Classifications

• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
  • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
  • G01T1/16—Measuring radiation intensity
  • G01T1/1603—Measuring radiation intensity with a combination of at least two different types of detectors
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
  • A61B6/02—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
  • A61B6/03—Computed tomography [CT]
  • A61B6/037—Emission tomography
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
  • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
  • G01T1/29—Measurement performed on radiation beams, e.g. position or section of the beam; Measurement of spatial distribution of radiation
  • G01T1/2914—Measurement of spatial distribution of radiation
  • G01T1/2985—In depth localisation, e.g. using positron emitters; Tomographic imaging (longitudinal and transverse section imaging; apparatus for radiation diagnosis sequentially in different planes, steroscopic radiation diagnosis)

Definitions

• TITLE PET AND SPECT MULTI PHOTON IMAGING WITH SINGLE RADIOACTIVE
• Embodiments disclosed herein relate generally to systems and methods for nuclear imaging, including Positron Emission Tomography and Single- Photon Emission Computed Tomography,
• Positron Emission Tomography employs a radioisotope that emits a positron (antiparticle of electron) when it undergoes radioactive decay.
• the positron subsequently annihilates with an electron in the surroundings, typically resulting in two relatively high energy photons (gamma rays) travel!ing along a straight line in opposite directions (approximately 180 degrees apart).
• gamma detectors employs a radioisotope that emits a positron (antiparticle of electron) when it undergoes radioactive decay.
• the positron subsequently annihilates with an electron in the surroundings, typically resulting in two relatively high energy photons (gamma rays) travel!ing along a straight line in opposite directions (approximately 180 degrees apart).
• gamma detectors two relatively high energy photons
• Single Photon Emission Computed Tomography is also based on the detection of photons.
• the radioisotopes typically used in SPECT emit a photon directly during radioactive decay.
• the gamma cameras are typically unable to determine depth information. Instead, a number of two- dimensional images are captured from a number of different angles, and a computer is used to generate a tomographic reconstruction of these multiple 2-D images ⁇ or projections).
• FIG. 1 is a plot of fractional energy changes for 51 1 keV PET photons and 142 keV 99m Tc SPECT photons as a function of angle;
• FIG. 2 is a schematic representation of the detection of a photon emitted as a result of radioactive decay of a radionuclide
• FIG. 3 is a schematic representation of the detection of photons emitted as a result of positron annihilation
• FIG. 4 is a schematic representation of the detection of a photon emitted as a result of radioactive decay of a positron emitting radionuclide and photons emitted as a result of positron annihilation according to one embodiment
• FIG. 5 is a schematic representation of the detection of photons emitted as a result of radioactive decay of a radionuclide according to one embodiment
• FIG. 6 is a schematic block diagram of a detector system according to one embodiment
• FIG. 7 shows the neutron cross section (a measure of production rate) for 123 Xe plotted against proton energy
• FIG. 8 shows the neutron cross section (a measure of production rate) for 73 Se plotted against proton energy.
• Positron Emission Tomography PET
• SPECT Single Photon Emission Computed Tomography
• PET Positron Emission Tomography
• SPECT Single Photon Emission Computed Tomography
• Positron Emission Tomography employs a radioisotope which emits a positron (the antiparticle of electron). The positron subsequently annihilates with an electron in the surroundings, resulting in two 511 keV photons (gamma rays) travelling along an approximately straight line in opposite directions (i.e. 180 degrees apart).
• the PET technique registers the two photons entering the detectors to reconstruct the line of flight of the photons.
• the origin of the photons (e.g. the location of the positron annihilation) can be estimated to about a centimeter precision by employing the time of flight information of the detected photons. For example, if two 51 1 keV photons are detected at two detectors within a predetermined time interval, they are assumed to have originated from the same annihilation event; this annihiiation event is assumed to have taken place along a reconstructed line of flight of the photons (e.g. a fine between the two detectors). Also, the annihilation event can be estimated to have taken place closer to the detector that detected the first of the two detected photons, this estimated distance being proportional to the elapsed time between the detection of the two photons.
• PET assumes that the point of photon emission is same as the location where the radioisotope decays - e.g. the location of the emission of the positron.
• positrons emitted during radioactive decay are moving at high speeds, almost 30-70% of speed of light. This would mean that in many cases, the positron annihilation point is not the same as the location where the positron was "born" (i.e. emitted).
• a location of a radioisotope determined by a PET scan may not represent the precise location of the radioisotope from which it is emitted.
• PET utilizes pairs of photons emitted from positron-electron annihilation events
• Single Photon Emission Computed Tomography makes use of one photon emitted directly from gamma-emitting radionuclides. While a SPECT photon originates at the decay vertex of the isotope - i.e. there is no material difference in the location of the photon emission and the location of the radioisotope at the time of this emission - the one photon imaging does not provide depth information.
• CT Computed Tomography
• CAT Computed Axial Tomography
• multiple two- dimensional images of photon emission - each taken at a different angle with respect to the target sample being scanned - are post-processed in order to create a three-dimensional volumetric image (also referred to as a tomographic reconstruction, or simply a tomogram) of the target sample.
