photons 12 traveling through the FOV.

[0040] Therefore, generating transmission scan data using emission from a stationary gamma source 130 in the absence of any patient in the FOV produces a blank transmission scan sinusoid. By comparing the blank transmission sinusoid with a transmission scan from a transmission scan with a target volume in the FOV, a mu map (spatial map of photon attenuation coefficients) of the target volume can be constructed for attenuation correction of the emission PET scan data of the target volume.

[0041] Multiple radioactive isotopes can be used for gamma-ray source 130. Some examples are cesium-137, cobalt-60, and sodium-22. In some preferred embodiments, Cs-137 is used for gamma-ray source 130. Cs-137 is a good choice for gamma-ray source 130 because the gamma-ray energy is advantageous, slightly higher than the 511 keV commonly used in PET, but not so high that the gamma rays are outside the range commonly used in many PET detectors. Moreover, the 30-year half-life of Cs-137 is also advantageous, as the source will never need to be replaced during the expected lifespan of the PET scanner. Furthermore, only one gamma ray is produced per decay, while other isotopes produce multiple gamma rays with different energies. Finally, the Cs-137 isotope is widely available because it is one of the main components of nuclear reactor waste after it has been stored for several years.

[0042] According to this disclosure, a PET system configured to perform a method for simultaneously acquiring PET emission and transmission scan data by incorporating a stationary gamma-ray source and a time-of-flight (TOF) coincidence measurement includes... Figure 1B and 1C The PET detector ring assembly 100 shown in the figure includes a plurality of PET detectors arranged in a ring structure around a central opening, wherein the plurality of PET detector ring assemblies are arranged along a longitudinal axis L defined by the central opening (see figure). Figure 1A The PET detector assemblies are coaxially aligned. The patient tunnel T extends through the central opening of the coaxially aligned PET detector ring assembly 100. The plurality of PET detector assemblies are coaxially aligned along the length of the patient tunnel T.

[0043] Each detector electronics assembly 110 includes a PET detector block 214. Each detector block includes a detector 213 and one or more associated scintillator crystals 216. Multiple gamma-ray sources 130 are positioned behind the PET detector block 214, such as... Figure 1B As illustrated in the diagram. "Located behind the PET detector block 214" refers to any location on the side of the detector block 214 opposite to the patient tunnel T. For practical reasons, the gamma source 130 is preferably located within the PET stage 210 to shield and contain the gamma rays within the PET scanner.

[0044] Still referencing Figure 1CUsing this arrangement of the gamma-ray source 130 and the PET detector blocks 214, gamma photons 11 enter the first detector block 214 assembly from the rear. Some of the gamma photons undergo Compton scattering within the scintillator crystal 216 of the first PET detector block 214 assembly and are forward-scattered across the FOV of the PET scanner toward the second PET detector block 214 assembly on the opposite side of the FOV. The scattered gamma photons are then stopped when they encounter the scintillator crystal 216 associated with the second PET detector block assembly, and each scattered photon generates a flash that serves as a stop signal for the re-aligned electronics. The scattered photons are illustrated by arrow 12. This provides the basis for transmission imaging when the target subject is positioned within the patient tunnel T and within the FOV.

[0045] In some preferred embodiments of the disclosed system, the gamma source 130 may be provided with a collimating shield, allowing gamma photons directed toward the PET detector 214 to irradiate at full intensity, while gamma photons directed in other directions are absorbed by the shield, thereby reducing potential radiation exposure to personnel near the scanner. Specific configurations of the collimating shield can be designed for the specific radiation source selected for a given system to optimize the achieved transmission imaging.

[0046] Figure 2 and Figure 3 This is a schematic diagram of an example of a gamma-ray source assembly 130A and an associated PET detector ring 100. The gamma-ray source assembly 130A is positioned such that the gamma-ray source 130 is positioned in the gap between two PET detector blocks 214 within the PET detector ring assembly 100. The gamma-ray source assembly 130A includes a gamma-ray source 130 and a radiation shielding housing 136 for the gamma-ray source 130. The gamma-ray source assembly 130A is configured to hold the gamma-ray source 130 and switch between an ON configuration and an OFF configuration. In the ON configuration, except for gamma photons from the gamma-ray source 130 traveling in the direction toward the PET detector block 214, the remaining gamma photons from the gamma-ray source 130 are substantially absorbed by the housing 136. In fact, radiation absorption is never 100%. Therefore, "substantially absorbed" means that the gamma photons are strongly attenuated. The radiation shielding housing 136 may be made of a radiation-attenuating material such as tungsten.

