JP6921402B2 - Noise removal method and radiation source position estimation method - Google Patents

Description

The present invention relates to a noise removing method and a radiation source position estimation method. More specifically, it is preferably applied in Positron Emission Tomography (PET).

PET is a PET drug containing proton excess nuclei such as ¹¹ C and ¹⁸ F in the living ^{body, and two positrons generated by β +} decay are generated 180 degrees in the opposite direction when the positrons are paired with the electrons in the body. It is a nuclear medicine diagnostic method that obtains the distribution of PET drugs in the body by measuring 511 keVγ rays, and in principle, even minute early cancers can be diagnosed.

However, although some of the whole-body PET devices currently on the market have a position resolution of 1 to 3 mm according to the catalog value, the PET image actually obtained by a doctor is greatly blurred. The actual position resolution is about 10 mm for a good product and about 50 mm for most commercial products.

Further, regarding a known PET device, for example, there is a technique described in Patent Document 1 below.

Japanese Unexamined Patent Publication No. 2016-133333

By the way, the event that 511keVγ rays are scattered in Compton in the measuring instrument and emitted light in multiple places in the measuring instrument can be used for tomographic image reconstruction if the first γ-ray incident position is known. The vessel must be subdivided and the amount of light emitted in each small block must be measured accurately.

Typical block sizes for existing PET devices range from 10 mm to 50 mm, which is much larger than the required position resolution. Therefore, in the existing PET apparatus, events including Compton scattering cannot be used for tomographic image reconstruction and must be discarded.

It is considered that a position resolution of about 3 mm is sufficient for a PET device for systemic cancer diagnosis, but from the viewpoint of earlier treatment, and also from the viewpoint of the necessity of diagnosing abnormalities in the brain and determining the efficacy of animal experiments at the time of new drug development. There is a demand for a PET device having a high position resolution that can be indicated by a digit of 0.1 mm.

It is shown that the resolution of the reconstructed tomographic image is determined by the statistic in the case of PET having a position resolution of γ-ray measurement higher than 1 mm. ^{In a normal PET examination, the concentration of 18} F-FDG is set to about 3 MBq / L so that the subject's exposure dose is about 3 mSv. In this case, the event density of collapse during the inspection time is about ^{3000 events / mm 3.} Of these, 10% of the two γ-rays are both emitted to the outside without Compton scattering in the living body, 50% of which is incident on the measuring instrument, 80% reacts on the measuring instrument, and 4% is two. Since both are photoelectric absorption, the event density emitted from normal tissue and observed as photoelectric events is 5 events / mm ³ . In addition to this, there are about the same number of events that cause small-angle Compton scattering in the living body but are photoelectrically absorbed with energy of 420 keV or more in the measuring instrument, so the event density of normal tissue observed by the measuring instrument is 10 events / It is mm ^3. On the other hand, since glucose is accumulated in the cancer tissue at a concentration five times higher than that in the normal tissue, the event density of the cancer tissue is 25 events / mm ³ . This difference is 2.5 times the standard deviation, and a 1 mm cancer can be diagnosed at the last minute. In order to diagnose a cancer of 0.5 mm, it is necessary to set the exposure dose of the subject to 20 mSv or to actively utilize the Compton scattering event that must be discarded in the calculation of a normal PET reconstructed tomographic image.

Therefore, in view of the above problems, the present invention provides a noise removing method and a radiation source position estimation method that reduce data waste and achieve high position resolution.