• Common radionuclides used for nuclear imaging include 1 C, 13 N and 13 F for PET and 99m Tc, 67 Ga, 111 ln and 123 l for SPECT.
• radionuclide also referred to as a radioactive tracer
• FDG fluorodeoxyglucose
• PET and SPECT have their own limitations that can result in ambiguities and misdiagnoses. These limitations include photon absorption and scattering.
• Attenuation correction must be made in both PET and SPECT and the process for correction is different for each.
• attenuation in PET is different from attenuation in SPECT because PET requires both photons from an annihilation event to reach the detectors for a coincident event to be registered.
• the probability that the two PET photons emitted from a single annihilation event will be attenuated i.e. a product of the attenuation probability for each photon, each attenuation probability being less than one
• the probability that one single photon emitted during radioactive decay e.g. in SPECT
• the attenuation is a sensitive function of photon energies.
• SPECT photons are usually of much lower energy ( ⁇ 200 keV) while the PET photons are of 51 1 keV.
• the lower energy SPECT photons are more likely to be absorbed rather than scattered, resulting in a loss of intensity or information rather than a misidentifi cation.
• the emitted photons can be scattered (i.e. deflected) on their way to the gamma detectors. This can result in a misidentification of their point of origin. While collimators are typically used to help reduce image artifacts, their use may lower the contrast resolution of the detector due to a decreased signal-to-noise (SNR) ratio. Also, while large angle scattering of photons may be discarded as the photon energy will be quite different from the expected PET/SPECT photon energies, the finite resolution of detectors may admit scatters as large as 30 degrees for PET detectors, and even much larger angles for SPECT detectors.
• SNR signal-to-noise
• Figure 1 shows fractional energy changes for 51 1 keV PET photons and 142 keV 99 m Tc SPECT photons as a function of angle.
• photons can be absorbed (e.g. lost) between emission and the detectors. This loss of information can set a limit on the achievable contrast. Also, it is often unclear if the photon is absorbed in the human body or somewhere else on its way before its gets into the sensitive volume of the detector.
• the result could effectively be a single isotope imaging of either the PET or SPECT variety, unless one administers a relatively high dose of the radioisotope (PET or SPECT isotope) with the shorter half-life to achieve an effective combined imaging over the life of the radioisotope with the longer half-life, and even then it is likely that the relative intensities of the two radioisotopes will result in imaging based on one of the two energies being dominant (e.g. either the PET or SPECT variety).
• PET or SPECT isotope the radioisotope
• SPECT/CT an additional disadvantage is that the SPECT photon (typically having an energy of about 140 to about 200 keV) and CT X-ray photons (typically having an energy of about 100 to about 120 keV) may not constitute distinct enough energies to provide significant attenuation information.
• Patients may be administered higher radiation doses (due to the exposure to two or more separate radiation sources).
• Embodiments of the present invention provide an improved method for determining an approximate location of a radionuclide at a time of its radioactive decay.
• the method involves the use of a single radioisotope consisting essentially of radionuclides that, at the time of their radioactive decay, emit both a positron and a primary photon.
• a radioisotope that emits positrons accompanied by photons emitted from a daughter product of the decay path - the positron and a primary photon being emitted within a relatively small time interval (e.g. a separation of less than a picosecond) - a result is a method of dual energy imaging that may be highly localized in both space and time. That is, during this time interval, the positron- emitting source typically moves less than a micron, while the positron itself may migrate up to a few millimeters from the location of the emission of the primary photon before annihilation.
• the positron annihilation photons and the primary photon are both temporally and spatially correlated.
• the emission point may be localized with greater spatial resolution, this may lead to 3-D imaging that is more accurate - and in some cases may be much more accurate - than, for example, single photon X-ray imaging, single photon SPECT imaging, or two photon PET imaging.
• embodiments described herein may reduce the amount of radiation administered to a patient during medical imaging. For example, data that would ordinarily require both a PET and separate SPECT scan to be conducted (which would normally require two separate radioisotopes to be administered) may be acquired using only a single radioisotope. Also, data acquisition time may be reduced (as only one scan - not two -is required), which may make the process less strenuous to the patient. The technical staff time may also be decreased.
• the use of two or more separate radioisotopes in addition to requiring injections of two isotopes, may entail inefficient and excessive dosages of isotopes to be administered, as the two isotopes will likely not have identical half-lives.
• the two sets of measurements on patients would typically result in separate data sets (one for PET and another for SPECT) which are uncorrelated with each other. That is, the intensities and half- lives of the two sources are uncorrelated, and there is no temporal or locational correlation between separate emission events. By detecting temporally and spatially correlated emission data, the accuracy or precision or both of the determined locations or decay events may be improved.
• the isotope production method as well as other chemical processing required to prepare an isotope sample to be administered to a patient may be simplified.
• Use of a single radioisotope will typically involve only one production method and one chemical processing system, rather than two separate production channels and chemical treatments that would typically be needed to prepare and administer two separate isotopes.
