JP2012514734A - High resolution PET breast imager with improved detection efficiency - Google Patents

Abstract

A highly efficient PET breast imager for whole breast lesion detection, including those located near the patient's chest wall. The breast imager comprises a ring-shaped imaging module surrounding the imaging breast. Each imaging module includes a tilted imaging light guide inserted between the gamma sensor and the photoelectric detector. The tilted light guide allows the gamma radiation sensor to be placed in close proximity to the skin of the chest wall, thereby extending the sensitive area of the imager to the base of the breast. Several types of photoelectric detectors have been proposed for use in detector modules, and a preferred choice is a compact silicon photomultiplier for high compactness. The geometry of the detector head, the arrangement of the detector ring, significantly reduces the dead area and improves the detection efficiency of lesions located near the chest wall.
[Selection] Figure 5

Description

  The present invention relates to chest images, particularly high resolution PET chest imaging systems for lesions located near a patient's chest wall with improved detection efficiency.

  Breast cancer biology appears to indicate that breast cancer metastasis begins with a lesion size of about 2 mm. X-ray mammography, on the order of 50 micrometers, provides much better intrinsic spatial resolution, but the type of suspicious structure seen in mammograms for this anatomical imaging therapy (cancer or benign, Etc.) are less specific. In some patients, this x-ray imaging is completely useless for dense breast tissue, implants, or scar tissue as a result of previously performed surgery. Breast MRI can provide more structural information than mammography, but it also provides only non-specific information regarding the type of tissue imaged. Imaging techniques that provide information about the type or ecology of anatomical structures present in the breast, especially any suspicious lesions, are based on biomarkers that are sensitive to the molecular species present in the tissue. Examples of these functional or molecular imagers are the Positron Emission Tomography (PET) scanner and gamma camera. These standard imaging devices, as described above, are not capable of achieving the desired spatial resolution in the range of 2 mm or more in the breast. Therefore, it is important to develop functional imaging techniques that can provide 1 mm resolution scale imaging power.

  A breast-specific functional imager has been proposed that includes a dedicated compact gamma camera placed near the breast, but the distance between the collimator surface from the parallel hole mechanical collimator and the imaged features There is a physical limit in resolution due to the effects of The radioactive tracer used for single gamma breast imaging using a standard gamma camera and dedicated breast-specific instrument is typically Tc-99m-Sestamibi, which emits gamma rays at 140 keV and is labeled with Tc99m. Usually, the best average linear spatial resolution obtained with compressed breasts by these instruments using parallel collimators is on the order of 5-6 mm. A partial remedy to this problem is to simultaneously image the compressed breast from both sides, thus changing the average distance from any feature in the breast volume to the nearest collimator. However, even in this case, the 2 mm limit is not actually available due to the correlation between collimator resolution and sensitivity. Parallel collimators that provide 2 mm spatial resolution at 2.5 cm distance are available, but their sensitivity is too low and the resulting image shows too much statistical noise, so efficient detection of small cancer lesions is not possible Can not. Although other imaging methods having a pinhole collimator have been proposed to obtain better spatial resolution, there is a problem of low sensitivity due to the function of the pinhole gamma collimator.

  An FDA approved dedicated breast PET imager is available from Navisscan PETSystems, San Diego, California. These instruments can achieve a spatial resolution of <2 mm FWHM (full width at half maximum) with an F-18 positron emission radiation identification label used, for example, on the F-18-fluorodeoxyglucose (FDG) biomarker. Is possible. However, they involve another problem associated with the physics process of coincidence PET coincidence imaging, which results in low sensitivity in the part of the breast near the chest wall. PET detection efficiency is rapid at the detector end due to the geometric requirement of simultaneous detection of two consecutive 511 keV annihilation gamma rays produced by the action of positron annihilation by one of the electrons in the surrounding breast tissue. To drop. For this reason, when approaching the front end surface of the system, a dead edge, that is, a sharp decrease in detection efficiency occurs. In comparison, single photon imaging does not suffer from the chest wall effect to the same extent and, for example, a slant collimator solution can be used that further reduces the dead zone effect. This geometric effect has been defined in a recent clinical trial study as the cause of an undetectable lesion located near the base of the breast, near the wall chest. This is therefore a serious limitation and a recognized defect of dedicated PET mammography currently implemented in the planar scan module.