[0047] In the illustrated example, a gamma-ray source 130 is attached to the end of a hinge rod 134, which allows the gamma-ray source 130 to be positioned within a housing 136 between an ON and OFF configuration. The radiation shielding housing 136 may include a channel 138, and in the OFF configuration, the gamma-ray source 130 can be retracted into the channel 138, as shown below. Figure 3As shown, this causes gamma photons emitted from the source to be absorbed by the housing 136. In the ON configuration, the hinge rod 134 extends the gamma source 130 beyond the channel 138. In the ON configuration, the gamma source 130 can be positioned in the opening 135, which opens toward the detector block 214.

[0048] Figure 4A This is a flowchart 300 illustrating a method for using one or more stationary gamma-ray sources 130 in a PET scanner as a transmission scanning radiation source to generate scan data (list pattern data) that can be used to generate attenuation maps. The method includes: (a) providing one or more stationary gamma-ray sources 130 in each PET detector ring assembly 100, wherein the stationary gamma-ray sources are positioned outside the detector ring assembly 100 (box 310); (b) identifying forward-scattered gamma photons 12, which are gamma photons emitted from the stationary gamma-ray sources 130 that have been forward-scattered by a scintillator crystal in a first set of detector blocks 214 in the PET detector ring assembly, through the FOV of the PET scanner, and detected via the scintillator crystal in a second set of detector blocks in the PET detector ring assembly through the PET scanner's reconciliation electronics (box 320); (c) from empty... (d) Acquire list pattern data from a blank transmission scan (i.e., when there is no activity in the FOV) (box 330); (e) Acquire list pattern data from a transmission scan with the target body in the FOV (box 340); (f) Simultaneously with step (d), acquire list pattern data from a PET scan when the target body is in the FOV (box 350); (f) Generate an attenuation map (mu map) by comparing the list pattern data from the blank transmission scan with the list pattern data from the transmission scan with the target body in the FOV (box 360); and (g) Apply the attenuation map to the list pattern data from the emission scan in step (e) to attenuate correct the list pattern data from that emission scan (box 370).

[0049] On the other hand, a PET scanner system 200 is disclosed, configured to perform the methods described in flowchart 300. The PET system includes: a plurality of detector ring assemblies 100, each including a plurality of scintillator crystals; a machine-readable storage medium 250; and a system controller connected to and communicating with the detector ring assemblies, wherein the machine-readable storage medium is encoded using computer program code such that, when executed by the system controller 290, the system controller performs a method including the following steps: (a) providing one or more stationary gamma-ray sources in each PET detector ring assembly, wherein the stationary gamma-ray sources are positioned outside the detector ring assembly; (b) identifying forward-scattered gamma photons emitted from the stationary gamma-ray sources, which have been forward-scattered by the scintillator crystals in a first set of detector blocks in the PET detector ring assembly, passing through… (c) The FOV of the PET scanner is detected by the scintillator crystal in the second detector block assembly in the PET detector ring assembly via the PET scanner's superimposed electronics; (d) List pattern data is acquired from a blank transmission scan (i.e., no activity in the FOV); (e) List pattern data is acquired from a transmission scan with the target body in the FOV; (f) Simultaneously with step (d), list pattern data is acquired from a PET scan when the target body is in the FOV; (g) An attenuation map (mu map) is generated by comparing the list pattern data from the blank transmission scan with the list pattern data from the transmission scan with the target body in the FOV; and (e) The attenuation map is applied to the list pattern data from the emission scan in step (e) to perform attenuation correction on the list pattern data from the emission scan.

[0050] In some embodiments, time-of-flight (TOF) considerations are applied to improve the quality of the transmission scan signal from forward-scattered gamma photons. Reference Figure 4BFlowchart 400, which considers the Time of Flight (TOF) when applied to a transmission imaging scan using a stationary gamma-ray source 130, includes: (a) calculating the time of flight of a scattered photon to reach the second scintillator crystal in the second detector block 214b based on the distance between the two scintillator crystals, the scattered photon being a photon from the gamma-ray source that has already undergone Compton scattering in the first scintillator crystal in the first detector block 214a (see box 410). This step calculates the TOF that should take for a gamma photon to travel across the FOV from the first scintillator crystal to the second scintillator crystal on the opposite side in a given PET scanner. In other words, the distance of interest here is from the first detector block 214a (specifically, the scintillator crystal associated with the first detector) where Compton scattering of gamma photons occurs to the second detector block 214b (specifically, the scintillator crystal in the second detector) where coincident scattered gamma photons are detected. Next, (b) defining a time window, wherein the time window has a width centered on the calculated TOF (see box 420). Next, (c) the TOF of actual scattered gamma photons originating from the first scintillator crystal is measured using the scanned target object in the FOV of the PET scanner (see box 430). Next, (d) the measured TOF from (c) is compared with the calculated TOF, and the measured TOF within that time window is identified (see box 440). Then, (e) the scattered gamma photons corresponding to those measured TOFs within that time window are identified as transmission source events originating from Compton scattering in the first scintillator crystal, thereby distinguishing the transmission type data from gamma emission annihilation events and random events within the scanned target object (see box 450). The resulting transmission scan data can be used to generate an attenuation map for correcting the master PET emission scan data.