The radiation source position estimation method according to one aspect of the present invention that solves the above problems is arranged on a flat scintillator, a first optical fiber layer arranged on one surface of the scintillator, and on the other surface of the scintillator. in the second and the optical fiber layer, the radiation source position estimation method to be applied in the radiation detecting apparatus in which a plurality stacked scintillator blocks cylindrical scintillator unit comprising a plurality of light receiving elements disposed on the side surface of the scintillator, the Therefore, when the emitted light is not completely reflected in the scintillator, only the emission position of the emitted light in the scintillator unit is recorded. When the emitted light is totally reflected in the scintillator, the emission position of the emitted light in the scintillator unit and The radiation source position is estimated based on the step of recording a plurality of light receiving time sets of the light receiving element, the step of removing the light emitting position and the light receiving time set in the second light emission, the light emitting position, and the light emitting position and the light receiving time set. It is provided with a step to do.

As described above, according to the present invention, it is possible to provide a noise removing method and a radiation source position estimation method that achieve high position resolution by reducing data waste.

It is a figure which shows the outline of the radiation detection apparatus which concerns on embodiment. It is a figure which shows the outline of the scintillator block of the radiation detection apparatus which concerns on embodiment. It is a figure which shows the outline of the scintillator unit of the radiation detection apparatus which concerns on embodiment. It is a figure which shows the image of the light emission in a scintillator unit. It is a figure which shows the image in the case of a plurality of light emission in a scintillator block.

Hereinafter, embodiments of the present invention will be described in detail with reference to the drawings. However, the present invention can be implemented in many different embodiments and is not limited to the examples described in the embodiments shown below.

(Device)
FIG. 1 shows an outline of a cross section of an apparatus using the radiation source position estimation method according to the present embodiment.

As shown in this figure, the radiation detection device (hereinafter referred to as “the device”) 1 according to the present embodiment receives the radiation γ emitted from the radiation source, estimates the radiation source C, and uses this radiation source. It is a device that can create a visualized image by accumulating a plurality of images. That is, the structure includes a plurality of scintillator blocks 2 and a control device 3 connected to the plurality of scintillator blocks 2. Specifically, the scintillator block 2 receives radiation and outputs it as a signal, and the control device 3 performs various signal processing to create an image. Specifically, as described above, this device is the most applicable device to be a PET device.

In this device, as shown in this figure, a plurality of scintillator blocks 2 are arranged in a tubular shape. More specifically, the measurement target person is laid on a bed or the like in the center, and the position of the radiation source of the radiation emitted from the measurement target person's body is specified. As described above, the measurement subject has been previously administered with a PET drug containing a proton excess nucleus such as ¹¹ C or ^{18 F, and radiation is emitted from this drug.}

Further, FIG. 2 is a diagram showing an outline of the scintillator unit 21 used in this device. As shown in this figure, in the scintillator block 2, a plurality of scintillator units 21 are stacked and arranged. Further, as shown in this figure, it is preferable that a plurality of scintillator units 21 are arranged even in the same layer. By doing so, the scintillator unit 21 can be subdivided and the resolution can be improved. Then, each of the scintillator units 21 outputs a signal to the control device 3 and is subjected to signal processing by the control device 3.

Further, FIG. 3 shows a more specific structure of the scintillator unit 21 in the present device 1. As shown in this figure, the scintillator unit 21 includes a flat plate scintillator 211, a first optical fiber layer 212 arranged on one surface of the scintillator 211, and a second scintillator unit 21 arranged on the other surface of the scintillator 211. The optical fiber layer 213 and a plurality of light receiving elements 214 arranged on the side surface of the scintillator 211 are provided, and as a result, the light emitting position and the light receiving time in the scintillator 211 can be specified.

Further, the scintillator 211 in the scintillator unit 21 has a flat plate shape and emits light when radiation (γ-ray) is incident. The material of the scintillator 211 is not limited as long as it has this function, and may be, for example, an organic scintillator in which a fluorescent substance is dispersed in a plastic such as polystyrene or polyvinyltoluene. It is preferable to have.

Here, the sintered scintillator is formed by sintering an inorganic solid substance, and when it is exposed to radiation, it can emit light by absorbing the energy of the radiation. .. The material of the sintered scintillator is not limited as long as it has the above functions, and for example, La-GPS ((Gd _0.75 La _0.24 Ce _0.01 ) ₂ Si ₂ O ₇ ) or the like can be used. ..

