Signal processing system, positron emission tomography apparatus, and positron emission tomography method

Abstract

The present invention addresses the problem of providing a signal processing system, a positron emission tomography device, and a positron emission tomography method that achieve an improved S / N value by improving detection efficiency compared to what is conventional, reducing the total amount of radiation through the improvement in detection efficiency, and consequently reducing accidental coincidence counting. This signal processing system creates image data on the basis of a group of electrical signals output from a radiation detector. The signal processing system recognizes the group of electrical signals as signals to be processed. The group of electrical signals includes at least a portion of a group of electrical signals that satisfies the following requirements: the group of electrical signals correspond to gamma rays having an energy of less than or equal to 375 keV and have a signal value in a predetermined range, wherein the predetermined range is 50%-80% of a 100% signal value, the 100% signal value being a signal value that is detected when a gamma ray having an energy of 511 keV enters a radiation detecting element in the radiation detector and is totally absorbed by the radiation detecting element.

Description

[Technical Field]

[0001] The present invention relates to a positron emission tomography (PET) apparatus for use in a high count radiation detection device, for example. [Background technology]

[0002] Positron emission tomography (PET) and other procedures use positron emission tomography devices that measure high-energy radiation such as gamma rays. In nuclear medicine, for example, a substance (tracer) in which some elements of molecules localized in cancer cells are replaced with positron-emitting nuclides (radiation sources) is injected into the patient, and the radiation generated from the tracer inside the patient's body is measured and the location of its generation is identified, making it useful in diagnostic devices that predict the location of cancer cells. Generally, a positron emission tomography apparatus includes a scintillator unit having a scintillator that receives radiation and emits electromagnetic waves such as visible light, a conversion output unit that receives the electromagnetic waves emitted by the scintillator, converts the received electromagnetic waves into electrical signals, and outputs the electrical signals, and a signal processing system that converts the electrical signals output by the conversion output unit into image data. Positron emission tomography (PET) systems utilize the phenomenon in which two radiation beams are emitted simultaneously and in opposite directions due to electron-positron annihilation in the tracer or other components in the subject. When radiation beams are incident on and detected by two scintillators at different locations at approximately the same time (within a specified time interval), it is considered that "two radiation beams were generated at the midpoint of the line connecting the two scintillators according to the above principle," and the position of the radiation source is calculated from the positions of the scintillators. (Hereinafter, this detection method will be referred to as the "coincidence counting" method, and the number of coincidence counts will be referred to as the "number of coincidence events.") Furthermore, by recording the frequency of radiation beams generated at each location where radiation is presumed to have been generated, the relative concentration distribution of the radiation source in the subject can be measured, and an image can be synthesized from this information. The specified time interval will also be referred to as the "time window."

[0003] In positron emission tomography (PET) systems, in addition to coincidence counts (sometimes referred to as true coincidence counts) resulting from electron-positron annihilation near the tracer, various noises that appear as coincidence coefficients are generated. Examples of noise include radiation emitted by the scintillator itself (hereafter sometimes referred to as "intrinsic background radiation"), background radiation generated both inside and outside the system, "backscatter" (measured radiation scattered by the system walls before entering the scintillator), "accidental coincidence" (two radiation beams generated at different locations accidentally strike different scintillators at the same time), and "scattered coincidence" (two radiation beams generated simultaneously and in opposite directions from the same radiation source undergo Compton scattering, resulting in changes in angle and energy). These noises interfere with accurate measurements and must be appropriately filtered out. However, scattered coincidence can also be used as a signal, although the positional information is subject to uncertainty.

[0004] For this reason, methods have been reported to accurately obtain the location information of the radiation source by improving the signal-to-noise ratio (S / N value), and methods have also been reported to reduce the subject's exposure by increasing the number of events per total radiation dose. For example, in Non-Patent Document 1, since the energies of inherent background radiation (gamma rays), backscattered gamma rays, and gamma rays generated in scattered coincidence events differ from the energy of gamma rays during true coincidence counting (511 keV), a circuit has been developed that automatically treats luminescence with an energy below a certain threshold as noise and does not collect data. [Prior art documents] [Non-patent literature]

[0005] [Non-Patent Document 1] IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 63, (2016) pages1327 - 1334 Summary of the Invention [Problem to be solved by the invention]

[0006] However, in the previously reported coincidence measurement devices, the signal processing system removes, as noise, electrical signals that are very close to the signal value of the electrical signal corresponding to the initial energy of the radiation generated from the radiation source (hereinafter sometimes referred to as the initial energy), for example, when the initial energy is 511 keV, electrical signals that are equal to or less than about 500 keV. As a result, there are cases where scattered coincidence counts whose energy has been reduced by Compton scattering cannot be used as a signal, and as the number of events generated per radiation dose decreases, it is necessary to increase the total radiation dose to obtain clear images, which increases the subject's exposure.In addition, if the total radiation dose exceeds a certain value, the S / N value may drop sharply due to a sudden increase in random coincidence counts. Alternatively, when the time window is set large to increase the number of events, random coincidence counts cannot be properly excluded, resulting in a problem of a decrease in the S / N value. Therefore, an object of the present invention is to provide a signal processing system, a positron emission tomography apparatus, and a positron emission tomography method that improve detection efficiency compared to conventional systems, and further reduce the total radiation dose due to the improved detection efficiency, thereby reducing random coincidence counts and improving the S / N value. [Means for solving the problem]

[0007] As a result of intensive research in light of the above problems, the inventors discovered that by using a signal processing system that processes electrical signals whose signal values ​​are within a predetermined range, or at least a portion of electrical signals corresponding to gamma rays having an energy within a predetermined range, it is possible to increase the number of events per generated radiation dose and reduce the radiation exposure of the subject, leading to the completion of the present invention.

[0008] That is, the gist of the present invention includes the following. [1] A signal processing system that creates image data based on a group of electrical signals output from a radiation detector, the signal processing system recognizes the group of electrical signals as a processing target; a signal processing system, wherein the set of electrical signals includes at least a portion of the set of electrical signals that meets the following requirements: A group of electrical signals corresponding to gamma rays having an energy of 375 keV or less, whose signal values ​​are within a predetermined range, the predetermined range is 50% or more and 80% or less of a signal value of 100%, The 100% signal value is a signal value detected when a gamma ray having an energy of 511 keV is incident on a radiation detection element in a radiation detector and is totally absorbed by the radiation detection element. [2] A signal processing system that creates image data based on a group of electrical signals output from a radiation detector, the signal processing system recognizes the group of electrical signals as a processing target; A signal processing system, wherein the group of electrical signals includes at least a portion of a group of electrical signals corresponding to gamma rays having energy values ​​within a predetermined range, the predetermined range being equal to or greater than 232 keV and equal to or less than 340 keV. [3] A positron emission tomography apparatus comprising the signal processing system according to [1] or [2] and a radiation detector unit. [4] The positron emission tomography apparatus according to [3], wherein the radiation detector unit has the following elements: A scintillator section including a scintillator that receives radiation and emits electromagnetic waves, and a conversion output section that receives the electromagnetic waves emitted from the scintillator, converts the received electromagnetic waves into pulsed electrical signals, and outputs the signals. [5] The positron emission tomography apparatus according to [4], wherein the scintillator satisfies the following characteristics: The intensity of the inherent background of the scintillator is 200 Hz / cm in the range of signal value 10% to 120%, where the signal value of the pulsed electric signal when a gamma ray having an energy of 511 keV is incident on the scintillator and is totally absorbed by the scintillator is 100%.3 The following is the result. [6] A positron emission tomography apparatus according to [4] or [5], wherein the time window in the conversion output section is 180 ns or less. [7] The positron emission tomography apparatus according to any one of [4] to [6], wherein the scintillator has a fluorescence decay time (DT) of 25 ns or less when irradiated with gamma rays. [8] The positron emission tomography apparatus according to any one of [4] to [7], wherein the gamma ray absorption rate of the scintillator is 70% or more. [9] The positron emission tomography apparatus according to any one of [4] to [7], wherein the gamma ray absorption rate of the scintillator is 50% or less.