[0051] In some embodiments, step (c) in flowchart 400 may further include simultaneously acquiring PET emission scan data of the target object in the FOV of the PET scanner.

[0052] On the other hand, a PET scanner system 200 is disclosed, configured to perform the method described in flowchart 400. The PET system includes: a plurality of detector ring assemblies 100 comprising a plurality of scintillator crystals; a machine-readable storage medium 250; and a system controller connected to and communicating with the detector ring assemblies, wherein the machine-readable storage medium is encoded with computer program code such that, when executed by the system controller 290, the system controller performs a method comprising the following steps: (a) calculating the time-of-flight (TOF) of a scattered photon reaching a second scintillator crystal in a second detector block based on the distance between two scintillator crystals, the scattered photon being a photon from a gamma-ray source that has already undergone Compton scattering in a first scintillator crystal in a first detector block. This step calculates the TOF that should take for a gamma photon to travel across the field of view (FOV) from the first scintillator crystal to the second scintillator crystal on the opposite side in a given PET scanner. In other words, the distance of interest here is the distance from the first detector block (specifically, the scintillator crystal associated with the first detector) where Compton scattering of gamma photons occurs to the second detector block (specifically, the scintillator crystal in the second detector) where coincident scattered gamma photons are detected. Next, (b) a time window is defined, wherein the time window has a width centered on the calculated TOF. Next, (c) the TOF of actual scattered gamma photons originating from the first scintillator crystal is measured using the scanned target object in the FOV of the PET scanner. Next, (d) the measured TOF from (c) is compared with the calculated TOF, and the measured TOF within the time window is identified. Then, (e) the scattered gamma photons corresponding to those measured TOFs within the time window are identified as transmission source events originating from Compton scattering (forward scattering) in the first scintillator crystal, thereby distinguishing the transmission type data from gamma emission annihilation events and random events within the scanned target object.

[0053] The use of stationary gamma-ray sources to obtain improved transmission scan data can be applied to PET / CT scanners as well as PET / MR scanners.

[0054] Using Cs-137 as a gamma-ray source, the inventors experimentally confirmed that, compared to using LSO background radiation, forward-scattered Cs-137 gamma rays provide improved transmission imaging in PET scanners by generating transmission imaging data more quickly to produce attenuation μ maps without CT hardware. The experimental work is presented below.

[0055] [Experimental Data]

[0056] The disclosed invention was demonstrated using a 6-ring Biograph Vision PET / CT scanner with LSO scintillation and SiPM-based detectors. Each ring of the scanner comprises 38 detector blocks measuring 64 mm × 32 mm, and each block contains 200 crystals measuring 3.2 x 3.2 x 20 mm. The annihilation radiation (511 keV) measured was sealed with 23 MBq. 68 The Ge source (germanium) is produced in a thin steel tube. The Cs-137 source, sealed in a lightweight housing, has a nominal strength of 115 MBq.

[0057] List pattern data (LM) is obtained over five scans and can be referenced. Figure 5A and 5B Let's explain. Scan 1 was acquired at 30 minutes, in which there was no activity near the scanner except for the LSO background, and no object or even patient bed was in the FOV. Scan 2 was acquired at 30 minutes, in which the germanium source was located near the center of the scanner's transaxial axis at position (A), but at the end of the axial range or just beyond the end of the axial range. Scan 3 was a 15-minute scan, in which the cesium source was located at position B, approximately 25 cm beyond the detector ring. Scan 5 was identical to Scan 4, except that the patient bed was placed in the FOV to provide a simple phantom. The detected radiation corresponds to various physical phenomena, including: normal PET overlap ( Figure 5A (Line DE in the middle); due to backscattering caused by 511 or 662 keV photons from A ( Figure 5A The DF line in the middle); forward scattering of photons from the Cs-137 source at point B ( Figure 5A The line GH in the middle); and the LSO scintillation crystal background ( Figure 5A The line CI in the middle). LSO is distributed around the PET detector ring 100, but in Figure 5A Only one exemplary location is indicated as (C).

[0058] The measurements were performed under special PET scanner settings. The energy window was widened to accept events with energies ranging from 160 to 730 keV. Furthermore, the coincidence time window was set to 6.64 ns, corresponding to a maximum chord length of approximately one meter. Both are limited in normal PET imaging.