• our methods may be implemented using the latest generation of PET machines without significant changes to the hardware.
• Current generation PET machines typically operate using "list-mode" data collection, in which a time-indexed list of photon detections is generated. For a number of radionuclide decay events, it is possible that not all photons of interest from that decay event will be detected.
• not all emitted photons may reach a photon detector, or some photons may hit' a detector but go unrecorded due to the detectors having less than 100% detection efficiency, or both.
• list-mode may provide flexibility to classify and/or group detection events separately as detection events where three photons are detected (the primary photon and a pair of positron annihilation photons) and as detection events where only two or one of the three photons emitted during a radioactive decay of a radionuclide are detected.
• data validity can be cross checked and several artifacts can be examined.
• Using a technique referred to as "software cuts” multiple images - each image being generated based on specific logic conditions - may be created to critically evaluate physical and geometrical artifacts to facilitate false-negative and false-positive diagnostics.
• a method of determining an approximate location of a radionuclide at a time of radioactive decay of the radionuclide comprising: providing a plurality of photon detectors directed towards a target region, introducing into the target region the radionuclide that, at the time of its radioactive decay, emits a positron and a primary photon, detecting the primary photon at a first one of the plurality of photon detectors and detecting at least one secondary photon emitted from an annihilation of the positron at at least a second, one of the plurality of photon detectors, and determining the approximate location of the radionuclide at the time of its radioactive decay based on a location of the first one of the plurality of photon detectors, a location of the second one of the plurality of photon detectors, and a presumed common point of origin of the detected primary photon and the detected at least one secondary photon.
• the primary photon is emitted from a daughter product of the radionuclide following the emission of the positron.
• the plurality of photon detectors is capable of characterizing a detected photon as either primary or secondary.
• the primary photon has a primary expected energy
• the at least one secondary photon has a secondary expected energy
• the characterizing is based on a comparison of a detected energy of the detected photon and at least one of the primary and secondary expected energies.
• the determining is further based on a time of detection of the primary photon and a time of detection of each of the at least one secondary photon.
• the detecting at least one secondary photon comprises detecting two secondary photons.
• the approximate location of the radionuclide at the time of its radioactive decay is determined based on a selected two of the three detected photons.
• the approximate location of the radionuclide at the time of its radioactive decay is determined based on a selected one of the three detected photons.
• the detecting of the primary photon and the at least one secondary photon is performed during a predetermined time interval.
• the plurality of photon detectors comprises a ring of photon detectors disposed about the target region.
• the introducing comprises introducing a molecule or compound tagged with a plurality of radionuclides into the target region.
• the method further comprises repeating the detecting over a plurality of predetermined time intervals, and for substantially each predetermined time interval in the plurality of predetermined time intervals during which one of the plurality of radionuclides undergoes radioactive decay, determining the approximate location of that radionuclide at the time of its radioactive decay based on a presumed common point of origin of photons detected during that predetermined time interval.
• the method further comprises determining a volumetric reconstruction of a concentration of the plurality of radionuclides within the object based on a statistical reconstruction of the determined approximate locations of substantially each of the plurality of radionuclides.
• the radionuclide is 123Xe, 73Se, 75Br, 174Ta, 99Rh, or 184lr. In some embodiments, the method of claim 14, wherein the radionuclide is 123Xe. In some embodiments, the radionuclide is 73Se.
• a system for determining an approximate location of a radionuclide that, at the time of its radioactive decay, emits a positron and a primary photon comprising: a plurality of photon detectors directed towards a target region, the plurality of photon detectors capable of detecting the primary photon at a first one of the plurality of photon detectors and detecting at least one secondary photon emitted from an interaction of the positron and an electron at least a second one of the plurality of photon detectors, and a computing device capable of determining the approximate location of the radionuclide at the time of its radioactive decay based on a location of the first one of the plurality of photon detectors, a location of the second one of the plurality of photon detectors, and a presumed common point of origin of the detected primary photon and the defected at least one secondary photon.
• a method of localizing a radioisotope within a target region at a time of radioactive decay of the radioisotope comprising: providing at least one photon detector directed towards the target region, introducing the radioisotope into the target region, the radioisotope comprising a radioisotope that, at the time of radioactive decay, emits a positron and at least one primary photon, detecting the at least one primary photon and at least one secondary photon emitted from an interaction of the positron and an electron using the at least one photon detector, and localizing the radioisotope based on the detected at least one primary photon and the detected at least one secondary photon.
• Embodiments of the present application provide methods for determining an approximate location of a radionuclide that, at a time of its radioactive decay, emits both a primary photon and a positron.
• specific embodiments are described below as an example of the general method; it is to be understood that alternate embodiments are feasible.
• different embodiments may, in variant implementations, be associated with radioisotopes not specifically mentioned, or using other types of photon detection equipment.
• FIG. 2 is an illustrative schematic example of single photon imaging in which a number of photon detectors 200 are directed towards a target region 205 that contains a radionuclide 210.
