The present invention relates to systems and methods for emission tomography and, more particularly, to systems and methods for inter-detector, scatter-enhanced positron emission tomography that provides an increase in the performance of current positron emission tomography scanners by allowing the counting of coincidences caused by inter-detector, scattered photons.
There are a variety of emission tomography imaging systems and methods. One clinically important example is positron emission tomography (PET), which, generally, utilizes an administered radionuclide to acquire two-dimensional and three-dimensional tomographic images of a target area or organ of interest in a subject. More specifically, such radionuclides are employed as radioactive tracers called “radiopharmaceuticals” by incorporating them into substances, such as glucose or carbon dioxide. These radiopharmaceuticals are then administered to the patient where they become involved in biological processes such as blood flow; fatty acid and glucose metabolism; and protein synthesis. Through a respective biological process, the radiopharmaceuticals accumulate in, or otherwise target, the area or organ of interest in the subject. By measuring or identifying photons emitted from the area or organ of interest by the accumulated or targeted radiopharmaceutical, clinically useful biological and physiological information can be acquired from the area or organ of interest.
For example, in PET, as the injected radioactive tracer decays, it emits positrons. The positrons travel a very short distance before they encounter an electron and, when this occurs, the positrons are annihilated and converted into two photons, or gamma rays. This annihilation event is characterized by two features that are pertinent to PET imaging. Namely, each gamma ray has an energy of 511 keV and the two gamma rays are directed in substantially opposite directions. An image is created by determining the number of such annihilation events at each location within the scanner's field of view.
To create such an image, typical PET scanners consist of one or more rings of detectors that are positioned to encircle the patient. Coincidence detection circuits connected to the detectors record only those photons that are detected simultaneously by two detectors located on opposite sides of the patient and that fall within an energy acceptance window around 511 keV. The number of such simultaneous events indicates the number of positron annihilations that occurred along a line joining the two opposing detectors. Within a few minutes, hundreds of millions of events can be recorded to indicate the number of annihilations along lines joining pairs of detectors in the ring. These numbers are employed to reconstruct an image using well-known tomographic reconstruction techniques.
For example, current clinical (and most preclinical) PET scanners and systems include a ring of block detectors for detecting emitted photons, typically in circular, such as the array shown in FIG. 1, or in hexagonal or octagonal arrays. Block detectors include a piece of scintillator material that converts the energy deposited by gamma rays into visible light. The scintillator material is usually segmented into many scintillation crystal elements configured in an array, which is read out by a number of individual photo-multiplier tubes (PMTs) or a position-sensitive photo-multiplier tube (PS-PMT) that convert the light emitted by the scintillation material into electrical signals having a magnitude proportional to the energy deposited by the gamma rays in the scintillator material. By combining the output signal of the PMTs or PS-PMT of the block detector, it is possible to determine the single crystal in which the detected photon interacted and the energy deposited by such photon.
Although block detectors have been demonstrated as the most cost-effective solution for the implementation of PET scanners, these detectors also present some drawbacks. For example, since each detector element is a block, if several photons interact simultaneously on the same block and the added energy of those photons is within a predefined energy acceptance window (around 511 keV), it is not possible to determine from the output signals of the detector if they were produced by the interaction of a single photon (thereby presenting useful information) or by the interaction of multiple photons (thereby presenting distorted or non-useful information).
In addition, as shown in FIG. 1, the ring of block detectors of a PET scanner includes individual detectors that are operated in coincidence with a fan beam of block detectors on the opposite side of the ring. The inner circle formed by edges of all such fan beams defines the useful field of view. Data is usually recorded simultaneously for all possible fan beams, and the PET scanner will produce an output whenever two photons are detected in opposite block detectors of a fan beam within a specified coincidence timing window (for example, in the range of hundreds of picoseconds to tens of nanoseconds) and when both events fall into a predetermined energy window (511 keV±ΔE, where ΔE is a function of the energy resolution of the block detectors). Any such events are called prompt coincidences, but can be of three specific types: true coincidences, scatter coincidences, and random coincidences.
True coincidences occur when two photons produced from the same annihilation are detected within the time and energy windows of the system, as shown in FIG. 2A. Scatter coincidences occur when at least one of the photons undergoes scattering in the object under study, such as Compton scattering, where the photon loses a fraction of its total energy in the scatter interaction with the object before its detection. The scatter coincidence is, thus, detected in a pair of detectors that are non-collinear with the originating annihilation, as shown in FIG. 2B. Random coincidences, also known as accidental coincidences, occur when annihilation photons from two unrelated positron annihilation events are detected in opposite detectors, as shown in FIG. 2C. True coincidences produce valid information, while both scatter coincidences and random coincidences produce distorted information. In particular, scatter and random coincidences yield incorrect positional information, as shown by the dotted lines in FIGS. 2B and 2C, and contribute to a relatively uniform background noise in the resulting image, which results in a loss of contrast.