  Another example of a prior art chest imaging PET system is RR Rayman et al., Development of a Dedicated Positron Emission Tomography System for the Detection of Biotechnology, Non-patent Document 2 1). The imaging system includes two sets of planar PET imager modules mounted on a gantry that rotates under the patient's bed. Therefore, the two pairs of modules of this proposed system have only limited reach towards the patient's chest wall and cannot image the breast part closest to the chest wall.

Development of a Dedicated Positron Emission Tomography System for the Detection and Biopsy of Breast Cancer, NIMA, 2006, 564 (2): 291-295

  A method and apparatus is proposed to ameliorate the problem with a dedicated PET breast imager, a missing detection lesion located in a portion of the breast near the patient's chest wall.

  A highly efficient breast imager comprises a ring-shaped detection module surrounding the breast to be imaged. In each imaging module, a tilt imaging light guide is inserted between the gamma ray scintillator sensor and the photoelectric detector. The tilted light guide element allows the gamma sensor to be placed in very close proximity to the skin of the chest wall, thereby extending the sensitive area of the imager to the base of the breast. Several types of photoelectric detectors are proposed for use in the detector module, together with a new and compact silicon photomultiplier tube, due to the preferred choice and high compactness.

(Purpose and advantage)
Several advantages are achieved with the breast imager of the present invention including:
(1) A breast imager is made up of a plurality of imager modules, each module having a tilted optical fiber optic light guide inserted between a scintillator sensor and a photoelectric detector, with the scintillation sensor very close to the chest wall. It can be placed almost touching the patient's skin.
(2) The breast imager is placed in the form of a ring, and the ring is implanted in a properly formed patient bed, allowing the end of the detection module to be positioned directly on the chest wall.
(3) Geometrical arrangement of detector heads, arrangement of detector rings significantly reduces the dead area between individual detector heads compared to conventional PET systems and hence allows static imaging .
(4) The detector ring is attached to a disc attached to a scanning system that allows a limited angle of rotation about the chest and a vertical scan towards or away from the chest wall. And enables imaging of the entire breast.

  These and other objects and advantages of the present invention will be better understood by reading the following description with reference to the drawings.

  FIG. 1 is an end view of one variation of the imager system of the present invention that includes a set of detector modules that look at the laterally recumbent breast, where the detector head can efficiently detect lesions in the breast near the chest wall. Is lifted beyond the edge of the chest wall.

  FIG. 2 is a side view of a preferred imaging module according to the present invention comprising a scintillator array coupled with a Position Sensitive PMT with optical light guides made of plastic fibers arranged in an array.

  FIG. 3 is a cross-sectional view through the imager ring and is part of the patient's bed illustrating the thrust of the imager module relative to the chest wall.

  FIG. 4 is a cross-sectional view of an imaging system according to the present invention, in which a patient's sagging breast is inserted into a tubular portion of a patient's bed, and a circular imaging module surrounds the tubular body. .

  FIG. 5 is a side view of an imaging system according to the present invention including two rings of imaging modules.

  FIG. 6A is a plan view of an imaging system according to the present invention including a ring-shaped imaging detector surrounding a thin-walled tube, where the detector ring is scanned or moved along the vertical axis to fill the entire breast volume. And the detector ring can be rotated slightly around the thin-walled tube, providing a view of all the angles of view necessary for 3D tomography reconstruction.

  FIG. 6B is an imaging system according to the present invention that includes two ring sectors and the detector is rotated around the breast to provide a plan view of all necessary angles required for 3D tomography reconstruction To do.

  FIG. 7 is a cross-sectional view of a scintillation module having four standard square or round PMTs in an array coupled with a scintillator block in standard direct coupling mode.