Further, in the above organic scintillator, the fluorescent substance is not limited, but is preferably an organic fluorescent substance or an inorganic fluorescent substance, for example. In the case of an organic fluorescent substance, for example, anthracene, stilbene and the like can be exemplified, but the present invention is not limited thereto. In the case of inorganic fluorescent substances, for example, gadolinium silicate, cesium iodide, bismuth germanium, lutetium silicate, thallium-activated sodium iodide, cerium fluoride, lead tungstate, barium fluoride, lead tungstate, and foot Examples include, but are not limited to, barium silicate, lead fluoride, and the like.

The size of the scintillator unit 21 can be appropriately adjusted according to the size of the device and the required accuracy, but is preferably a rectangle having a side of 20 mm or more and 100 mm or less. The thickness is not limited as long as the light receiving element can be arranged, but is preferably 1 mm or more and 10 mm or less, and more preferably 5 mm or less.

Further, the first optical fiber layer 212 is arranged on one surface of the scintillator 211 as described above, and can take out the light emitted by the scintillator 211 to the outside. As is clear from this figure, the optical fiber layer 212 has a plurality of optical fibers arranged in parallel. Then, when light is generated in the scintillator 211, the light is incident on the optical fiber in the vicinity of the scintillator 211 and is guided to the outside. As a result, it is possible to roughly confirm which region is shining in the uniaxial direction.

The material of the optical fiber layer 212 is not particularly limited, but is preferably a wavelength conversion fiber. Here, the "wavelength conversion fiber" is a fiber capable of absorbing incident light and emitting light having a wavelength different from the wavelength of the incident light. By performing wavelength conversion, light can be output as a signal to the outside more efficiently.

Further, in the present device 1, the optical fiber layer 212 may be arranged as a single optical fiber layer, or may be laminated in a plurality of layers. By doing so, it is possible to further improve the detection accuracy.

Further, in the present device 1, the second optical fiber layer 213 is substantially the same as the first optical fiber layer 212 except that the arrangement (axis) direction of the optical fibers is different. More specifically, the optical fiber of the first optical fiber layer 212 and the optical fiber of the second optical fiber layer 213 are preferably in the intersecting direction, and more preferably in the substantially vertical direction. By doing so, it is possible to obtain the vertical coordinate value on the vertical axis and the abscissa value on the horizontal axis, and it is possible to estimate that this intersection is a region close to the light emitting position. More detailed processing will be described later.

Further, in the present device 1, a plurality of light receiving elements 214 are arranged on the side surface of the flat plate scintillator 211. The light receiving element 214 is preferably one capable of extracting the light emitted by the scintillator 211, and for example, a photon measuring device such as MPPC (Multi-Pixel Photo Counter) may be used. preferable. Further, although the light receiving element 214 in the present device 1 is not limited, it is preferable that the light receiving element 214 is arranged on the entire side surface, or on all four side surfaces if it is rectangular. By doing so, the light emitted by the scintillator 211 can be detected without exception. Further, as will be described later, when light is detected by the light receiving element 214, the time is also recorded in the control device 3.

The control device 3 of the present device 1 is connected to the scintillator block 2, more specifically, the optical fiber and the light receiving element of each scintillator unit 21, and can perform data processing based on the output signal. be. The configuration of the control device 3 of the present device 1 is not limited as long as it has the above functions, but is not limited to a recording device such as a hard disk, a central computing device (CPU), a temporary recording device such as a memory, a keyboard, and the like. A so-called computer having an input device such as a mouse and a display device such as a liquid crystal display can be used. The specific flow of this data processing will be described later.