[0059] The LM data includes the following information for each coincidence event: the energy signal of each photon, discretized in a bin with a width of 2.8 keV; the crystal identifier of each photon; the time-of-flight (TOF); and a bit indicating whether the coincidence was immediate or delayed. Data from each scan was categorized for two analyses. First, without limiting the TOF, a two-dimensional (2D) energy histogram was created using a 200x200 bin for immediate coincidence, and a matching histogram was created for delayed coincidence. In these histograms, the energy E of a photon in a detector is represented by... A Represented on the horizontal axis, and the energy E of another coincident photon B It is represented on the vertical axis. Secondly, without limiting the energy, a sine curve of net true coincidence was created, where events are used only when the TOF corresponds to the distance between the two crystals (±215 ps).

[0060] For transmission imaging, emitted photons from the stationary Cs-137 source 130 will be used efficiently, and they will be allowed to be scattered from almost all the crystals in the PET detector ring 100, so that virtually all response lines are sampled. The Cs-137 source should be close to the PET detector 214, but far enough to maintain an acceptable single-event rate. One way to do this is to distribute the stationary Cs-137 source 130 throughout the PET detector ring 100, such as... Figure 1C As shown in the diagram. Because the Biograph Vision PET scanner uses a nineteen-detector electronic assembly 110, Figure 1C The diagram shows nineteen angular positions. At each angle, two axial positions will be used, so the total number of sources should be 38.

[0061] The 38-source geometry was not used in the experiment described above, but the resulting sine curve can be estimated using the following calculations. Let S... bg-blank and S cs1-blank These are the sine curves derived from scan 1 and scan 4, respectively. These are 3D sine curves, but in the analysis, it will be sufficient to treat them as 2D functions of radial and angular sine curve coordinates, thus summing over all slices and all sine curve segments. Assume different source intensities, such as those that might be used in the new design, rather than the 115 MBq in the test described here. In this case, the 38-source sine curves, including the LSO background radiation, will be described by the following equation:

[0062] .

[0063] Because different scan times were used for the two input sine waves, S was calculated before the computation was performed. bg-blankDivide by two. In this equation, R is an operator that rotates the sine graph by one of nineteen angles, and the factor of 2 takes into account the source placement in both axial positions. This rotation must be performed carefully. The sine graphs are transformed from 180 degrees to 360 degrees, then (1) is applied, and then they are returned to the 180-degree representation.

[0064] It can be determined how fast transmission scanning can be performed using the cesium method, thus providing roughly the same image quality as in long scans using the LSO background method. This gain is approximated as:

[0065] .

[0066] Since (1) is linear, the gain can also be calculated simply without actually performing the rotation, as follows:

[0067] .

[0068] The following discussion section presents some points to note regarding the use of (1,2,3).

[0069] [result]

[0070] Figure 6 The table shows 2D energy histograms from scans 1, 2, 3, and 4. In these histograms, the energy range is 160 to 730 keV per photon, as mentioned above. Immediate coincidence is shown in the bottom row, and net true coincidence (immediate minus delay) is shown in the top row. The middle row shows the net true histogram where the LSO background is subtracted, with corrections for different scan times. The column representing scan 1 shows the expected energy spectrum, including the spectra of beta and gamma radiation in one detector, and a 202 or 307 keV peak in the other detector. The column representing scan 2 shows a normal PET coincidence with several features of interest. Near the upper right corner, small circular patterns are seen, representing the 511 keV peak in each detector. Below and to the left of this, notations for scattered radiation from one of the sources or detectors are seen, including another prominent feature at the lower end of the energy range, representing 511 keV backscattering through an angle close to 180 degrees. In part 2-B, the locus of the following points, also shown as diagonal boxes, is given by E. A + E B = 511 keV, E A and E BThe values ​​represent the energies of the coincidence event pairs, referred to as A and B. These represent the backscattering of 511 keV photons, as shown by the DF lines in Figure 5. Most of the backscattering overlaps with the LSO background. Columns of scan 3, where one photon is emitted at a time from the cesium source, show random coincidences in the bottom row but disappear in the top row, where delayed coincidences have been subtracted. In this case, for each pair of coincidence events A and B generated from 662 keV photons from the cesium source, the values ​​representing E are shown. A + E B = Another diagonal trajectory of 662 keV. These represent the backscattering of 662 keV photons from a cesium source. Again, this notation for backscattering strongly overlaps with the LSO background. The columns of scan 4 show the forward scattered radiation along lines such as GH in Figure 5. Again, the notation representing E is shown. A + E B = 662 keV point trajectory. It is worth noting that the notation of these photon pairs does not overlap with the LSO background radiation as strongly. In addition to the fact that the source does not need to be placed inside the FOV (which is the space best reserved for the patient), the more favorable energy range is a potential advantage of the forward scattering method.