  FIG. 8 is a cross-sectional view illustrating the application of a tilted fiber optic light guide that can be applied to the four standard square or round PMTs of FIG. 7 to enable the use of standard PMTs in a dedicated PET breast imager according to the present invention. .

  FIG. 9 represents a basic first 12.5 mm silicon PMT imaging module composed of sixteen 3 mm × 3 mm readout pixels in an array or pads arranged in a 4 × 4 array.

  FIG. 10 shows a 2 × 2 array of the basic four-side buttable modules of FIG. 9 that are joined together to form an approximately 1 inch square photo detector.

  FIG. 11 shows a larger module that can be implemented by arranging 16 (4 × 4 array) basic imaging modules.

  FIG. 12 is a perspective view of an imaging system according to the present invention, and a small basic form (prototype) that forms an active field of view that exceeds 2.5 cm and is 5 cm higher by using a commercially available detection element. Including detectors.

  Turning to FIG. 1, a first embodiment of a chest imaging system 20 of the present invention is shown and includes a set of photodetector modules 22 viewing a breast 24 in a lateral position, with the detector head covering the edge of the chest wall 26. Being lifted beyond, it allows for efficient detection of lesions in the breast portion near the chest wall. The imaging system has a lateral imaging geometry in which the detector head 22 is placed further apart, and the detector head 22 is lifted beyond the edge of the chest wall 26 so that the breast portion is close to the chest wall 26. It is designed to enable efficient detection of lesions. A lightweight patient bed 28 (desirably made of hard plastic or carbon fiber matrix) has a shape that approximates the portion of the cylinder that facilitates placing the detector head 22 above the chest wall 26.

  A second and preferred embodiment of the imaging system of the present invention, including a ring-shaped detection module surrounding the breast to be imaged, and designed to minimize the dead space at the ends. Turning to FIG. 2, an imaging or detector module 22 based on a scintillator array 30 coupled to a position sensitive PMT 32 via optical light guides 34 made of plastic fibers arranged in an array is shown. The slanted shape of the light guide 34 allows the end of the scintillation sensor element 30 to rest on a surface 36 that defines the upper end 38 of the photodetector module 22. Thus, the detector module 30 protrudes toward the patient's chest wall. In each imaging module 22, a tilt imaging light guide 34 is inserted between the gamma radiation sensor or scintillator 30 and the photoelectric detector 32. The light guide 34 allows the gamma sensors 30 to be placed almost directly beside the skin of the chest wall (see FIG. 3) so that they can extend the sensitive area of the imager to the base of the breast. Desirably, the tilted imaging light guide 34 is a multi-fiber optic light guide having an inclination angle of 5 to 45 degrees, most preferably an inclination angle of 10 degrees. An electronic reading device 39 is provided at one end of the PMT 32.

  Looking at the imaging system or breast imager 40 in FIG. 3, the imager or scanner ring 42 and a portion of the patient bed 44 showing that the imager module 46 is positioned to be thrust against the patient's chest wall 26. A cross-sectional view through is shown. A forming bed 44, an inner tube 48, and a side housing (not shown) surround the scanner ring 42. Within that enclosure, the scanner ring 42 can be rotated by a small angle and scanned up and down as indicated by the directional arrows 49 to sufficiently cover the imaging breast 24 and provide angular sampling by the photodetector module 46. Desirably, the patient bed 44 is constructed of a thin composite material or plastic. The patient bed 44 includes cushioning at the top for patient comfort. The cushioning material at the end of the inner tube 48 is shaped to allow a thin imager ring 42 to be placed proximate to the chest wall 26. Alternatively, as shown in FIG. 4, a conventional table 47 can be used with the mammography system of the present invention, eg, available from Lorad, Danbury, Connecticut, used in core biopsy procedures. Like a table.