With the above configuration, the present device 1 can realize a radiation source position estimation method.
As described above, in the present device 1, the number of laminated scintillator units can be adjusted as appropriate. For example, the inorganic scintillator is 34 mm × 34 mm × 3 mm, 8 × 8 = 64 sheets, one layer, and eight layers, one block. It is a preferable example to use a whole body PET device with 6 blocks. Further, in PET for heads and small animals, it is preferable that one layer is 4 × 4 = 16 sheets or 5 × 5 = 25 sheets. Further, it is a preferable example that a wavelength conversion fiber sheet having a diameter of 0.2 mm is adhered to the upper and lower surfaces of the inorganic scintillator plate to measure the light emitting position.

(Radiation source position estimation method)
Next, a radiation source position estimation method (hereinafter referred to as “the estimation method”) applied to the present apparatus will be described. This estimation method includes a noise removal step (S1) and a radiation position step (S2). Each step will be described below.

As is clear from the above description, this estimation method is executed by the control device 3. More specifically, it can be realized by executing a program recorded on a recording medium such as a hard disk in the control device 3 as an information processing device by an operation of a user.

The noise removing step (S1) is also a noise removing method, and is a step of recording a plurality of sets of a light emitting position of light emission in the scintillator unit and a light receiving time of a light receiving element, and (S12) a plurality of the light emitting positions and the light receiving time. A step of determining that the light is emitted at the second location based on the set of the above, and (S13) a step of removing the set of the emission position and the light receiving time in the second emission.

First, (S11) The step of recording a plurality of sets of the light emitting position of the light emission in the scintillator unit and the light receiving time of the light receiving element is clear from the structure of the present device described above, but the position where the scintillator shines in the scintillator unit (light emitting position). ) Is estimated based on the light emitting positions of the first optical fiber layer and the second optical fiber layer. Specifically, as shown in FIG. 4, when radiation is incident on a scintillator and emits light, the light spreads from this light emitting position and is applied to a part of each of the first optical fiber layer and the second optical fiber layer. Light is incident and output to the outside. Then, by processing the output signal, it is possible to identify the shining optical fibers of the first optical fiber layer and the second optical fiber layer and estimate the light emitting position. In particular, when the first optical fiber layer and the second optical fiber layer are arranged so as to intersect, a region is specified by this intersection, and a predetermined calculation is performed based on the extent of each light emitting region in the scintillator. It is possible to estimate the light emitting position of.

Further, in this light emission, when the light is totally reflected in the scintillator, the light is repeatedly totally reflected, and finally the light is incident on the light receiving element and output as a signal to the outside. Then, the time when the light is incident on the light receiving element is also recorded, and is assembled and recorded together with the information on the light emitting position. That is, as a result, it becomes possible to confirm at what time and at what position the light was emitted. In this light emission, if the light emitted in the scintillator does not satisfy the total reflection condition, only the information on the light emission position is recorded.

In addition, this process sequentially records this set each time it emits light. By doing so, as will be described later, when two places of light are emitted by one radiation, it becomes possible to recognize which is the first and which is the second.

Next, the present estimation method includes (S12) a step of determining that the light is emitted at the second location based on the set of the plurality of light emitting positions and the light receiving time. Specifically, the set of the light emitting position and the light receiving time recorded sequentially in the above step is compared, and if the two sets satisfy a predetermined relationship, it is determined that one radiation emits light at two places, and the light emitted behind is determined. Can be determined to be the second emission. The predetermined relationship can be simply the difference in the light receiving time, and when the light is emitted within the predetermined time, it can be considered that the light is emitted at two nearby locations. That is, it is preferable to determine whether or not the light receiving time between the light emitting position and the light receiving time set is within a predetermined time. FIG. 5 shows an image diagram in the case of emitting light at these two locations. In the case of this figure, an example is shown in which two different scintillator units emit light.