[10] A signal processing method for creating image data based on a group of electrical signals output from a radiation detector, comprising: The signal processing method includes recognizing the group of electrical signals as a processing target; a signal processing method, wherein the set of electrical signals includes at least a portion of the set of electrical signals that meets the following requirements: A group of electrical signals corresponding to gamma rays having an energy of 375 keV or less, whose signal values ​​are within a predetermined range, the predetermined range is 50% or more and 80% or less of a signal value of 100%, The 100% signal value is a signal value detected when a gamma ray having an energy of 511 keV is incident on a radiation detection element in a radiation detector and is totally absorbed by the radiation detection element.

[11] A signal processing method for creating image data based on a group of electrical signals output from a radiation detector, comprising: the signal processing method recognizes the group of electrical signals as a processing target; A signal processing method, wherein the group of electrical signals includes at least a portion of a group of electrical signals corresponding to gamma rays having energy values ​​within a predetermined range, the predetermined range being equal to or greater than 232 keV and equal to or less than 340 keV.

[12] A positron emission tomography method comprising at least the following steps (a), (b) and (c): (a) a scintillation step in which radiation is converted into electromagnetic waves using a scintillator that receives radiation and emits electromagnetic waves; (b) a conversion / output step of receiving electromagnetic waves emitted from the scintillator, converting the received electromagnetic waves into a pulsed electric signal, and outputting the signal; and (c) A signal processing step including a step of performing signal processing by the signal processing method according to

[10] or

[11] .

[13] The positron emission tomography method according to

[12] , wherein the scintillator satisfies the following characteristics: The intensity of the inherent background of the scintillator is 200 Hz / cm in the range of signal value 10 to 120%, where the signal value of the pulsed electric signal when a gamma ray having an energy of 511 keV is incident on the scintillator and is totally absorbed by the scintillator is 100%. 3 The following is the result.

[14] The positron emission tomography method according to

[12] or

[13] , wherein the time window in the conversion output unit is 180 ns or less.

[15] The positron emission tomography method according to any one of

[12] to

[14] , wherein the scintillator has a fluorescence decay time (DT) of 25 ns or less when irradiated with gamma rays.

[16] The positron emission tomography method according to any one of

[12] to

[15] , wherein the gamma ray absorption rate of the scintillator is 70% or more.

[17] The positron emission tomography method according to any one of

[12] to

[15] , wherein the gamma ray absorption rate of the scintillator is 50% or less. [Effects of the Invention]

[0009] The present invention provides a signal processing system, a positron emission tomography apparatus, and a positron emission tomography method that have improved detection efficiency compared to conventional systems. Furthermore, the improved detection efficiency reduces the total radiation dose, thereby reducing random coincidences and improving the S / N ratio. Furthermore, by using a scintillator material with a small intrinsic background, it is possible to provide a positron emission tomography device with a dramatically improved S / N value, and furthermore, by eliminating the need for radiation absorption in the scintillator and reducing the thickness of the scintillator, it is possible to provide a miniaturized positron emission tomography device and positron emission tomography method. [Brief explanation of the drawings]

[0010] [Figure 1] 1 is a cross-sectional view of a simulation space simulating a radiation imaging apparatus according to an embodiment of the present invention, taken along a Y-axis-Z-axis plane. [Figure 2] 1 is a cross-sectional view of a simulation space simulating a radiation imaging apparatus according to an embodiment of the present invention, taken along an XY plane. DETAILED DESCRIPTION OF THE INVENTION

[0011] The following describes in detail the embodiments of the present invention, but these descriptions are examples (typical examples) of the embodiments of the present invention, and the present invention is not limited to these contents as long as it does not deviate from the gist of the invention. In this specification, the term "signal value" refers to a parameter expressed as a relative numerical value, so-called signal intensity, obtained by logarithmically expressing the time integral value of a pulsed electric signal. In this specification, a numerical range expressed using "to" means a range that includes the numerical values ​​written before and after "to" as the lower and upper limits, and "A to B" means A or more and B or less. In addition, in this specification, "plurality" means "two or more." In addition, several embodiments will be described below, but the conditions of each embodiment can be applied to other embodiments to the extent that they are applicable.

[0012] <Positron emission tomography equipment> In one embodiment, the present invention is a positron emission tomography apparatus. Hereinafter, a positron emission tomography apparatus according to one embodiment of the present invention may be simply referred to as a "positron emission tomography apparatus." A positron emission tomography apparatus includes a plurality of radiation detectors that convert radiation into electrical signals, and a signal processing system that creates image data based on a group of electrical signals made up of the radiation (and converts the group of electrical signals into an image by signal processing).

[0013] (Radiation detector section) In the positron emission tomography apparatus, the radiation detector receives radiation and outputs an electrical signal. The configuration of the radiation detector is not particularly limited, but it usually includes a radiation detection element that receives radiation and outputs an electrical signal, and is preferably an array including a plurality of such radiation detection elements from the viewpoint of obtaining an image based on positional information of the radiation. In this specification, a radiation detection element refers to the smallest unit (e.g., a unit equivalent to one pixel) that has the function of converting radiation into an electrical signal.

[0014] In one embodiment, the radiation detector unit has at least a scintillator unit including a scintillator that receives radiation and emits electromagnetic waves, and a conversion output unit that receives the electromagnetic waves emitted from the scintillator, converts the received electromagnetic waves into pulsed electrical signals, and outputs the signals. In this embodiment, the radiation detection elements include one scintillator (scintillator element) corresponding to each radiation detection element, and a conversion output section (conversion output element) corresponding to the one scintillator.

[0015] In another embodiment, the radiation detector unit includes at least (a) one or more radiation detection elements, each of which has a semiconductor member that receives radiation and generates electrons or holes, and a pair of electrodes arranged to sandwich the semiconductor member, and (b) a transmission unit that transmits the electrons or holes detected by each of the radiation detection elements as an electrical signal to a signal processing system. By applying a voltage between the two electrodes of the radiation detection element, the electrons or holes generated in the semiconductor element reach the electrodes, and radiation can be detected as an electrical signal. In this embodiment, the radiation detection element is an element that includes a semiconductor member made of a semiconductor material that can form the smallest unit for radiation detection, and an electrode corresponding to the semiconductor member, similar to a single scintillator (scintillator element) in the embodiment having the scintillator portion, and the element may further include a transmission section corresponding to the semiconductor member.

[0016] The radiation detector unit is preferably arranged to surround the expected position of the subject, and may be concentrically arranged around the subject's position, or on a curved surface equivalent to a cylinder with an equidistant radius around the subject's central axis, or may be arranged on a curved surface that matches the shape of the subject, and may be movable on these curved surfaces during imaging, or may be of a type that allows its arrangement to be changed depending on the subject. The distance between the surface of the subject and the radiation detector unit during use is not particularly limited, but is usually 50 cm or less, preferably 30 cm or less, and more preferably 10 cm or less. The lower limit is not particularly limited, but is usually 1 mm or more. The closer the distance to the subject, the more compact the device can be. In addition, the positron emission tomography apparatus may have an outer wall made of a material with high radiation blocking ability, such as lead, to prevent background radiation from entering from outside and radiation generated within the apparatus from penetrating outside the apparatus.

[0017] Below, as an example of a positron emission tomography apparatus, a case will be illustrated in which the radiation detector unit includes a scintillator unit having a scintillator that receives radiation and emits electromagnetic waves, and a conversion output unit that converts the electromagnetic waves into an electrical signal, but the present embodiment is not limited to this. Furthermore, the arrangement of each device section in the positron emission tomography apparatus is not particularly limited, and can be determined appropriately following the example of known positron emission tomography apparatuses.

[0018] <Scintillator section> The scintillator section (hereinafter sometimes simply referred to as "scintillator section") in a positron emission tomography apparatus is not particularly limited as long as it includes a scintillator and emits electromagnetic waves upon receiving radiation to be measured. The scintillator section may include multiple scintillators. The scintillator section may also be made up of scintillators of the same or different types stacked together, and suitable positional resolution can be obtained even when different types are stacked together. Furthermore, a reflective layer that is reflective to the electromagnetic waves emitted by the scintillator may be provided between the different types of scintillators and / or on the radiation incident surface of the scintillator.