  With reference to FIG. 4, a first embodiment of a chest imaging system 50 according to the present invention is shown. The patient inserts a breast 24 that hangs down into a cylindrical thin tube 48 in an annular imaging module 46 surrounding the tube 48. In this embodiment of the breast imaging system 50, a ring 42 of photodetector modules 46 is provided, which scans or moves along the vertical axis 49 along the vertical line 49 to scan the entire volume of the breast 24. The photodetector module 46 can rotate slightly around the thin-walled tube 48.

  A second embodiment of the mammography system 60 is represented in FIG. In the embodiment depicted in FIG. 5, a two-ring imaging module 46 is provided that includes an upper imaging ring 62 and a lower imaging ring 64, and a set of aging rings 62 and 64 to cover the entire volume of the breast 24. The whole can be moved. In this particular embodiment, a small physical gap 68 occurs between the two imaging rings 62 and 64 due to the dead edge of the PMT perimeter 66. The effect of gap 68 is minimal in the final quality of the reconstructed image. If desired, three or more rings (not shown) can be added with additional detector rings to eliminate the need for vertical scanning and cover the entire breast volume without requiring detector ring movement. A mammography system with

  In the chest imaging system of the present invention, the breast to be imaged hangs through the inner tube 48 of the patient bed 44 as shown in FIG. An alternative embodiment for the imaging procedure would include breast compression, especially when the breast is fixed for an image-guided biopsy. Several arrangements are available for breast immobility. Place the breast in a plastic container or bag, apply sub-atmospheric pressure with a small pressure differential, etc. The pressure differential will be created by a low power pump. An alternative device for breast immobilization, as proposed by Karimian et al., Includes a vacuum pump to fix the saphenous breast in the center of the field of view (FOV) (“CYBPET” Physics Research A545 (2005) 427-435 “ Chest image cylinder PET system, Nuclear Instruments, and methods). In that example, a nominal diameter of approximately 20 centimeters (suitable for almost all adult female breast sizes) surrounds an individual posterior breast under moderately reduced pressure in a thin soft plastic membrane. .

  FIG. 6A shows a plan view of a second and preferred embodiment of an imaging system 70 according to the present invention. The chest imaging system 70 includes a ring 71 of an imaging module 72 that surrounds a thin tube 73 in which a breast (not shown) is placed. Each imaging module 72 includes a scintillator sensor 74, a fiber optic light guide 75, an optical window 76 at either end of the light guide 75, and a PMT enclosure 77 mat contains a PMT and associated electronics.

  FIG. 6B shows a top view of a third embodiment of an imaging system 78 similar to the imaging system of FIG. 6A, but with the imaging module 72 in two imaging ring sectors 79A and 79B. The imaging module 72 rotates around the breast (not shown) to provide all the angular fields necessary for 3D tomography reconstruction. This embodiment of the breast imager 78 is an economical arrangement consisting of only two imaging ring sectors 79A and 79B. This embodiment of the breast imager 78 provides a convenient side port 80 for installing a live device. The imager 78 provides the convenience of performing live procedures while imaging the breast. By using a radioactive tip biopsy needle (not shown), the biopsy needle tip is guided to the lesion by the imaging system 78.

Referring to FIG. 2, a dedicated breast PET imager according to the present invention can comprise several imaging techniques. A preferred general type of solution has a scintillator sensor 30 as a 511 keV annihilation gamma ray sensor / energy converter, where various photoelectric detectors 32 are used as detectors of scintillation light generated by 511 keV gamma rays absorbed. Can be helpful. The scintillator sensor 30 can be made from a plate or pixelated array crystal scintillator material. Most desirably, the scintillator sensor 30 LSO, LYSO, LuAP, LFS, GSO, GYSO, BGO, NaI (Tl), or pixelated such LaBr _3, made from a high stopping power crystal scintillator material. Photoelectric detector 32 may be a standard or multi-element photomultiplier tube or a position sensitive, flat panel or microchannel plate based photomultiplier tube. Alternatively, the photoelectric detector 32 may be an avalanche photodiode array or a large avalanche photodiode having a resistive readout device. As another alternative, the photoelectric detector 32 may be a silicon photomultiplier tube. Examples of commercially available flat panel photomultiplier tubes are model H8500 or H9500 from Hamamatsu, Bridgewater, NJ or Phototonis USA, Inc. of Starbridge, Massachusetts. Several models of the 85000 series available from: Each of these position sensitive photomultiplier tubes has an active surface area of approximately 5 cm × 5 cm.