• a primary photon 220 having an energy ⁇ (referred to herein as a primary energy) is emitted and subsequently detected at detector 200 ! .
• an approximate location of the radionuclide at the time of its radioactive decay (i.e. the source of the primary photon emission) can be estimated as being somewhere along a line of flight 222.
• the approximate location of the photon emission event may be estimated as being within the three dimensional region 224, which may be a substantially cylindrical or conical area.
• the estimation of an approximate location for radionuclide 210 based solely on the detected primary photon 220 is similar to SPECT imaging. In particular, without further information (e.g. depth information) it is impractical, if not impossible, to further localize the position of the photon emission event.
• FIG. 3 is an illustrative schematic example of dual photon imaging in which a pair of photons is emitted from an interaction (e.g. annihilation) of a positron and an electron.
• a positron is emitted from radionuclide 211.
• This positron travels until it loses enough of its kinetic energy to interact with (i.e. annihilate) an electron; the positron may travel a distance of up to a few millimeters from its point of emission.
• This positron- electron annihilation event results in a pair of secondary photons 230a and 230b travelling in opposite directions (e.g.
• the energies of photons 230a and 230b are expected to be approximately 51 1 keV; that is, photons 230a and 230b each have a secondary expected energy of, in this example, approximately 51 1 keV.
• these secondary photons are subsequently detected at detectors 200 3 and 200g, respectively. Based on the locations on photon detectors 200 3 and 200 9 at which these secondary photons 230a and 230b are detected, an approximate location of the positron -electron annihilation event (i.e. the source of the secondary photon emission) can be estimated as being somewhere along a line of flight 232.
• the estimated location of the positron-electron annihilation event can be used as a proxy for an approximate location of the radionuclide 21 1 at the time of its radioactive decay.
• the average distance a positron travels from its point of emission before it is annihilated can be estimated based on the expected decay energies of radionuclide/isotope 2 1 ).
• d 3 _ 9 is the distance between the two detectors 200 3 and 200g in Figure 3
• t 3 is the time at which the photon 230a is registered at detector 200 3
• t 9 is the time at which the photon 230b is registered at detector 200 9
• x be the distance between the detector 200 3 and the point at which the photons originated and let tong be the time of the annihilation event (i.e. the time at which the photons originate).
• photon 230a travels a distance x in the time t 3 - t or igi n
• photon 230b travels a distance d 3 - 9 - x in the time t 9 - Wigm to reach detector 200 9 .
• c is the speed of a photon (i.e. approximately 3 x 10 ⁇ 8 m/s, or 0.3 m/ns).
• d 3- 9 is the fixed distance between the detectors (i.e. a known quantity)
• the distance between the detector 200 3 and the point at which the photons originated i.e. the point of emission of the photons along the line of flight 232
• the distance between the detector 200 3 and the point at which the photons originated can be determined based on the time difference of the detection of the two photons 230a and 230b at the respective detectors 20O3 and 200 9 ,
• the approximate location of the positron-electron annihilation event may be estimated in three dimensions as region 234. which may be a substantially spherical area. That is, the estimation of a location for radionuclide 211 is based on detected secondary photons 230a and 230b.
• positron emission and photon emission occur at effectively the same place (since an excited 'daughter' nucleus barely moves before the photon is emitted).
• the positron itself may move up to about a few millimeters before it annihilates with an electron, resulting in the emission of two 51 1 keV photons.
• an approximation of the location of a radionuclide can be made by presuming a common emission time and a common point of origin for a primary photon and at least one secondary photon emitted from the annihilation of the positron.
• the detector location and timing information for the secondary photons can be used to approximate a locus for the positron annihilation region
• the detector location and timing information for the primary photon can be used to approximate a cylindrical or conical region for the location of the radionuclide at the time of its radioactive decay (i.e. at the time of the emission of the primary photon).
• the intersection of these two regions may localize the decay point of the radionuclide to a much narrower region.
• Such an approximation can be expected to be a significant improvement in comparison to estimates based solely on detection of a primary photon (e.g. SPECT imaging), or solely on detection of at least one secondary photon (e.g. PET imaging).
• FIG. 4 This can be seen, for example, in Figure 4, in which a number of photon detectors 200 are directed towards a target region 205 that contains a radionuclide 212. As this radionuclide 212 undergoes radioactive decay, a primary photon 220 having an energy ⁇ (referred to herein as a primary energy) is emitted and subsequently detected at detector 200i. As explained above with reference to Figure 2, based on the location of photon detector 200i at which this primary photon 220 is detected, an approximate location of the radionuclide 212 at the time of its radioactive decay (i.e. the source of the primary photon emission) can be estimated as being within the three dimensional region 224.
• a primary energy a primary photon 220 having an energy ⁇
• an approximate location of the radionuclide 212 at the time of its radioactive decay i.e. the source of the primary photon emission
• positron is also emitted from radionuclide 212.
• This positron travels until it loses enough of its kinetic energy to interact with (i.e. annihilate) an electron; the positron may travel a distance of up to a few millimeters from its point of emission.