[0071] As for scan 4, Figure 7 The distribution of single-event rates across the 228 detector blocks of the scanner is shown using grayscale representation. The horizontal direction corresponds to the circumference of the scanner, and the vertical direction corresponds to the axial dimension of the scanner. In the region near the cesium source, the maximum single-event rate is approximately 1 x 10⁻⁶. 5 The maximum single-particle rate is approximately 1 x 10^10^12 particles per block, compared to regions where single particles primarily originate from the LSO background radiation. 4 Count / block. The shape of the distribution confirms that the source location is approximately 25 cm beyond the detector. The actual amount of radiation impacting the detector is expected to be more than this, since the single-event rate only includes events where the deposited energy is above the arming threshold of approximately 150 keV, and roughly half of the 662 keV Compton scattering is expected to result in recoil electrons with energies below that threshold. Simple simulations show that approximately 40% of the incident radiation is detected in single-event mode. Other relevant effects include absorption and scattering by materials between the detector crystal and the source, such as electronics and the 3 to 8 mm thick aluminum channel to retain the coolant that stabilizes the detector during use.

[0072] Figure 8The net true sine curves from scans 1, 4, and 5 are shown (parts A, B, and C of the figure), as well as the ratio of the sine curve of scan 5 to that of scan 4 (part D). The sine curves have 520 radial bins and 399 angular bins. The LSO background sine curve from scan 1 is largely featureless, except for the expected pattern of high and low values ​​at the edges of the detector block, and a lower count in the middle, due to variations in stereo angle with position. The cesium spectrum from scan 4 is characterized by prominent diagonal orientation features due to forward-scattered radiation. As expected, the sine curve regions corresponding to small-angle scattering are reduced in brightness because less than 150 keV is applied to the crystal in which scattering occurs; that is, the electron recoil energy is too low to be detected. The transmission sine curve and the ratio sine curve from scan 5 show the gamma-ray shadow of the patient bed and illustrate the possibility of estimating the attenuation coefficient based on these measurements.

[0073] To illustrate and quantify the sensitivity gain due to cesium, the sine curves from scans 1 and 4 were reclassified from 520 radial chambers to 52x19 using 399 angular chambers. The ratio of the reclassified sine curves was calculated in each coarse chamber, again taking into account different scan durations. This ratio varied between 1.00 and 4.77. Figure 9 ). 38 source sine curves are estimated according to (1) and in Figure 10 The figure is shown in the diagram, which illustrates that its uniformity is similar to the sine curve from scan 1. The estimated 38-source sine curve has a count multiplier of 19.6 in the middle of the radial range and a count multiplier of 23.2 at the ends, relative to the sine curve from scan 1. Using (3), the sensitivity gain of the 38-source geometry is calculated as:

[0074] .

[0075] In these experiments, it is of interest to quantify the efficiency of detecting radiation from the LSO background or from the cesium source. Table I lists the source activity and sinusoidal count rates for scans 1 through 4. To estimate the corresponding efficiency, let NΩ / 4π be... 76 Lu or 137 The number of transmitted photons from each decay of a Cs nucleus, multiplied by the geometric efficiency of radiation, will provide the start signal for the recombinant electron device. In the case of LSO, the start signal will be generated by a cascade of beta and gamma radiation, and N=2, because as Figure 6 As shown, photons of 202 or 307 keV can be detected in coincidence. In this case, Ω is set to 4πsr, which is due to 176The Lu atom is surrounded by a scintillator. In the case of cesium, the initial signal will be due to the scattering of 662 keV gamma rays. In this case, N=1 is set because only scattered photons can be detected in the coincidence, and Ω is calculated as the sum of the stereo angles of all detector blocks facing the source from the position of the cesium source. The table lists NΩ / 4π and the quantity referred to herein as the coincidence efficiency. This is defined as:

[0076] .

[0077] Table 1. Normalized Sine Count Rates of Activities

[0078] LSO background (scan 1) Cesium (scan 4) Activity (Bq) 2.24E+06 1.15E+08 Sine curve rate (counts / second) 1.67E+04 8.87E+03 Sine curve rate / activity 7.44E-03 7.73E-05 NΩ / 4π 2 0.063 Overlap efficiency 3.72E-03 1.24E-3

[0079] [Discussion of Experimental Data]

[0080] The experimental data presented in this paper demonstrate the feasibility of acceleration in transmission imaging based on coincidence (such as coincidence from the LSO background). The study shows that forward-scattered cesium gamma rays can provide such acceleration. The use of the forward-scattered gamma photonic method disclosed herein offers advantages such as: the hardware, the stationary gamma-ray source can be directly integrated into the PET scanner; only a few simple moving parts may be required for the gamma-ray source housing; potentially no CT scanner is needed; and nothing obstructs the PET FOV or limits its diameter.