[0019] <Scintillator> The scintillator (hereinafter sometimes simply referred to as "scintillator") provided in the scintillator section is excited by receiving energy and emits electrons or emits light. When emitting light, it is preferable that the scintillator emits light in a wavelength range of 160 nm or more and 700 nm or less, from the viewpoint of increasing the conversion efficiency from electromagnetic waves to electrical signals in the conversion output section. Furthermore, the scintillator has an emission peak wavelength in a wavelength range of preferably 300 nm or more, more preferably 350 nm or more, preferably 600 nm or less, more preferably 550 nm or less. Examples of energy related to the excitation include electromagnetic waves, electron beams, and ionizing radiation. Examples of the ionizing radiation include X-rays, gamma rays, beta rays, alpha rays, and neutron rays, with gamma rays being preferred.

[0020] The radiation incident surface of each scintillator element is not particularly limited and may be, for example, flat or curved. The area of ​​the radiation incident surface of the element is not particularly limited, but is usually 36 mm 2 Less than 9mm, preferably 2 Less than 4mm, preferably 2 The reason is as follows: The smaller the radiation incident surface of the scintillator element, the more improved the spatial resolution can be achieved. Conventionally, the thickness of the scintillator has had to be set large because it is necessary for the scintillator to absorb the entire energy of the radiation, but in this embodiment, the thickness of the scintillator is sufficient as long as it can absorb at least {1-(1 / e)} of the entire energy of the radiation. The thickness of the scintillator refers to the length of the scintillator in the direction of incidence of the radiation. Specifically, the thickness of the scintillator capable of absorbing at least {1-(1 / e)} of the total energy of the radiation is determined appropriately based on the energy and type of the radiation, and the density and effective atomic number of the scintillator, but is usually 1 mm or more, preferably 5 mm or more, more preferably 10 mm or more, and even more preferably 20 mm or more, and is usually 100 mm or less, preferably 60 mm or less, more preferably 50 mm or less, and even more preferably 40 mm or less. More specifically, for example, when the radiation is a gamma ray with an energy of 511 keV and the density and effective atomic number of the scintillator are approximately the same as those of LSO, that is, the density is 7 to 8 g / cm. 3 When the material has an effective atomic number of 60 to 70, the thickness of the scintillator is usually 1 mm or more, preferably 2 mm or more, more preferably 5 mm or more, even more preferably 10 mm or more, and particularly preferably 15 mm or more, and usually 60 mm or less, preferably 50 mm or less, more preferably 40 mm or less, and more preferably 30 mm or less. By being at or above the lower limits of these ranges, radiation absorption efficiency can be ensured, and suitable radiation detection efficiency can be obtained. Furthermore, by being at or below the upper limits of these ranges, suitable position resolution can be obtained. Note that, as described above, the above ranges can be appropriately changed and applied based on the energy and type of radiation, as well as the density and effective atomic number of the scintillator, and therefore, the present embodiment is not limited to the above ranges.

[0021] A scintillator may itself emit radiation. In this specification, such radiation is referred to as intrinsic background radiation. Examples of the intrinsic background radiation include α-rays, β-rays, γ-rays, and X-rays. In addition, indirectly generated radiation, such as when a scintillator emits protons and positrons generate γ-rays, is also considered in this specification as intrinsic background radiation. The inherent background intensity is the intensity of the scintillator 1 cm in a certain energy range. 3 The frequency of radiation generated per second (Hz / cm 3 ) can be expressed as The intensity of the inherent background radiation is typically 350 Hz / cm in an energy band of 0.1 to 1.2 times the initial energy of the radiation to be observed by the positron emission tomography device. 3 Below 250Hz / cm, preferably 3 , more preferably 200Hz / cm 3 or less, more preferably 150 Hz / cm 3 Below 100Hz / cm, particularly preferably 3 The lower limit is not particularly limited, and the lower the value, the better. 3 or more, 10Hz / cm 3 It may be more than that. For example, when gamma rays derived from positron-emitting nuclides are used, the initial energy of the radiation observed by a positron emission tomography device is 511 keV, so the energy band that is 0.1 to 1.2 times this initial energy is 51.1 to 612.1 keV.

[0022] The intensity of the endogenous background radiation can also be observed in the form of a pulsed electric signal that is output after the endogenous background radiation is converted by the conversion output unit. When observed in this manner, the intensity of the endogenous background radiation is typically 350 Hz / cm in a signal value range of 10 to 120%, assuming that the signal value of the pulsed electric signal generated when a gamma ray having an energy of 511 keV is incident on a scintillator and is totally absorbed by the scintillator is 100%. 3 Below 250Hz / cm, preferably 3 Less than or equal to 200Hz / cm 3 or less, more preferably 150 Hz / cm 3 Below, particularly preferably 100 Hz / cm 3 Below 50Hz / cm, particularly preferably 3 Below, particularly particularly preferably 10 Hz / cm 3 or less, most preferably 5 Hz / cm 3 The lower limit is not particularly limited, and the lower the value, the better. 3 or more, and 3 It may be more than that. The smaller the inherent background radiation, the less noise is mixed in, and the lower the energy scattered coincidence counts are, the less noise is confused with the noise, or there is no need to set a high range of the signal to be used as the signal to remove noise.By efficiently using scattered coincidence counts, the number of events relative to the total radiation dose can be increased.

[0023] The intrinsic background intensity can be measured by performing the measurement in a sealed container made of a material with high radiation stopping power, which minimizes environmental radiation such as cosmic rays, and by integrating the number of simultaneous events that emit light due to radiation from the scintillator itself. By calibrating the light emission amount and radiation energy in advance, the radiation energy can be calculated from the light emission amount during measurement, and the intrinsic background intensity for each radiation energy can be determined. The sealed container can be, for example, a container with a lead layer about 10 cm thick and an oxygen-free copper layer about 1 cm thick inside the lead layer.

[0024] <Scintillator fluorescence decay time> The fluorescence decay time (DT) of the scintillator when irradiated with gamma rays is usually 50 ns or less, preferably 35 ns or less, more preferably 25 ns or less, even more preferably 20 ns or less, particularly preferably 15 ns or less, and most preferably 12 ns or less. There is no particular lower limit, but it is usually 0.1 ns or more. The shorter the DT, the better the time resolution of the signal, making it possible to set a shorter time window (described later), thereby eliminating random coincidences and improving the S / N value.

[0025] The DT of the scintillator can be determined by the following method. That is, the intensity of the electromagnetic wave emitted from the scintillator exponentially decays over time, and the electromagnetic wave is converted in the converter / output unit, and the signal value of the output electrical signal also decays exponentially. Therefore, the DT can be determined by plotting the signal value versus time and fitting the curve consisting of the signal value over time using an exponential function.

[0026] The short DT of the scintillator allows the time window to be shortened, and if the intrinsic background intensity is sufficiently small, the time interval at which the intrinsic background radiation occurs becomes sufficiently long compared to the time window, so that background radiation can be excluded from detection in coincidence measurement. In other words, since the presence of the intrinsic background can be ignored, even when radiation with an energy value of less than 511 keV and a certain level or higher, excluding relatively low-energy radiation such as backscattered radiation, is detected, it can be calculated as true coincidence. Specifically, when the DT is 50 ns or less and the intrinsic background intensity is 250 Hz / cm, 3 The more preferred ranges for DT and background intensity are as described above.

[0027] When a scintillator is irradiated with gamma rays, the fluorescence intensity 100 ns after the time when the fluorescence intensity reaches its maximum value is not particularly limited, but is usually 5% or less, preferably 4% or less, more preferably 3% or less, even more preferably 2% or less, and particularly 1.5% or less, with the maximum fluorescence intensity being 100%. The lower limit is not particularly limited, and is usually 0% or more, or 0.001% or more. Since the fluorescence decay is thus very rapid and the fluorescence intensity is sufficiently small after the lapse of a predetermined time, a scintillator material that is useful for radiological testing with high time resolution can be provided.

[0028] <Energy absorption rate for radiation> In order to detect radiation, the scintillator preferably absorbs some or all of the energy of incident radiation. If the proportion of the energy of incident radiation absorbed by the scintillator is defined as the energy absorptivity, in one embodiment of the present invention, similar to conventional positron emission tomography devices, a high energy absorptivity of the scintillator can achieve high radiation detection efficiency. From this perspective, the energy absorptivity of the scintillator, in particular the gamma ray absorptivity, is preferably 50% or more, preferably 70% or more, more preferably 80% or more, and even more preferably 90% or more, with no particular upper limit, which may be 100%.