• This positron-electron annihilation event results in a pair of secondary photons 230a and 230b travelling in opposite directions (e.g. 180 degrees apart), each having an energy of Y2 (expected to be approximately 511 keV).
• the time of emission of the primary photon 220 and the time of emission of the secondary photons 230a and 230b is highly correlated.
• the location of emission of the primary photon 220 and the location of emission of the secondary photons 230a and 230b is also highly correlated. If the primary photon and the secondary photons are considered to originate from the same location (i.e. have a common point of origin), the approximate location of this point of origin can be determined based on the intersection 244 of areas 224 and 234. That is, an approximate location for radionuclide 212 at the time of its radioactive decay can be determined based on both the primary photon 220 and the secondary photons 230a and 230b.
• Figures 2 to 5 are only illustrative schematic examples, and actual detector systems may vary. For example, there may be more or fewer photon detectors 200 in each detector ring, and there may be more than one detector ring located about a common axis, forming a cylinder of detector elements.
• Radionuclides may be introduced into a target region of a detector array, and that a volumetric image may be reconstructed based on an analysis of approximate locations determined for a number of decay events detected within the target region.
• a detector system similar to what is currently being used in PET imaging may be used.
• a cluster of scintillation detectors or solid state detectors or combinations thereof 200 ⁇ , 2002, 200N may be used to generate fast time signals and linear energy signals.
• a schematic block diagram of such a detector system 300 is shown in Figure 6.
• the blocks 200i , 2 ⁇ 2, 200 M are photon detectors, each sending analog signals carrying energy information to amplifiers 310i , 310 2 310 N , which may be linear amplifiers.
• Amplifiers 310 shape and amplify the signals received from the detectors 200, and preferably maintain a linear relationship between the output amplitude and the detector output.
• the output of amplifier 310 - generally a slow pulse of a few microseconds width ( ⁇ 2-5 microseconds)- with appropriate time delay conditions is fed as input to an Analog Digital Converter (ADC), as discussed further below.
• ADC Analog Digital Converter
• the photon detectors 200i , 200 2 , 200 N also each send fast logic signals with time information to discriminators 320i, 320 2 , 320 ⁇ . These - discriminators generate a fast output ( ⁇ 10 nano seconds) logic signal upon receipt of input from the detector.
• a threshold level can be programmed or otherwise set. The threshold level is a voltage above which the discriminator is operative to generate an output, and in some embodiments may be as low as a few tens of milli Volts.
• Discriminators 320 generate two or more outputs, at least one of which will be analyzed by the majority logic unit (MLU) 330 to determine the number of detectors outputting signals.
• MLU majority logic unit
• the MLU 330 accepts the logic signals from the discriminators 320.
• the timing resolution of MLU 330 is generally determined by the pulse widths of the input signals received from the discriminators, which may be approximately 10 nanoseconds.
• MLU 330 may comprise a Field Programmable Gate Array (FPGA) for versatility of logical analyses in a simple setting.
• FPGA Field Programmable Gate Array
• the logic condition of MLU 330 may be configured such that it generates output for a photon detection event that corresponds to a preset logic requirement. For example, MLU 330 may generate an output signal if 1 , 2, or 3 photons are simultaneously (or near-simultaneously) detected - that is, a detection event may occur when 1 , 2, or 3 logic inputs are received from discriminators 320, Alternatively, MLU 330 may be configured to generate an output signal if a minimum number of time adjusted input signals from the discriminators are received (e.g. if at least 1 , 2, or 3 input signals are received). The conditions for satisfying a detection event may also include a minimum energy level output by a detector 200, or other conditions.
• the MLU signal is fed to a Gate Generator (GG) 340 which may perform pulse shaping (polarity, height, shape and width) on the signal received from the MLU 330 to serve as trigger pulses to the Time to Digital Converter (TDC) 360.
• GG 340 may be integrated into MLU 330.
• TDC 360 registers time information for the photon detection events.
• TDC 360 is triggered by the signals received from GG 340, and also receives time- adjusted logical signals from the discriminators 320 ! , 320 2 ... 320N as input signals.
• the timing information registered by TDC 360 is relative to the GG trigger pulse and arrival times of all individual photon detection signals and an address identifying the detector at which the photon detection signal was detected (e.g. at which detector 200 2 ⁇ 2, ... 200 the signal was registered) are recorded.
• the output signal from GG 340 is also fed as a trigger to the Analog Digital Converter (ADC) 350, which also receives time-adjusted linear signals from the amplifiers 310i, 310 2 , 310 N . Receipt of a trigger pulse from GG 340 sets ADC
• ADC 350 in an "active mode" for a predetermined time interval, during which ADC 350 receive pulses from amplifiers 310 as inputs.
• the conversion times of the ADC and TDC modules may be of the order of a few microseconds.
• the predetermined time interval for recording a detection event may last from about 100 nanoseconds to about a few microseconds.