Further, the present estimation method includes a step of removing the set of the light emitting position and the light receiving time in the second light emission (S13). In this way, if it is determined that the light is emitted at the second location, the second location becomes noise, so this is removed. Thereby, it is possible to provide a highly accurate radiation source position estimation method. When light is emitted at two locations in this way, the accuracy will decrease if both locations are used as data, while the accuracy cannot be improved if data is deleted at both locations. On the other hand, in this estimation method, even if it is grasped that the light is emitted at two places, it can be dealt with by deleting only one of the data. Methods and radiation source location estimation methods can be provided.

Further, the present estimation method includes a radiation position step (S2) for estimating the radiation source position based on the above-determined light emission position and light reception time. Specifically, radiation is emitted from the radiation source, but the PET apparatus emits two radiations in the opposite directions of 180 degrees. That is, the light emitting positions exist at approximately the same time in the scintillator blocks close to the positions facing each other across the tubular hollow portion. Therefore, the position of the radiation source can be estimated by connecting these two light emitting positions. As this method itself, a known method can be adopted.

As described above, this estimation method can provide a radiation source position estimation method that achieves high position resolution by reducing data waste. In addition, according to this estimation method, by using the majority of Compton scattering events for image reconstruction while maintaining the position resolution of the photoelectric absorption reaction of 0.2 mm (FWHM), the sensitivity is higher than that of a PET device that does not use this method. It is also possible to achieve 10 times.

In addition, since the amount of light emitted is measured independently by this device by blocking a large number of minute light received on the side surface, the time of incidence of γ-rays is also measured independently, and the time resolution is 1/5 that of the conventional device. Can be achieved. Generally, if there are multiple independent measurements, the error of the mean value is inversely proportional to the square root of the number of measurements. In the product of the present invention, a total of about 32 minute light receiving elements are adhered to the side surface, and the light emitting time is measured independently by them. Therefore, even if the same light receiving element as the conventional device is used, the time resolution is improved to 1/5 or less. do. One of the drawbacks of the current PET device is that the tomographic image calculation can be started only after the data collection is completed. Since the time required for the tomographic image calculation is longer than the data collection time, the diagnosis result can be known and the subject can be notified. It will be from the day after the inspection to the next week. If the time resolution is 100 psec or less, the tomographic image can be calculated in parallel with the data collection, and the tomographic image can be displayed immediately after the inspection is completed. The doctor can immediately determine the need for a work-up.

The present invention can be applied to a head office line detection device such as a PET device, and can be used in a noise removing method and a radiation source position estimation method applied thereto, and has industrial applicability.

Claims (4)

A flat scintillator, a first optical fiber layer arranged on one surface of the scintillator, a second optical fiber layer arranged on the other surface of the scintillator, and a plurality of surfaces arranged on the side surface of the scintillator. a radiation source position estimation method to be applied in the radiation detecting apparatus in which a plurality stacked scintillator blocks cylindrical scintillator unit and a light receiving element,
When the emitted light is not totally reflected in the scintillator, the step of recording only the emission position of the emitted light in the scintillator unit.
When the light emitted is totally reflected in the scintillator, a step of recording a plurality of sets of the emission position of the emission in the scintillator unit and the reception time of the light receiving element.
A step of removing the set of the light emitting position and the light receiving time in the second light emission,
A radiation source position estimation method comprising a step of estimating a radiation source position based on a set of the light emission position, the light emission position, and the light reception time . Before the step of removing the second emission,
The radiation source position estimation method according to claim 1, further comprising a step of determining that the light is emitted at a second location based on a set of the plurality of light emitting positions and the light receiving time. The step of determining that the light is emitted at the second location based on the set of the plurality of light emitting positions and the light receiving time is
The radiation source position estimation method according to claim 2, wherein it is determined whether or not the light receiving time is within a predetermined time between the light emitting position and the light receiving time set. The radiation source position according to any one of claims 1 to 3, wherein the average value of the light receiving times measured independently by the plurality of light receiving elements is the light receiving time of the light receiving element corresponding to the light emitting position of the light emission in the scintillator unit. Estimating method.

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