[0029] In another embodiment, conversely to the above, the low energy absorptivity of the scintillator allows the thickness of the scintillator to be thin, thereby enabling the radiological imaging device to be miniaturized. From this perspective, the energy absorptivity of the scintillator, particularly the gamma-ray absorptivity, is typically 90% or less, preferably 70% or less, more preferably 50% or less, and even more preferably 30% or less, with the lower limit typically being 10% or more. In a conventional positron emission tomography device, if the energy absorptivity is not 100%, the scattered radiation generated by the scintillator becomes noise, resulting in a reduced S / N ratio. However, in some embodiments of the present invention, the scintillator is made of a material that has a certain degree of radiation detection efficiency, a short fluorescence lifetime, and negligible intrinsic background, making it possible to ignore the scattered radiation generated by the scintillator. Therefore, a high S / N ratio can be ensured even when the energy absorptivity is low as described above.

[0030] The luminescence output of the scintillator is usually 1000 Ph / MeV or more, preferably 5000 Ph / MeV or more, and more preferably 20000 Ph / MeV or more. There is no particular upper limit, and the higher the upper limit, the more improved the radiation detection sensitivity.

[0031] A high density and a large effective atomic number of the scintillator increase the radiation energy absorption rate and improve the radiation detection efficiency. From this perspective, the density of the scintillator is usually 4 g / cm 3 More than 6g / cm 3 More preferably, 7 g / cm 3 More preferably, 7.5 g / cm 3 More particularly, particularly preferably 8 g / cm 3 There is no particular upper limit, but it is usually 12 g / cm 3 The following is the result. Furthermore, the effective atomic number of the scintillator is usually 30 or more, preferably 40 or more, more preferably 45 or more, even more preferably 50 or more, particularly preferably 55 or more, and especially preferably 60 or more, and there is no particular upper limit, and it can be, for example, 100 or less. The effective atomic number can be determined based on the composition of the scintillator, with reference to the description in Medical Physics, 39 (2012), p. 1769.

[0032] The type of scintillator is not particularly limited as long as it has the properties required of a scintillator described above and does not impair the essence of the present invention; for example, when the radiation is gamma rays, BGO, plastic scintillators, organic scintillators, lutetium orthosilicate (LSO) and yttrium- or gadolinium-substituted scintillators of the LSO (LYSO, LGSO), hafnium oxide-based scintillators (BaHfO3, SrHfO3, CaHfO3, etc.), LuBr3, Nd-doped LaF3, or Yb-doped garnet (YAG: Yb, YbAG) can be used. Among these, in terms of the low intensity of the inherent background, hafnium oxide scintillators, plastic scintillators, BGO, Nd-doped LaF3, or Yb-doped garnet (YAG: Yb, YAG) are preferred, as they all have a background intensity of 200 Hz / cm 3 Among them, from the viewpoint of short DT, Nd-doped LaF3 (DT = 20 ns or less), Yb-doped garnet (DT = 50 ns or less), and hafnium oxide scintillators (DT = 10 to 20 ns) are particularly preferred. Note that Yb-doped garnet has a density of 4.56 g / cm 3 , effective atomic number 32.6, hafnium oxide scintillator density 8.1 g / cm 3 , and effective atomic number 64, which is preferable from the viewpoints of density and effective atomic number. Only one type of these scintillators may be used, or two or more types may be used in any combination. Also, part of the composition of the scintillator may be replaced with another element. The scintillator can be produced by a known method, or a commercially available product may be used.

[0033] <Conversion output section> The conversion output unit (hereinafter, sometimes simply referred to as the "conversion output unit") in the positron emission tomography apparatus according to this embodiment is not particularly limited as long as it can receive the electromagnetic waves emitted by the scintillator and output an electrical signal corresponding to the received waves, and any known conversion output unit can be used. Preferably, an electrical signal having a signal value corresponding to the energy of the electromagnetic waves is output, and more preferably, the signal value is proportional to the energy. In this case, the positron emission tomography apparatus can discriminate radiation energy information according to the signal value, and can distinguish between background radiation, scattered radiation, scattered coincidence counts, true coincidence counts, etc. The conversion output unit may also include a mechanism for amplifying the received signal value in order to improve sensitivity. The conversion output unit may include a light receiving unit (which may be a photoreceiver) that receives electromagnetic waves emitted from the scintillator, and a signal output unit that outputs the electromagnetic waves received by the light receiving unit as an electrical signal. The light receiving unit and signal output unit may be provided as separate components, or may be provided as an integrated component. The form of the light receiving unit is not particularly limited, and known components may be used, or commercially available products may be used. Furthermore, the form of the signal output unit is not particularly limited, and an appropriate combination of, for example, a circuit that accumulates electric charge and a circuit that outputs electric charge as an electrical signal according to a certain timing or amount of accumulated electric charge may be used, and known components may be used, or commercially available products may be used. The type of conversion output unit is not particularly limited, and known or commercially available units may be used. Furthermore, the time resolution of the conversion output unit is preferably such that it can distinguish signals with short time intervals equal to or greater than the DT of the scintillator. For example, a photomultiplier tube, a Si-avalanche photodiode, or a Si-Geiger mode avalanche photodiode may be used. Among these devices, a position-sensitive conversion output unit, such as a multi-anode type or an array type, is more preferable in order to prevent a reduction in the detectable area due to gaps between the conversion output units.

[0034] <Time window> In the coincidence method, the time window in the converter / output unit is represented by τ, measured in nanoseconds (ns). When the measurement start time is t=0, for t=τ, 2τ, 3τ, and so on, if two different detectors measure radiation within the time period (n-1)τ≦t≦nτ (n is a natural number), it is considered that one coincidence has occurred (one event has occurred). The time window can be, for example, 200 ns or less, preferably 180 ns or less, more preferably 160 ns or less, even more preferably 140 ns or less, and particularly preferably 120 ns or less. There is no particular lower limit, but it is usually 10 ns or more. Setting the time window to a short value reduces random coincidences, eliminates noise, improves the S / N ratio, and shortens the measurement cycle time.

[0035] <Transmission section> As described above, the positron emission tomography imaging apparatus according to this embodiment may include a transmission unit that transmits electrons or holes generated in the semiconductor material as an electrical signal to a signal processing system. The transmission unit, also referred to in the art as an electrical signal output unit, may be a known configuration, or a commercially available product may be used. For example, the transmission unit may include a charge storage circuit such as a capacitor. Furthermore, the transmission unit may include an amplifier circuit (such as an amplifier or an integrating amplifier circuit) that amplifies the charge or electrical signal information, a glitch removal circuit such as a sample-and-hold circuit, a ground that can be connected to a circuit that can store charge, a switch that switches the connection between the ground and the circuit on and off, a filter circuit (such as a low-pass filter or a high-pass filter) that removes unnecessary low-frequency and high-frequency noise, and the like. By appropriately including these circuits, sensitivity can be improved, noise can be removed, or residual charge can be removed after the electrical signal is output, thereby improving the accuracy of the electrical signal.

[0036] <Radiation detector and surrounding structure> The radiation detector section may be integrated with or arranged around another member capable of detecting or shielding radiation, and used in combination with the other member. The separate member is not particularly limited as long as it is a member or device capable of detecting or blocking radiation. For example, a radiation conversion member or a radiation shielding member can be used. The radiation conversion member may be, for example, a semiconductor that converts radiation into an electrical signal. Or, as with the scintillator of one embodiment, a general scintillation material that converts radiation into electromagnetic waves can be used. These members can be used, for example, as an active collimator. Examples of radiation shielding members include collimators that transmit light only in a specific direction using lead or tungsten, or coded collimators. However, other materials and shapes that shield radiation may also be used. The separate member may be positioned between the subject and the scintillator unit (before the scintillator) or on the opposite side of the subject from the scintillator unit (after the scintillator). Such a positioning allows the separate member to be used as a signal selection member, trigger, or the like.