• ADC 350 While ADC 350 is in active mode, it converts the analog pulses from the amplifiers 310 containing energy information into digital pu!se signals that may be stored along with the corresponding detector addresses and arrival times registered by the TDC 360. This information - i.e. the arrival time(s), detector location(s), and energy information for each detection event that satisfies the logic conditions of MLU 330/GG 340 - may be fed to a computer 370 for further data processing.
• Figure 6 is only an illustrative schematic example, and actual detector systems and electronic arrangements may vary.
• FPGAs or other hardware or software logic elements may be used in combination to detect and record photon emissions from target region 205.
• detector system 300 may include a processor or processors (not shown) and computer readable storage media (not shown) with instructions stored thereon to instruct the processor(s) to register the energy and time of arrival information of photons at each detector, with data being recorded in list-mode. For example, this information may be sent to a list-mode data buffer. In the list-mode, the following information may be recorded for each detection event: i) the number of photon detectors outputting an energy signal indicative of a photon detection; ii) the identity (e.g.
• the hardware logic of a typical PET detector system may be modified to allow for list-mode or other forms of data collection.
• each detection event may be recorded in, for example, a data array.
• This data array may include the following information; Number (e.g. ID) of each detector registering a photon, the time of arrival of that photon at that detector, the energy deposited by the photon at the detector.
• Number e.g. ID
• a person skilled in the art will recognize that more or less information may be recorded.
• the data may be analyzed to distinguish detection events based on, for example, the relative time of arrival of two or more photons, the energies of detected photons, or both.
• this detected photon is a primary photon (e.g. analogous to a SPECT photon) or one of a pair of secondary photons (e.g. one of a pair of PET photons, the other having not been successfully detected) based on the energy information recorded for that detection event.
• the relative energies of the detected photons can be used to confirm that a primary photon and a pair of secondary photons - presumed to originate from the same decay event - have been detected.
• a primary photon and at least one secondary photon are detected within the predetermined time interval (i.e. their emission is temporally correlated), as discussed above they may be presumed to share a common point of origin ⁇ i.e. their emission locations are spatially correlated) and an approximate location for the radionuclide at the time of its radioactive decay can be determined based on this presumption.
• a volumetric reconstruction e.g. tomographic reconstruction
• such a data set may reduce, or correct for, one or more imaging 'artifacts' of concern to radiologists that may be present in images reconstructed in typical PET or SPECT systems.
• the list-mode data may be parsed to separate the detection information for photons having the primary energy from the detection information for photons having the secondary energy (e.g. 51 1 keV). In this way, either a SPECT image or a PET image (or both) may be separately reconstructed from the data set collected during a single scanning procedure.
• the secondary energy e.g. 51 1 keV
• the list-mode data set may be analyzed to identify detection events during which three photons were detected, and these three-photon detection events may be further analyzed to identify detection events during which one primary photon and two secondary photons were detected.
• a volumetric reconstruction e.g. tomographic reconstruction
• a volumetric reconstruction of such a data set can be expected to provide a more accurate representation - based in the improved approximate locations of the radionuclides - than images reconstructed in typical PET or SPECT systems.
• Such a comparison may be useful in identifying imaging artifacts or other undesirable imaging sensitivities that may not be identifiable by traditional imaging systems.
• the data set of three-photon (one primary, two secondary) detection events can be analyzed to remove or otherwise ignore the information regarding the secondary photons and reconstruct a volumetric image based only on the detected primary photons (as in typical SPECT imaging) and compare this one-photon PET image with the three-photon image. Again, such a comparison may be useful in identifying imaging artifacts or other undesirable imaging sensitivities that may not be identifiable by traditional imaging systems.
• the image analysis performed will be similar to that of typical SPECT and PET software.
• various exemplary types of imaging are outlined as follows.
• a first type of imaging [Type A] one photon is detected during the predetermined time interval. If the detected photon has a secondary energy (i.e. the photon is a product of a positron-electron annihilation) an approximate location of the positron at the time of its annihilation may be determined assuming that the other secondary photon went undetected. While the approximate location based on a single photon may be determined as discussed above with reference to Figure 2, it may be possible to take into consideration the different expected attenuation for a 51 1 keV photon (compared to the expected attenuation for a lower energy photon). Reconstructing a volumetric image based on detected photons having a secondary energy (e.g. 51 1 keV) may be regarded as Type A (high energy) SPECT imaging.
• Type A high energy
• the system may employ a SPECT imaging mode - i.e. determine approximate locations for decay events based on one detected photon per decay event - and the resulting image will be similar to a conventional SPECT image.
• Reconstructing a volumetric image based on detected photons having a primary energy may be regarded as Type A (low energy) SPECT imaging,
• Comparing and/or combining Type A (high energy) and Type A (low energy) images may therefore constitute a form of dual energy SPECT imaging.
• a second type of imaging [Type B] two photons are detected during the predetermined time interval. If both detected photons have the expected energy of a secondary photon (e.g. 51 1 keV), the system will employ a PET imaging mode - i.e. determine approximate locations for decay events based on two detected photons per decay event - and the resulting image will be similar to a conventional PET image. For example, as shown in Figure 5, if secondary photons 230a and 230b are detected at detectors 200 3 and 200 7 , respectively, but primary photon 220 is not detected, an approximate location for the radionuclide may be estimated in three dimensions as region 234a.