[0081] In a carefully designed implementation, the source will be provided with a collimating shield, allowing gamma rays guided toward the detector to irradiate at full intensity, while gamma rays guided in other directions will be absorbed by the shield, thereby reducing radiation exposure to personnel near the scanner.

[0082] There is reason to assume that the acceleration or sensitivity gain predicted by equation (4) underestimates the advantages of the forward scattering method. First, the scan 1 sine curve used in this analysis combines the results of... 176 LU emits 202 and 307 keV gamma rays; however, experience has shown that using only 307 keV gamma rays may be better. Secondly, most forward-scattered gamma rays from cesium have higher energies than lutetium gamma rays. This can be seen from... Figure 6 This is evident in part 4B. The higher energy produces advantages, including lower attenuation as the radiation penetrates the phantom and the patient's body. Moreover, less scattering and less confusion are expected due to the backscattering of the 511 keV radiation from the patient.

[0083] When using forward-scattered radiation, the physical effects at the lower end of the energy range are important. This is due to... Figure 8Part (B) is shown. The angular distribution of Compton scattering at 662 keV favors small angles, but scattering at the minimum angle leaves very little energy in the scintillator crystal, and this may be below the low-energy threshold of the scanner. Therefore, most of the scattered radiation cannot be detected. The use of multiple cesium sources (such as...) Figure 1C (As shown in the diagram) produces a sine graph without clearly defined edges. This is caused by... Figure 10 As shown, nineteen of the source locations have been combined with the calculation.

[0084] A figure of merit, known as the coincidence efficiency, is introduced to account for the relative efficiency of the two methods—LSO background and cesium forward scattering. Equation (5) and Table I indicate that the LSO background radiation can be utilized efficiently because 176 Lu-beta decays were detected in almost 100% of all decays, with two gamma rays emitted in each decay. In this experiment, the cesium source was positioned close to the PET detector ring, resulting in a 6.3% stereo angle with 4π sphericity. Taking all factors into account, this is comparable to 3.72 x 10⁻⁶ gamma rays against an LSO background. -3 In comparison, the forward scattering method is shown to be quite efficient, with a resolution of 1.24 x 10⁻⁶. -3 To explain why some potentially detectable coincidences are not actually observed, three effects should be noted. First, some of the 662 keV gamma rays may have already passed through the detector without any interaction, or been scattered or absorbed by the material between the source and the scintillator. In the case where the detector block is closest to the cesium source, these effects are estimated to account for approximately 40% of the loss. Second, some of the gamma rays may have entered the detector and been scattered at an angle of less than 40 degrees, leaving an amount of energy below the detection threshold. Based on a simple analysis of the Compton scattering formula, using a weighted average of the sine of the scattering angle, this effect is estimated to be another factor reducing the detection rate by approximately 40%. These two effects themselves account for most of the missing coincidences. A third type of effect also occurs, where, after scattering, most photons are not directed toward the detector on the other side of the FOV. Because this overall geometric efficiency factor would be similar in the case of LSO background emission, it is not easy to predict the ratio of the two coincidence efficiencies in Table I. In summary, the ratio is close to the amount expected based on simple calculations.

[0085] For forward scattering to be used practically in a PET scanner, there is a trade-off between the actual or perceived radiation hazard and the quality of the transmitted image. Therefore, the optimal amount of gamma-ray source material will be determined. In the example where Cs-137 is used, 38 sources can be considered, each with As = 30 MBq, totaling 1110 MBq, which exactly matches the total Cs-137 content used in the ECAT ART PET scanner. Equation (4) shows that the sensitivity gain in this case will be 6.5.

[0086] Experiments have confirmed that, using 662 keV gamma rays from Cs-137 as an example, forward scattering of gamma rays can be used for transmission imaging in PET scanners. In this application, the energy window and coincidence timing window should be wide. The resulting sine map can even be added to the sine map of the LSO background radiation. If such a PET scanner is manufactured with a total of 1110 MBq distributed among 38 cesium sources, transmission imaging over a typical 3-minute duration acquired by PET can have approximately the same quality as a 20-minute scan based solely on the LSO background.

[0087] The photon energy E detected from each pair of coincidence events A and B from forward-scattered gamma photons A and E B The total initial energy E of the gamma photons emitted from the gamma-ray source 130 was reached. I E A + E B = E I This can be used to enable PET systems to reject almost all background noise, such as background noise from false coincidence events. For a given gamma-ray source, the Egamma of gamma photons from that source is known. I For example, for Cs-137, E I It is 662 keV. Therefore, if Cs-137 is used as the gamma-ray source 130, then E is required. A + E B = E I = 662 keV can be used to screen for true coincidence events generated by forward-scattered Cs-137 gamma photons to improve the quality of the identified transmission scan signal.