[0037] <Signal Processing System> In one embodiment, the present invention is a signal processing system. Hereinafter, the signal processing system according to this embodiment may be simply referred to as a "signal processing system." The signal processing system may be used in a general positron emission tomography apparatus, or may be used as a component of the positron emission tomography apparatus according to one embodiment of the present invention described above. The signal processing system creates image data based on the electrical signals output from the radiation detector. That is, when two radiation detection elements detect signals above a threshold within a certain time period (time window), it starts collecting data on the signal values ​​of the electrical signals, and calculates the energy value from the signal values ​​of the electrical signals and the radiation source location (limited to one dimension) from the positions of the two radiation detection elements that received the radiation. In this process, time information on the detection may also be acquired. Alternatively, when a signal is detected by only one radiation detection element, time information may also be acquired and compared with the time information of another radiation detection element to determine whether the signals were detected simultaneously. Then, the position (point) of the radiation source is identified by accumulating statistics on the calculated radiation source location. Furthermore, for the position information identified by the above method, the likelihood of the detected position can be calculated based on the position resolution of the detector, etc. Finally, the final image can be reconstructed by reflecting the above positions and their likelihoods. A known method can be used to perform the reconstruction, such as the Ordered Subset Expectation Maximization Method (OSEM). Furthermore, in the signal processing system, for electrical signals derived from radiation having an energy value below the inherent energy value of the radiation, for example, 511 keV or less in the case of gamma rays from positron-emitting nuclides, image reconstruction may be performed by weighting the electrical signals in terms of the likelihood of position based on the energy information. As a method for evaluating the likelihood of a position based on the energy information, for example, in the case of scattered coincidence counting, the degree of angle change due to Compton scattering can be predicted based on the energy information, and the region where the radiation generation position may exist can be predicted based on this prediction. The form of the image data created by the signal processing system is not particularly limited, and for example, for each event detected as simultaneous measurement, the position of generation of the positron that is the origin of the radiation can be estimated, and an image can be created in which the concentration distribution of the positron source (tracer) at each position is plotted. By using the above method, the S / N ratio can be further improved.

[0038] In medical positron emission tomography, the tendency of sugars such as glucose to localize in cancer cells is utilized to detect positron-emitting nuclides, such as radioactive isotopes of fluorine. 18 By using fluorodeoxyglucose substituted with F as a tracer, the location of radiation emission can be determined as the location where cancer cells may exist. Positron-emitting nuclides are radioactive isotopes of oxygen, carbon, and nitrogen, respectively. 15 O. 11 C, or 13Examples of tracers to which a positron-emitting nuclide is attached include, but are not limited to, glucose, water, acetic acid, and the like. The tracer is injected or inhaled into the subject or patient. The amount of radioactivity injected is usually between 1MBq and 1GBq, but because lowering the dose reduces the impact on the patient, it is necessary to be able to obtain images reliably with a low dose.

[0039] In one embodiment, the signal processing system is a signal processing system that creates image data based on electrical signals output from a radiation detector, and the signal processing system recognizes the group of electrical signals as a processing target; a signal processing system, wherein the set of electrical signals includes at least a portion of the set of electrical signals that meets the following requirements: A group of electrical signals corresponding to gamma rays having an energy of 375 keV or less, whose signal values ​​are within a predetermined range, the predetermined range is 50% or more and 80% or less of a signal value of 100%, The 100% signal value is the signal value detected when a gamma ray having an energy of 511 keV is incident on a radiation detection element in a radiation detector and is totally absorbed by the radiation detection element. A signal processing system that recognizes such a group of electrical signals as the processing target can improve detection efficiency compared to conventional systems, and further, the improved detection efficiency can reduce the total radiation dose, which in turn reduces random coincidence counts and improves the S / N value.

[0040] The predetermined range of the signal value is usually 30% or more, preferably 40% or more, more preferably 50% or more, even more preferably 52.5% or more, and even more preferably 55% or more of the 100% signal value, and is usually 80% or less, preferably 70% or less, more preferably 65% ​​or less, and even more preferably 60% or less. When the predetermined range is equal to or greater than the lower limit, electrical signals generated by low-energy radiation such as backscattered radiation caused by radiation scattering on the outer wall of the device can be appropriately excluded as noise, and when the predetermined range is equal to or less than the upper limit, more electrical signals can be used, including electrical signals derived from scattered coincidence counting.

[0041] From another viewpoint, in one embodiment, a signal processing system is a signal processing system that creates image data based on a group of electrical signals output from a radiation detector, the signal processing system recognizing the group of electrical signals as a processing target, The group of electrical signals includes at least a portion of a group of electrical signals corresponding to gamma rays having an energy value within a predetermined range, the predetermined range being 232 keV to 340 keV inclusive. A signal processing system that recognizes such a group of electrical signals as a processing target can improve detection efficiency compared to conventional systems, and further reduce the total radiation dose due to the improved detection efficiency, thereby reducing random coincidence counts and improving the S / N value. In this specification, an electrical signal "corresponding" to a gamma ray having a certain energy means an electrical signal that is generated when a gamma ray of the certain energy value is incident on and absorbed by a radiation detection element.

[0042] The predetermined range for the energy of the gamma rays is not particularly limited, but is usually 180 keV or higher, preferably 200 keV or higher, more preferably 220 keV or higher, even more preferably 232 keV or higher, and still more preferably 250 keV or higher, and is usually 420 keV or lower, preferably 375 keV or lower, more preferably 340 keV or lower, even more preferably 320 keV or lower, still more preferably 300 keV or lower, and particularly preferably 280 keV or lower. When the predetermined range for the energy value of the gamma rays is equal to or higher than the lower limit, electrical signals generated by low-energy radiation such as backscattered radiation caused by radiation scattering on the outer wall of the device can be appropriately excluded as noise. When the predetermined range for the energy value of the gamma rays is equal to or lower than the upper limit, more electrical signals can be used, including electrical signals derived from scattered coincidence counting.

[0043] Here, the signal processing system "including" a certain electrical signal in its processing targets means that the electrical signal is regarded as a true coincidence count, and gamma ray generation position information, energy information, etc. estimated from the electrical signal are used for the purpose of creating image data. In one embodiment, it is sufficient that the processing targets use at least a portion of the electrical signals that satisfy the above-mentioned condition related to the "predetermined range," and all electrical signals that satisfy the above-mentioned condition may be processed.

[0044] In one embodiment, the signal processing system recognizes, as targets for processing, all electrical signals output from the radiation detector that have a signal value equal to or greater than a certain threshold. The threshold is typically 50% or less, preferably 40% or less, more preferably 30% or less, and even more preferably 20% or less, of the signal value (assuming that signal value is 100%) obtained when a gamma ray having an energy of 511 keV is incident on a scintillator and is totally absorbed by the scintillator, and is typically 5% or more, preferably 10% or more, and more preferably 15% or more. The signal processing system may recognize, as targets for processing, electrical signals having a signal value equal to or less than the threshold within a range that achieves the effects of the present invention, for example, within a range that does not include noise generated by low-energy radiation such as backscattered radiation described below. When the threshold value is equal to or greater than the lower limit of the range, electrical signals generated by low-energy radiation such as backscattered radiation caused by radiation scattering on the outer wall of the device can be appropriately excluded as noise, and when the threshold value is equal to or less than the upper limit of the range, more electrical signals can be used, including electrical signals resulting from scattered coincidence counting. In conventional PET methods, only gamma rays of 511 keV are calculated as true coincidences, but when the positron emission tomography apparatus of this embodiment is used, events detected at 511 keV or less can also be calculated as true coincidences. In general, backscattered radiation is excluded from the processing target, so gamma rays of 232 keV or higher are the processing target.

[0045] <Other equipment> The positron emission tomography apparatus may be provided with apparatus units other than the above-mentioned apparatus units, such as a computed tomography apparatus or a cooling device.

[0046] There are no particular limitations on the method for manufacturing the positron emission tomography apparatus, and it can be manufactured according to a known method so that the above-mentioned device components are arranged in desired positions.