• a PET imaging mode i.e. determine approximate locations for decay events based on two detected photons per decay event -
• an approximate location for the radionuclide may be estimated in three dimensions as region 234a.
• a location of the radionuclide at least as accurate as in Type A may be determined. For example, as shown in Figure 5, if primary photon 220 is detected at detector 200i and secondary photon 230a is detected at detector 200 3 , but secondary photon 230b is not detected, an approximate location for the radionuclide at the time of its decay may be estimated in three dimensions as region 244a (i.e. the intersection of areas 224a and 224b).
• a third type of imaging [Type C] three photons are detected during the predetermined time interval, and information for each of the three photons is used to determine approximate locations of the decay events (as discussed above with reference to Figure 4). For example, as shown in Figure 5, if all three photons 220, 230a and 230b are detected at detectors 200i , 200 3 , and 200 7 , respectively, an approximate location for the radionuclide at the time of its decay may be estimated in three dimensions as region 244b (i.e. the intersection of areas 224a and 234b, region 234b being estimated based on the presumed correlation between the emission location of the primary photon and the positron annihilation event, which implies that the region 234b and 224a should intersect).
• region 244b i.e. the intersection of areas 224a and 234b, region 234b being estimated based on the presumed correlation between the emission location of the primary photon and the positron annihilation event, which implies that the region
• a fourth type of imaging [Type D]
• three photons are detected during the predetermined time interval, as in Type C.
• information from only two of the three detected photons is used to determine approximate locations of the decay events. That is, the detection information recorded for a primary photon may be suppressed or ignored, and an approximate location of the decay event may be based on the two secondary photons, as discussed above.
• detection information recorded for one of the two detected secondary photons may be suppressed or ignored, and an approximate location of the decay event may be based on the primary photon and the other one of the two detected secondary photons (as in Type B).
• Type D the detection information recorded for a primary photon or a secondary photon is not utilized in the image reconstruction.
• the differences in the images obtained in this manner and the images reconstructed using the Type C imaging may assist in the identification and/or classification of imaging artifacts.
• images produced using Type D imaging may be compared with images produced using Type B imaging to verify the robustness of the imaging algorithms - as the Type B and Type D images should be similar (if not identical) as they are based on the same information, consistency between the Type B and Type D images may increase the confidence in Type B imaging where the additional Type D data is unavailable.
• a fifth type of imaging [Type E] three photons are detected during the predetermined time interval, as in Types C and D.
• information from only one of the three detected photons is used to determine approximate locations of the decay events.
• the detection information recorded for the secondary photons may be suppressed or ignored, and an approximate location of the decay event may be based on the primary photon, as in SPECT imaging.
• detection information recorded for the primary photon and one of the two detected secondary photons may be suppressed or ignored, and an approximate location of the decay event may be based on the other secondary photon (as in Type A). That is, in Type E the detection information recorded for two of the three detected photons is not utilized in the image reconstruction.
• images produced using Type E imaging may be compared with images produced using Type A imaging to verify the robustness of the imaging algorithms - as the Type A and Type E images should be similar (if not identical) as they are based on the same information, consistency between the Type A and Type E images may increase the confidence in Type A imaging where the additional Type E data is unavailable.
• Imaging types may also contribute to improving detector assemblies and/or imaging algorithms, For example, images generated using data collected at a first section of a detector assembly (e.g. the upper half of a cylindrical detector array) may be compared to images generated using data collected from a second section (e.g. the lower half of the array). If Type A (high energy) images generated using data collected at the first and second sections appear identical (or nearly identical), but appear different from Type A (low energy) images generated using data collected at the first and second sections, we may attribute the differences to energy-dependent attenuation artifacts. However, if the Type A (high energy) image generated using data collected at the first section appears different from the Type A (high energy) image generated using data collected at the second section, detector geometry may be considered a contributing factor.
• the proposed imaging types may allow for cross-checking image information in multiple ways, which may reduce ambiguity in interpreting the data. For example, this may involve recovering PET-only and SPECT-only imaging in order to assess the consistency of the data and to assess energy dependent artifacts of scattering and absorption in different physiological zones of a patient.
• the primary (SPECT) photons are subject to absorption in bone, however they typically pass through soft tissue with scattering - not absorption - effects.
• the secondary (PET) photons of higher energy typically are not significantly attenuated by passing through soft tissue, Secondary photons are also likely to pass through bone also without significant absorption, but they may be subject to scattering.
• a medical practitioner may be presented with images based on one or more of the above imaging types on different monitors. The practitioner may then examine these images and critically evaluate the influence of imaging artifacts and arrive at more reliable diagnostic information. This may provide a better check of the computer imaging algorithms for their correctness and accuracy, and may provide an improved basis for the evaluation of influences of different physiological components on the passage of gamma radiation through tissue and bone. This information may reduce diagnostic errors relating to either false positives (scattering), false negatives (absorption), or other errors induced by imaging artifacts.