With respect to scatter coincidences, such events are typically assumed to occur only due to scattering within the patient, as shown in FIG. 2B, and current PET systems include scatter correction procedures based on this assumption. However, there are a large number of events in which Compton scattering occurs in the block detectors of the scanner, as shown in FIGS. 3A and 3B, depositing a fraction of the total energy of the photon in each interaction. In particular, FIG. 3A illustrates a scatter event where one of the photons from an annihilation event (photon A) interacts by photoelectric effect depositing energy in a detector within the acceptance energy window of the scanner (that is, 511 keV±ΔE), and the other photon (photon B) interacts by Compton scattering in another detector. Photon B deposits some of its energy in the detector it is incident upon, and the scattered photon (photon C) produced by the Compton scattering event deposits energy in another detector. FIG. 3B illustrates a scatter event where one of the photons from an annihilation event (photon A) interacts by photoelectric effect depositing energy in a detector within the acceptance energy window of the scanner, and the other photon (photon B) interacts by Compton scatter in another detector, where it deposits some of its energy, with the scattered photon (photon C) escaping from the detector ring.
In current clinical and preclinical PET scanners that include block detectors, no viable information is used from the scatter events shown in FIG. 3A because multiple detections are not identified by the coincidence system. That is, such events are rejected. Scatter events shown in FIG. 3B (that is, crystal scatter coincidences with two detection events) may be detected and processed in the same fashion as scatter events that have undergone scattering in the object (as shown in FIG. 2B). Thus, the data collected for events comprising more than two detections is thrown out and only data from prompt coincidences (including true coincidences, in-body scatter coincidences, random coincidences, and crystal scatter coincidences with two detection events) are used to compose images. This limits the potential sensitivity of the system and quality of the resulting images.
Approaches have been presented to make use of inter-detector scatter events (in particular, events as shown in FIG. 3A); however, such approaches have only been proposed using non-standard detector configurations, such as Compton cameras or high granularity detectors, and cannot be used with conventional block-detector type PET scanners. Such non-standard detectors can be very expensive and the corresponding systems must be combined with complicated mathematical models. Furthermore, some of these systems require the use of inter-detector scatter data to perform within the same range as those obtained in block detector-based scanners that do not use such data. Thus, although these non-standard detectors may be capable of detecting inter-detector scatter events, they do not produce higher quality images than traditional block-detector type PET.
For example, a Compton camera is a radiation detector that is usually composed of two detection planes, commonly made from semiconductor materials, which provide better energy resolution than radiation detectors using typical scintillation crystals. Photons emitted from a source are scattered in the first plane through Compton scattering and are absorbed in the second plane through photoelectric effect. In both planes, the position of the interaction and the energy deposited are measured. The detectors are operated in coincidence, so that only photons that interact with both detector-planes and deposit a total energy within a given window are recorded. In this case, and due to the disposition of the detectors, it is improbable that the first interaction would be detected in the second detector plane. Both the energy of the initial photon and the energy deposited in the first detector are known, and therefore the scattering angle can be calculated using the Compton formula. This defines a conic surface in which the origin of the initial photon is contained. The precision of this calculation is strongly related to the energy resolution of the detector being used for the construction of the Compton camera. The real location of the source is obtained as the intersection point of several of these conic surfaces. When several point sources or a continuous source distribution are imaged (for example, an organ containing a radioisotope), the reconstruction becomes more challenging, resulting in this type of device usually providing images of poor quality in comparison to traditional PET images.
In another example, scanners with high granularity detectors (detectors that, unlike block detectors, require that each detection element be read-out independently) are expensive and computationally complex. More specifically, an objective of such high-granularity detector systems is to determine the sequence of interaction points of each photon (for example, the first interaction point, the second interaction point, etc.) to find the appropriate line of response in a multiple coincidence event. For example, given the example of FIG. 3A, such systems must determine whether the first photon in the inter-detector scatter event was photon B (thereby illustrating an annihilation response line along A-B), or photon C (thereby illustrating an annihilation response line along A-C). These determined lines of response from multiple coincidence events are then combined with true coincidence lines to compose images. One system in particular requires the use of three-dimensional high-granularity semiconductor detectors. Unlike scintillation crystals used in block detectors, semiconducting detectors, such as cadmium zinc telluride (CZT), directly sense the ionization signal created by the annihilation photon absorption. CZT detectors have good energy resolution, but their stopping power for 511 keV is lower than most scintillation crystals, and the timing resolution is much worse than what can be achieved with block detectors. Furthermore, the composition of the semi-conductor detector and detection system design cause a large fraction of all photons (around 94%) undergo inter-detector scattering. For this reason, the ability to correctly position inter-detector scatter events strongly determines the performance of such a system. This results in a high price and high complexity of such scanners with little improvement in performance in comparison to block detector-type scanners, which can rely heavily on true coincidence events. As a result, to date, high-granularity camera scanners cannot compete with the performance and cost-effectiveness of block detector-type scanners.
Thus, current approaches for utilizing inter-detector scatter events require non-conventional scanners that are more expensive, are more computationally complex, and cannot, generally, achieve higher-quality images than current conventional block-detector scanners. Furthermore, the methods developed for interpreting and recording inter-detector scatter events using non-conventional scanner approaches cannot practically be applied to conventional block-detector scanners due to their inherent precision and performance characteristics.
Therefore, it would be desirable to have a system and method for block detector-based PET imaging that has increased sensitivity, such as by using data collected from inter-detector scatter events.