[0047] (Signal processing method) In one embodiment, the present invention provides a signal processing method for generating image data based on a group of electrical signals output from a radiation detector, the method comprising: The signal processing method includes recognizing the group of electrical signals as a processing target; The set of electrical signals includes at least a portion of the set of electrical signals that meets the following requirements: A group of electrical signals corresponding to gamma rays having an energy of 375 keV or less, whose signal values ​​are within a predetermined range, the predetermined range is 50% or more and 80% or less of a signal value of 100%, In this signal processing method, the 100% signal value is the signal value detected when a gamma ray having an energy of 511 keV is incident on a radiation detection element in a radiation detector and is totally absorbed by the radiation detection element. A signal processing method that recognizes such a group of electrical signals as the processing target can improve detection efficiency compared to conventional methods, and further reduce the total radiation dose due to the improved detection efficiency, which in turn reduces random coincidence counts and improves the S / N value.

[0048] In one embodiment, the present invention provides a signal processing method for use in radiography, which generates image data based on a group of electrical signals output from a radiation detector, the method comprising: the signal processing method recognizes the group of electrical signals as a processing target; In this signal processing method, the group of electrical signals includes at least a portion of a group of electrical signals corresponding to gamma rays having energy values ​​within a predetermined range, the predetermined range being 232 keV to 340 keV. A signal processing method that recognizes such a group of electrical signals as the processing target can improve detection efficiency compared to conventional methods, and further reduce the total radiation dose due to the improved detection efficiency, thereby reducing random coincidence counts and improving the S / N value.

[0049] The signal processing method can be realized, for example, by using the signal processing system. All conditions and features of the signal processing method, such as the equipment, devices, processing conditions, preferred ranges of signal values ​​or gamma ray energy values, and obtained effects, are not limited as long as the object of the invention can be achieved. For example, in the description of the signal processing system above, "signal processing system" may be read as "signal processing method," and expressions such as "comprise" and "include" may be read as "use," as appropriate. Moreover, the signal processing method can be used in the signal processing step (c) in the following positron emission tomography method.

[0050] (positron emission tomography) In one embodiment, the present invention comprises: A positron emission tomography method comprising at least the following steps (a), (b) and (c): (a) A scintillation step in which radiation is converted into electromagnetic waves using a scintillator that receives radiation and emits electromagnetic waves. (b) a conversion / output step of receiving the electromagnetic waves emitted from the scintillator, converting the received electromagnetic waves into a pulsed electric signal, and outputting the signal; (c) a signal processing step including a step of performing signal processing by the signal processing method.

[0051] The steps (a), (b), and (c) can be realized, for example, by using the positron emission tomography apparatus. All conditions and features of the steps (a), (b), and (c), such as the composition, structure, and properties of the devices or components used, the conversion output conditions, the signal processing conditions, and the effects obtained by the conditions, are not particularly limited as long as the object of the invention can be achieved. For example, in the description of the positron emission tomography apparatus, the terms "positron emission tomography apparatus," "scintillator unit," "conversion output unit," and "signal processing system" may be read as "positron emission tomography method," "scintillation step," "conversion output step," and "signal processing step," respectively, and "equipped" and "including" may be read as "utilizing," as appropriate. Furthermore, the positron emission tomography method may optionally include a step other than the image conversion step.

[0052] In addition, the positron emission tomography method may include a step of injecting a substance (tracer) into the patient to be imaged, in which some elements of molecules localized in cancer cells have been replaced with positron-emitting nuclides (radiation sources), or, in the case of a positron emission tomography device having a columnar space with a scintillator unit arranged in the wall forming the column, a step of guiding the patient into the space. The types of tracers described above in relation to the positron emission tomography apparatus can be similarly applied.

[0053] <Simulation conditions> The effects of the present invention can be verified by simulation. In the simulation, a positron imaging device and a subject are created in a virtual space, and a predetermined number of tracers that emit a predetermined type of radiation are created in a specific region of the subject. This makes it possible to simulate how the tracers emit a number of radiations corresponding to the number of tracers and how these radiations enter the scintillator in the positron imaging device. For example, a positron emission tomography device can be created in a virtual space using program code that calculates the interaction between matter and particles or photons using the Monte Carlo method. Here, the radiation can be set to alpha rays, beta rays, gamma rays, X-rays, etc., and by setting the initial energy of the radiation, it is possible to predict differences in results based on energy information. Only one type of radiation may be used, or two or more types may be used.

[0054] <Calculation program> The method for calculating the interaction between a substance and particles or photons can be any existing method, and is not particularly limited, but for example, the Monte Carlo method can be used. This makes it possible to calculate how the energy, angle, etc. of radiation emitted from the tracer changes as a result of scattering, diffraction, absorption, etc. in the subject, the space between the subject and the scintillator, the inside of the scintillator, the outer wall, etc. In addition, by setting the molar ratio, density, shape, or area of ​​the elements of each substance that makes up the subject, space, scintillator section, or outer wall, etc., it is possible to predict the behavior of each of the above substances when radiation is actually incident on them, and to predict results depending on the type, arrangement, thickness, or shape of the scintillator.

[0055] <Radiation detection unit and background radiation> For simplicity, the radiation detection section is provided with only a scintillator section, for example, a scintillator block (scintillator unit) with a specified size, and calculations can be performed to determine how, when radiation is incident on the scintillator unit, an interaction occurs between the radiation and the scintillator section, and how the radiation is detected with a predetermined probability. When multiple scintillator units are provided, the scintillator units are preferably spaced apart by 0.5 mm or more. Regarding background radiation, by setting the inherent background intensity of the scintillator and the background generated from each part, it is possible to predict the results depending on the background radiation and the processing method. The energy distribution of background radiation is arbitrary, but when a known material is assumed, literature values ​​can be applied. Furthermore, it is possible to set the energy resolution and the range of signal values ​​to be processed in the signal processing system, and it is also possible to predict the effects of changing these settings. Intrinsic background radiation refers to gamma rays, beta rays, etc. generated within the scintillator itself, which are not caused by the input of energy such as radiation from outside the scintillator, or gamma rays, etc. generated by pair annihilation within the scintillator of positrons emitted inside the scintillator. The intensity of these rays varies significantly depending on the type of scintillator, so values ​​can be set for each scintillator and reflected in the simulation. In addition to the inherent background, background radiation also includes continuous components that do not have a peak at a specific energy, which are generated by the Compton components of the beta rays and gamma rays. The occurrence probability (frequency) of these continuous components is usually 0.5 Hz / cm 3 Since this is very small, it can be ignored in this simulation. In addition, components derived from the environmental background generated by the environment outside the device, such as cosmic rays, may also exist as background, but the occurrence probability (frequency) of these components is usually 0.1 Hz / cm 3 Since this is very small, it can be ignored in this simulation.

[0056] <Detection method, detection efficiency, and S / N calculation> The total number of radiation beams detected by the scintillator unit is separated into radiation beams corresponding to electrical signals to be recognized as processing targets in a signal processing system, i.e., signals, and other radiation beams, i.e., noise, according to the energy information of each radiation beam. The number of signals (S) is divided by the number of noises (N), and the S / N value can be calculated as the value. The criteria for separating the number of signals and the number of noises can be appropriately determined depending on the detection method. For conventional detection methods, the criteria described in Comparative Example 1 in the Examples below can be adopted, while for the detection method of this embodiment, the criteria described in Example 1 in the Examples below can be adopted. Note that, for convenience of calculation in the simulation, it is also possible to use a method in which radiation is generated in only one direction, and the detection efficiency is determined by a single measurement technique, and then corrected to the detection efficiency and S / N value obtained by a simultaneous measurement technique.

[0057] <Energy window and energy resolution> In the simulation, in order to remove events detected in the energy band of backscattering, etc., only events in a specific energy band (hereinafter sometimes referred to as the energy window) are used as the signal (S) or noise (N) for single measurements, and the detection efficiency and S / N value can be calculated based on this. Furthermore, events with energies outside the energy window can be excluded from the calculation of the S / N value. The energy window can be set, for example, centered on the initial energy value of the radiation to be detected (511 keV in the case of a PET device), to a range of several times the energy resolution (FWHM value), for example, 2 to 6 times, preferably 3 to 5 times, and more preferably 4 times. The energy resolution (FWHM value) is arbitrary, but can be set, for example, to about 5 to 20% of the energy (e.g., 511 keV) that the radiation to be detected has at the time of generation, and can also be set to, for example, 10% or 15%. [Example]

[0058] Hereinafter, the present invention will be described in more detail with reference to examples, but the present invention is not limited to these examples. First, the items of a simulation in which the same settings were used in the examples and comparative examples will be described.