• a single radioisotope that emits positrons accompanied by photon emissions of the daughter product is used.
• a radioactive isotope should meet several criteria.
• the half-life of the radionuclides should be neither too short nor too long (optimal ones are about a few hours). If the half-life is too short, there is not enough time for the radionuclides to be processed and injected into a patient, and if the ha!f-life is too long, the residual radio-activities after imaging maybe too high and be of adverse health concerns.
• the production rate (which is a product of the isotope abundance and production cross section) should be high enough to ensure a sufficient amount of the isotope can be manufactured for each scan.
• the primary photon should be expected to have an energy sufficiently different from the secondary photons for contrast reasons.
• the primary photon has an expected energy lower than the expected energy of the secondary photons (e.g. less than about 511 KeV), as a photon with too high of an energy will simply pass through the bone and soft tissue without any contrast.
• the primary photon should have a comparable intensity to the secondary photons to accomplish an effective PET/SPECT photon imaging.
• the isotope should be compatible with the human body.
• it should be able to be incorporated into a larger radioactive tracer molecule (e.g. a metabolic substrate, receptor ligand, antibody, organic molecule, amino acid, or polypeptide), administrable directly (e.g. as a soluble ionic salt), or otherwise be incorporated into a radiopharmaceutical.
• a radioactive tracer molecule e.g. a metabolic substrate, receptor ligand, antibody, organic molecule, amino acid, or polypeptide
• administrable directly e.g. as a soluble ionic salt
• threshold energy for reaction energy of protons or gammas should be greater than threshold energy
• Attenuation Ratio 1 /
• the 123 Xe isotope 2.08 hours) can be produced by 127 l(p,5n) 123 Xe reaction in which protons wiil be bombarded on a 127 i target (100% natural abundance).
• the decay product 23 l is a commonly used isotope for thyroid cancer diagnosis.
• the threshold energy of protons for this reaction is 37.1 MeV.
• Figure 7 shows the neutron cross section (a measure of production rate) plotted against proton energy. As shown, it is reasonable to assume that sufficient amounts of 123 Xe can be produced with 55-60 MeV proton beam currents of several micro amperes.
• 123 Xe may improve the imaging compared to single photon imaging using the 142 keV photon of esm Tc, typically used in current SPECT imaging.
• xenon is administered to patients for xenon enhanced CT as it propagates in the human body.
• This technique is advocated for brain imaging as it is not limited by the blood-brain barrier (see for example N. iyazawa, M. Uchida ; A. Fukamachi, I. Fukasawa, H. Sasaki, and H. Nukui, American Journal of Neuroradiology, (1999), vol. 20, p. 1858-1862).
• the decay product 123 l is a commonly employed SPECT isotope for thyroid imaging.
• 23 Xe isotope has a great promise to be useful for whole body PET/SPECT imaging.
• the threshold for this reaction is 22.0 MeV, with a maximum at proton energies of about 34 MeV (see Figure 8).
• 73 Se In the decay of 73 Se, the positron emission is accompanied by 67 keV and 361 keV photons. Thus, it provides a very interesting contrast of very low energy (67 keV) : and somewhat high energy (361keV) photons in addition to the 51 1 keV photons which result from the positron annihilation. Accordingly, 73 Se could prove very useful in identifying soft tissue, dense material and bones distinctly.
• the 73 Se isotope looks promising for use in measuring and assessing the bile acid turnover in the intestines, based on the current use of SeHCAT (selenium homocholic acid taurine or tauroselcholic acid), a radiopharmaceutical incorporating the gamma emitter 75 Se. Also, the National Oncology Research Institute is carrying out extensive research on selenium as a selective chemotherapeutic agent, indicating that administering selenium to humans or animals has therapeutic applications. Thus, selenium isotopes appear to be good candidates for diagnostic purposes.
• SeHCAT selenium homocholic acid taurine or tauroselcholic acid
• Valette and coworkers produced Metabromobenzy!guanidine (MBBG), a bromine compound as a PET radiotracer (H. Valette, C. Loc'h, K. Mardon, B. Bendriem, P. Merlet, C. Fuseau, S. Sabry, D. Raffel, B. Maziere, and A. Syrota, The Journal of Nuclear Medicine, Vol. 34 (1993) p. 1739-1744.).
• Bromine belongs to the same group as fluorine, the radioisotope used in FDG for PET imaging. Therefore, as fluorine and bromine belong to the same group in periodic table, they can be substituted each other.
• MBBG does to heart imaging what FDG does to metabolically active tissues
• MBBG can be substituted employed for heart imaging, with the added advantage of the dual PET/SPECT imaging described above, as opposed to the typical PET imaging of FDG.
• the systems, processes and methods of the described embodiments are capable of being implemented in a computer program product comprising a non- transitory computer readable medium that stores computer usable instructions for one or more processors that cause the one or more processors to operate in a specific and predefined manner to perform the functions described herein.
• the medium may be provided in various forms, including as volatile or non-volatile memory provided on optical, magnetic or electronic storage media.

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