[0088] For example, in Figure 4A In the method summarized in flowchart 300, step (b) may include using requirement E A + E B =E I To identify forward-scattered gamma photons, where E A and E BE represents the photon energy detected from each pair of coincidence events A and B, where event A is the Compton scattering in the scintillator crystal of the first detector block set, and event B is the Compton scattering in the scintillator crystal of the second detector block set. I It is the initial energy of the gamma photon emitted from the gamma-ray source.

[0089] In another example, Figure 4B In the method summarized in flowchart 400, step (e) may include using requirement E. A + E B = E I To identify forward-scattered gamma photons, where E A and E B These are the photon energies detected from each pair of coincidence events A and B, where event A is Compton scattering in the first scintillator crystal and event B is Compton scattering in the second scintillator crystal, and where E I It is the initial energy of the gamma photon emitted from the gamma-ray source.

[0090] In some embodiments where the scintillator crystal of the PET detector is an LSO crystal, the LSO crystal background radiation can also be used as transmission scanning radiation to generate an attenuation map according to a method previously disclosed in U.S. Patent Application No. 14 / 172,980, filed February 5, 2014, the contents of which are incorporated herein by reference.

[0091] According to the present disclosure presented above, a positron emission tomography (PET) scanner system is provided, comprising: a gantry; a plurality of PET detector ring assemblies provided within the gantry, wherein each detector ring assembly includes a plurality of PET detectors arranged in a ring structure around a central opening, wherein the plurality of PET detector ring assemblies are coaxially aligned along a longitudinal axis defined by the central opening; a patient tunnel extending through the central opening of the coaxially aligned PET detector ring assemblies, wherein the plurality of PET detector assemblies are coaxially aligned along the length of the patient tunnel, wherein each of the plurality of PET detectors includes a detector and one or more scintillator crystals associated with the detector; and one or more stationary gamma-ray sources provided in each PET detector ring assembly within the gantry.

[0092] In a PET scanner system according to any of the above embodiments, one or more stationary gamma-ray sources in each PET detector ring assembly may be positioned behind the detector ring assembly, away from the patient tunnel. In a PET scanner system according to any of the above embodiments, the stationary gamma-ray source may be Cs-137, cobalt-60, or sodium-22. In some embodiments of the PET scanner system, the stationary gamma-ray source is Cs-137.

[0093] In any of the above embodiments of the PET scanner system, each gamma ray source may be provided in an assembly that provides radiation shielding.

[0094] In any of the above embodiments of the PET scanner system, the assembly providing the radiation shield may include an ON configuration and an OFF configuration, wherein when in the ON configuration, all gamma photons from Cs-137 except those traveling in the direction toward the detector assembly are substantially absorbed by the assembly providing the radiation shield. In any of the above embodiments of the PET scanner system, the assembly providing the radiation shield may be made of a material that at least partially comprises tungsten.

[0095] In any of the above embodiments of the PET scanner system, the PET scanner may be a PET / MR scanner. In any of the above embodiments of the PET scanner system, the PET scanner may be a PET / CT scanner.

[0096] Although the subject matter has been described with reference to exemplary embodiments, it is not limited thereto. Rather, the appended claims should be interpreted broadly to include other variations and embodiments that may be made by those skilled in the art.

Claims

1. A positron emission tomography (PET) scanner system, comprising: stand; Multiple PET detector ring assemblies are provided in a bench, wherein each detector ring assembly includes multiple PET detectors arranged in a ring structure around a central opening, wherein the multiple PET detector ring assemblies are coaxially aligned along a longitudinal axis defined by the central opening; A patient tunnel extending through a central opening of coaxially aligned PET detector ring assemblies, wherein the plurality of PET detector assemblies are coaxially aligned along the length of the patient tunnel, and each of the plurality of PET detectors includes a detector and one or more scintillator crystals associated with the detector; as well as One or more stationary gamma-ray sources are provided in each PET detector ring assembly within a gantry, wherein the one or more stationary gamma-ray sources are configured to emit forward-scattered gamma photons, which are forward-scattered by at least a portion of the one or more scintillator crystals provided in the corresponding PET detector ring assembly.

2. The PET scanner system of claim 1, wherein the one or more stationary gamma sources in each PET detector ring assembly are positioned behind the detector ring assembly, away from the patient tunnel.

3. The PET scanner system of claim 2, wherein the stationary gamma ray source is Cs-137, cobalt-60, or sodium-22.

4. The PET scanner system of claim 2, wherein the stationary gamma ray source is Cs-137.

5. The PET scanner system of claim 2, wherein each of the gamma-ray sources is provided in an assembly that provides radiation shielding.