[0059] <Program and coordinate design> A positron emission tomography (PET) imaging system with the following conditions was created in a virtual space using GEANT4, a program code that calculates interactions between matter and particles or photons using the Monte Carlo method. The position in the virtual space was a three-dimensional space consisting of three mutually perpendicular axes, x, y, and z, with the center of the subject set at the origin, i.e., (x, y, z) = (0, 0, 0). The direction in which the patient is inserted into the PET imaging system was defined as the z axis, and the PET imaging system was positioned concentrically around the z axis. This PET imaging system has a signal processing system capable of implementing the signal processing methods described in the following examples and comparative examples.

[0060] <Simulation conditions> Figures 1 and 2 show the arrangement of the subject 1, cancerous area 3, and scintillator section (here, scintillator unit, or scintillator) 2 in the simulation space. The subject is assumed to be a biologically equivalent substance (brain), and the contents are, in molar ratio, 64.44% hydrogen, 7.33% carbon, 0.95% nitrogen, 27.01% oxygen, 0.05% sodium, 0.08% phosphorus, 0.04% sulfur, 0.05% chlorine, and 0.05% potassium. As mentioned above, the subject is positioned at the origin, and is a cylinder with a radius of 8 cm and a height of 10 cm (5 cm in both the z and -z directions), with the height along the z axis, and a density of 1.03 g / cm. 3 It was decided. Next, we assumed that the cancerous area was located at the origin. The size of the cancerous area was 1cm, 1cm, and 1cm in the x, y, and z directions, respectively. In other words, the cancerous area was located within the area bounded by (x, y, z) = (-0.5cm, -0.5cm, -0.5cm) and (x, y, z) = (0.5cm, 0.5cm, 0.5cm). We assumed that the tracer was distributed evenly throughout the cancerous area, and that 511 keV gamma rays were emitted from the cancerous area with a uniform probability. The gamma rays were set to be generated at a rate of 100,000,000 per second, which corresponds to a tracer radioactivity of approximately 100 MBq.

[0061] <Scintillator placement and settings> For simplicity in the simulation, a scintillator unit was placed in place of the radiation detector section, and it was assumed that radiation was incident on this unit and interacted with the scintillator, resulting in the detection of radiation and energy information. Schematic diagrams of the subject and scintillator arrangements as viewed from the x-axis and y-axis directions are shown in Figures 1 and 2, respectively. First, 32 scintillator units were arranged symmetrically on a cylindrical surface with a radius of 30.4 cm centered on the z-axis on the x-y plane, with equal intervals between units at the inner circle of 0.5 mm. Next, the same set of 32 units was placed at positions shifted along the z-axis by ±60.5 mm, ±121 mm, ±181.5 mm, and ±242 mm, respectively. Furthermore, it was stated that the space was filled with air except for the positron emission tomography machine and the subject. The size of each scintillator unit was 60 mm in length, 60 mm in width, and L mm in thickness, and calculations were performed for L = 1, 2, 5, 10, or 20 mm. The scintillator material was also changed as appropriate in Comparative Examples 1 and 2 and Example 1 described below, and the results were compared.

[0062] <Direction of radiation emission and inherent background radiation> The calculation was performed assuming that all gamma rays emitted from the cancerous area were emitted in the x-axis direction. In this example, the radiation source, subject, and scintillator were all arranged in circular symmetry around the z-axis, so even if the direction of the emitted gamma rays was limited to one direction as described above, the results obtained were equivalent to those obtained when gamma rays were actually emitted in random directions, including the efficiency of radiation detection. The inherent background radiation intensity was set for each scintillator material and was assumed to be generated from the scintillator unit or all scintillator units from which gamma rays were to be detected.

[0063] <Calculation method of radiation detection efficiency and S / N value for simultaneous measurement> Based on the above conditions, we simulated how the emitted gamma rays and background radiation are absorbed, scattered, or transmitted within the subject, in the air, and by the scintillator. We then evaluated the probability of the radiation being detected by the scintillator for each energy value, and finally determined the number of gamma rays detected as a signal. Note that not all incident radiation is necessarily detected, and some is detected and some is not as a result of the interaction between the radiation and the scintillator. This interaction is calculated using the Monte Carlo method based on the program code GEANT4, with the radiation energy, the scintillator density, and the effective atomic number as parameters, and the detection efficiency for each radiation energy value is derived as a result. Next, the detected radiation is classified based on the energy value information. 1. The radiation to be detected as a signal, i.e., the signal (S) related to the single measurement; 2. Noise (N) associated with single counting that is detected because it has energy within the energy window but should not be considered a signal and should be used as the N value when calculating the S / N value; and 3. Noise that corresponds to the energy band of background radiation, etc., and is not counted in the N value when calculating the S / N value. The detection efficiency was calculated based on the S value, and the S / N value was calculated based on the S value and N value. Specifically, when 100,000 gamma rays with an energy of 511 keV were irradiated from the center of the cancerous area, the number of gamma rays detected in the scintillator unit (hereinafter sometimes referred to as the "scintillator unit that should detect gamma rays") placed in the direction of gamma ray emission (x-axis direction from the center) was measured, i.e., the S and N values ​​for single measurement. Then, the "S / 100,000]^ 2 was calculated as the detection efficiency for simultaneous measurement. 2 The S / N value was calculated by dividing the number of signals related to simultaneous measurement by the number of noises, where / 100,000 is the number of signals related to simultaneous measurement and N is the number of noises related to simultaneous measurement. With regard to N, if one scintillator unit detects gamma rays as noise, it is considered that a signal or noise is actually detected in the paired scintillator unit, and N for single measurement and simultaneous measurement are considered to have the same value. Note that conditions such as an energy window for distinguishing between signal and noise were set individually in Comparative Example 1 and Example 1 described below, and in some cases, gamma rays detected in all scintillator units other than the scintillator unit that should detect gamma rays were also considered as noise.

[0064] <Comparative Example 1> Simulations were performed using a similar material as the scintillator, lutetium orthosilicate (LSO), with an effective atomic number of 64 and a density of 7.4 g / cm. 3 The effective atomic number of LSO was calculated to be 64 based on Medical Physics, 39 (2012), p. 1769.

[0065] <Background radiation settings> The inherent background radiation intensity of the scintillator is 300 Hz / cm across the entire energy band, based on the literature value for the inherent background intensity of LSO (arXiv preprint arXiv:1501.05372, 2015 - arxiv.org). 3 It was assumed that LSOs emit inherent background radiation derived from radioisotopes, the main components of which are gamma rays with energies of 88, 202, and 307 keV, and beta rays with an energy maximum of 596 keV. Because these are generated simultaneously in the series of decays of the radioisotopes, gamma rays with an energy of 509 keV, the sum of 202 and 307 keV, are also apparently detected. The intrinsic background also includes continuous components that do not have peaks at specific energies due to the Compton components of the beta rays and gamma rays, but these components are not considered in this simulation because their occurrence probability (frequency) is very low. Note that although the intrinsic background actually originates from all scintillator units, for simplicity in Comparative Example 1, it was assumed that the background radiation originates only from the scintillator unit that is to detect gamma rays. In addition, components derived from the environmental background generated by the environment within the device and components generated when two radiation beams generated in different locations accidentally hit different scintillators at the same time may also exist as background, but these components are not considered in this simulation because their occurrence probability (frequency) is also very low.

[0066] <Calculation of detection efficiency and S / N value> First, calculations were performed using a single measurement method under conditions simulating the conventional signal processing method. That is, the energy resolution (FWHM value) was set to 51 keV, and the energy window was set to 511±102 keV. Gamma rays detected by the scintillator unit that should detect gamma rays and that had an energy value within the energy window were counted as the number of events used to calculate the detection efficiency and S / N value. Next, among the events, events related to gamma rays with an energy of 511 keV were considered as signals (S) related to single measurements, and all events related to gamma rays with energy values ​​other than 511 keV within the energy window were considered as noise (N) to be taken into account in calculating the S / N value. Finally, [S / 100,000]^ 2 is the detection efficiency for simultaneous measurement, and (S^ 2 / 100,000) / N was calculated as the S / N value for simultaneous measurement. In principle, conventional PET devices in reality can only use gamma rays with an energy of 511 keV as a signal, and gamma rays with other energy values ​​should be treated as noise. However, they cannot distinguish between gamma rays with an energy of 511 keV and gamma rays with other energies that lie within an energy band (energy window) with an energy value centered on 511 keV and a width several times, for example, four times, the energy resolution (FWHM value, for example, 50 keV, which corresponds to approximately 10% of 511 keV). Therefore, in actual conventional PET devices, the number of gamma rays detected as having an apparent energy of 511 keV, i.e., the apparent number of events, is likely to actually include noise within the energy window. In this simulation, since energy values ​​can be discriminated in 1 keV increments, all gamma rays with energy values ​​other than 511 keV within the energy window, which should normally be treated as noise, were treated as noise. The conditions are shown in Table 1, and the results are shown in Tables 2 and 3.

[0067] <Comparative Example 2> The scintillator material has an effective atomic number of 64, equivalent to LSO, and a density of 8.1 g / cm 3 The detection efficiency and S / N value for simultaneous measurement were calculated in the same manner as in Comparative Example 1, except that the inherent background in the material was assumed to be negligible. The conditions are shown in Table 1, and the results are shown in Tables 2 and 3.

[0068] Example 1 The detection efficiency and S / N value for simultaneous measurement were calculated in the same manner as in Comparative Example 2, except for the following changes. That is, in calculating the S / N value, all electrical signal groups within the range of energy values ​​of 250 keV or more among the detected gamma rays were considered to be gamma rays that can be used as signals. Next, for gamma rays with an energy of 232 keV or more and less than 250 keV among the detected gamma rays, half of the number of events detected by the scintillator unit to be detected were counted as signals (S) in single measurement, and the other half were counted as noise (N) used to calculate the S / N value. Furthermore, for gamma rays detected by scintillator units other than those to be detected, all gamma rays with an energy of 232 keV or more were counted as noise (N) used to calculate the S / N value. As mentioned above, all electrical signals corresponding to gamma rays with an energy value of 250 keV or more are used as signals. This reflects the fact that, while conventional PET methods calculate true coincidence counts only for gamma rays of 511 keV, if the fluorescence lifetime of the scintillator material is short, the time window can be shortened, and the inherent background can be ignored, radiation with an energy value below 511 keV that belongs to an energy band excluding relatively low-energy background radiation such as backscatter can also be calculated as true coincidence counts. Furthermore, the handling of gamma rays with energies of 232 keV or more but less than 250 keV is based on the fact that in actual devices, due to the energy resolution, approximately half of these can be detected as gamma rays with energies of 250 keV or more. In addition, in Comparative Examples 1 and 2, which are conventional methods, if the range of N is made the same as the range of N in Example 1, that is, if the S / N value is determined in the same manner as in Comparative Examples 1 and 2, but the range of the energy window that determines the noise is expanded to the processing range in Example 1 (232 keV or more), both the detection efficiency and the S / N value will be very small, and therefore the method cannot be applied to a positron emission tomography apparatus.

[0069] The simulation conditions are shown in Table 1, and the simulation results are shown in Tables 2 and 3.

[0070] [Table 1]

[0071] [Table 2]

[0072] [Table 3]

[0073] In all of the above-mentioned Example 1 and Comparative Examples 1 and 2, the energy of the radiation that took the signal value when a gamma ray having an energy of 511 keV entered a scintillator (scintillator unit) and was totally absorbed by the scintillator was 511 keV. Therefore, in the simulations of Comparative Examples 1 and 2 set under the above conditions, only the signal value for which the ratio of the minimum signal value in the signal to be processed to the signal value when a gamma ray having an energy of 511 keV entered the scintillator and was totally absorbed by the scintillator was 100% was recognized as the signal to be processed. On the other hand, in the simulation of Example 1 set under the above conditions, in the range of 232 to 511 keV, the ratio of the minimum signal value in the signal to be processed to the signal value when a gamma ray having an energy of 511 keV entered the scintillator and was totally absorbed by the scintillator was 45%. Therefore, in the simulation of Example 1, all signal values ​​for which the ratio of the minimum signal value in the signal to be processed to the signal value when a gamma ray having an energy of 511 keV in the signal to be processed enters the scintillator and is completely absorbed by the scintillator is 45% or more are recognized as signal values ​​to be processed.

[0074] As shown in the above tables, the positron emission tomography apparatus of Example 1 had high detection efficiency for simultaneous measurement. Furthermore, Example 1 also showed a high S / N value, despite the fact that the conditions were stricter than those of Comparative Examples 1 and 2 in that noise in all scintillator units had to be included. In particular, as can be seen from a comparison between Comparative Example 2 and Example 1, when the thickness of the scintillator was the same, the S / N value was more than two or three times higher. Furthermore, as can be seen by comparing Comparative Example 1 with Example 1, by changing the detection method and by using a scintillator material with low intrinsic background intensity, the S / N value was further improved, and when the scintillator thickness was the same, the S / N value was more than 8 times higher than when the conventional method and conventional material were used, and depending on the thickness, it was nearly 20 times higher.

[0075] As described above, the present invention can provide a signal processing system, a positron emission tomography apparatus, and a positron emission tomography method that have improved detection efficiency and / or S / N ratio compared to conventional systems. [Explanation of symbols]

[0076] 1. Subject 2. Scintillator 3 Cancer area

Claims

1. A signal processing system that creates image data based on a group of electrical signals output from a radiation detector, the signal processing system recognizes the group of electrical signals as a processing target; a signal processing system, wherein the set of electrical signals includes at least a portion of the set of electrical signals that meets the following requirements: a group of electrical signals corresponding to gamma rays having an energy of 375 keV or less, the signal values ​​of which are within a predetermined range; the predetermined range is 50% or more and 80% or less of a signal value of 100%, The 100% signal value is a signal value detected when a gamma ray having an energy of 511 keV is incident on a radiation detection element in a radiation detector and is totally absorbed by the radiation detection element.

2. A signal processing system that creates image data based on a group of electrical signals output from a radiation detector, the signal processing system recognizes the group of electrical signals as a processing target; The group of electrical signals includes at least a portion of a group of electrical signals corresponding to gamma rays having energy values ​​within a predetermined range, the predetermined range being equal to or greater than 232 keV and equal to or less than 340 keV.

3. A positron emission tomography apparatus comprising the signal processing system according to claim 1 or 2 and a radiation detector section.

4. The positron emission tomography apparatus according to claim 3 , wherein the radiation detector section comprises the following elements: A scintillator section including a scintillator that receives radiation and emits electromagnetic waves, and a conversion output section that receives the electromagnetic waves emitted from the scintillator, converts the received electromagnetic waves into pulsed electrical signals, and outputs the signals.

5. 5. The positron emission tomography apparatus of claim 4, wherein the scintillator satisfies the following characteristics: The intensity of the inherent background of the scintillator is 200 Hz / cm in the range of signal value 10% to 120%, where the signal value of the pulsed electric signal when a gamma ray having an energy of 511 keV is incident on the scintillator and is totally absorbed by the scintillator is 100%. ３ The following is the result.

6. 6. The positron emission tomography apparatus according to claim 4, wherein the time window at the conversion output section is 180 ns or less.

7. 7. The positron emission tomography apparatus according to claim 4, wherein the scintillator has a fluorescence decay time (DT) of 25 ns or less when irradiated with gamma rays.

8. 8. The positron emission tomography apparatus according to claim 4, wherein the gamma ray absorption rate of the scintillator is 70% or more.

9. 8. The positron emission tomography apparatus according to claim 4, wherein the gamma ray absorption rate of the scintillator is 50% or less.

10. A signal processing method for generating image data based on a group of electrical signals output from a radiation detector, the signal processing method comprising: recognizing the group of electrical signals as a processing target; a signal processing method, wherein the set of electrical signals includes at least a portion of the set of electrical signals that meets the following requirements: a group of electrical signals corresponding to gamma rays having an energy of 375 keV or less, the signal values ​​of which are within a predetermined range; the predetermined range is 50% or more and 80% or less of a signal value of 100%, The 100% signal value is a signal value detected when a gamma ray having an energy of 511 keV is incident on a radiation detection element in a radiation detector and is totally absorbed by the radiation detection element.

Images