6. The PET scanner system of claim 3, wherein each of the gamma-ray sources is provided in an assembly that provides radiation shielding.

7. The PET scanner system of claim 5, wherein the assembly providing the radiation shield includes an ON configuration and an OFF configuration, wherein when in the ON configuration, the remaining gamma photons from Cs-137, except for gamma photons from Cs-137 traveling in the direction toward the detector assembly, are substantially absorbed by the assembly providing the radiation shield.

8. The PET scanner system of claim 6, wherein the assembly providing radiation shielding includes an ON configuration and an OFF configuration, wherein when in the ON configuration, the remaining gamma photons from Cs-137, except for gamma photons from Cs-137 traveling in the direction toward the detector assembly, are substantially absorbed by the assembly providing radiation shielding.

9. The PET scanner system of claim 7, wherein the assembly providing the radiation shield is at least partially composed of tungsten.

10. The PET scanner system of claim 8, wherein the assembly providing the radiation shield is at least partially composed of tungsten.

11. The PET scanner system of claim 1, wherein the PET scanner is a PET / MR scanner.

12. The PET scanner system of claim 1, wherein the PET scanner is a PET / CT scanner.

13. A method for using one or more stationary gamma-ray sources provided in a PET scanner as transmission scanning radiation sources to generate scan data that can be used to generate attenuation maps, wherein the PET scanner includes a plurality of PET detector ring assemblies, the method comprising: (a) One or more stationary gamma-ray sources are provided in each PET detector ring assembly, wherein the stationary gamma-ray sources are positioned outside the detector ring assembly; (b) Identifying forward-scattered gamma photons, which are gamma photons emitted from a stationary gamma-ray source, which have been forward-scattered by a scintillator crystal in the first detector block assembly in the PET detector ring assembly, pass through the field of view (FOV) of the PET scanner, and are detected by the scintillator crystal in the second detector block assembly in the PET detector ring assembly via the PET scanner's reconciliation electronics; (c) Acquire list pattern data from blank transmission scans in which there is no activity in the FOV; (d) Obtain list pattern data from a transmission scan of a target body within the field of view; (e) An attenuation map is generated by comparing list pattern data from a blank transmission scan with list pattern data from a transmission scan containing a target body within the FOV; and (f) Apply the attenuation map to the list mode data from the transmit scan in step (e) to perform attenuation correction on the list mode data from the transmit scan.

14. The method of claim 13, further comprising: Simultaneously with step (d), list pattern data is acquired from a PET scan while the target body is in the field of view (FOV).

15. The method of claim 13, wherein step (b) comprises: Usage Requirements E A + E B = E I To identify forward-scattered gamma photons, where E A and E B E represents the photon energy detected from each pair of coincidence events A and B, where event A is the Compton scattering in the scintillator crystal of the first detector block set, and event B is the Compton scattering in the scintillator crystal of the second detector block set. I It is the initial energy of the gamma photon emitted from the gamma-ray source.

16. A method for improving the quality of a transmission scan signal from forward-scattered gamma photons by taking time-of-flight (TOF) into account when performing a transmission imaging scan using a stationary gamma-ray source, comprising: (a) The TOF of a scattered photon reaching the second scintillator crystal in the second detector block is calculated based on the distance between the two scintillator crystals, the scattered photon being a photon from a gamma-ray source that has undergone Compton scattering in the first scintillator crystal in the first detector block; (b) Define a time window, wherein the time window has a width centered on the calculated TOF; (c) Measure the TOF of actual scattered gamma photons originating from the first scintillator crystal using the scanned target object in the field of view (FOV) of the PET scanner; (d) Compare the measured TOF from (c) with the calculated TOF and identify the measured TOF within the time window; as well as (e) The forward-scattered gamma photons corresponding to those measured in the time window are identified as transmission source events originating from Compton scattering in the first scintillator crystal, thereby distinguishing the transmission type data from gamma emission annihilation events and random events within the scanned target object.

17. The method of claim 16, wherein step (c) further comprises simultaneously acquiring transmission scan data of the target object in the FOV of the PET scanner.

18. The method of claim 17, further comprising: Simultaneously with step (c), acquire PET emission scan data of the target object in the FOV of the PET scanner; as well as Attenuation maps are generated from transmission scan data to correct PET emission scan data.

19. The method of claim 16, wherein the step (e) comprises: Usage Requirements E A + E B = E I To identify forward-scattered gamma photons, where E A and E B These are the photon energies detected from each pair of coincidence events A and B, where event A is Compton scattering in the first scintillator crystal and event B is Compton scattering in the second scintillator crystal, and where E I It is the initial energy of the gamma photon emitted from the gamma-ray source.

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