  A more economical alternative solution to the basic detector module 22 of the external detector head is available from Photonis Inc. It can be based on a multi-element PMT such as a 2-inch square 9-element XP 1470 PMT with a small profile available from (USA). However, the intrinsic spatial resolution achieved cannot be as low as satisfactory for a 2 mm pitch scintillation pixel. In addition, these PMTs are much slower than the Hamamatsu H8500 flat panel PMT, mainly due to the time shift between the individual 9 inner channels. Another difficulty with the system built on the XP1470 PMT is that it is quite bulky.

  7 and 8, a standard square PMT can be used to form an edgeless detector design according to the present invention, and even a round PMT can be used. FIG. 7 is a cross-sectional view of scintillation module 81 in which an array of 2 × 2 standard round PMTs 82 are coupled to scintillator block 84 in a standard direct coupling mode. In this example, the dead edge portion is much larger when used in standard direct coupling mode than in the case of Position Sensitive PMT. Nevertheless, as shown in FIG. 8, the application of a tilted fiber optic light guide 86 can create a scintillator block 88 that can be used with the dedicated PET breast image of the present invention. FIG. 8 is a cross-sectional view depicting the application of a tilted fiber optic light guide 86 applied to the four standard round PMTs 82 of FIG. 7, allowing the use of the round PMTs 82 in a dedicated PET breast imager according to the present invention. . An example of a two-layer phosswitch scintillator (quoted from the Siemens high-resolution brain PET HRRT detector module design), which has a 2x2mm scintillation pixel of 10mm length and LYSO at the top and LSO at the bottom. 1 shows a light guide 89 that is molded.

  The tilted light guide 86 according to the present invention can be constructed from a variety of materials as described below. Glass, quartz, polystyrene, acrylic, polyvinyl toluene (PVT), or plastic with a liquid core light guide. Depending on the material, the light guide element may be straight or bent in a size or diameter range from 10 microns to 3 mm. A dense and regularly arranged array of these individual light guide elements or fibers forms a slanted light guide whose transverse size matches the size of the photodetector module. For example, in the case of an H8500 or H9500 flat panel PMT, a size of about 52 mm square is desirable. The length of the light guide 86 may be several centimeters to 20 centimeters or more. In addition, an optical window 89 is introduced between the scintillation element 84 and the light guide 86 and between the light guide 86 and the surface of the photoelectric detector 82 to form a fiber or other light guide 86. This array of elements can help receive, transmit and deliver the optical signal. For example, in the case of a scintillation array, the quantized nature of both scintillator 84 and light guide 86 helps such windows to spread the scintillation light from each scintillation array pixel to more fibers. , Pixel-to-pixel sampling uniformity can be improved. Preferred building materials for optical windows include glass, quartz, and plastic. The fiber optic window may be straight or tapered. The scintillation array 84 has a light-tight window (not shown) composed of about 1 mm thick aluminum or equivalent as a protective sheet or shell for light sealing and mechanical protection. Can be included. Desirably, a plastic or metal shell or box (not shown) with an opening in front of the scintillation array 84 is used to house the detector module. In a breast imager according to the present invention, all ring modules, such as rings 62 and 64 of FIG. 5, are protected by a thin outer epidermis (not shown) and built in the patient's bed.

  With reference to FIG. 3, a specific embodiment for a high-resolution chest imaging system with improved detection efficiency for lesions located near a patient's chest wall includes a ring 42 of individual detection modules. Each based on a Hamamatsu H8500 flat panel PMT 46 is coupled to an array 34 of 24 × 24 LYSO 2 × 2 × 15 mm pixels at a pitch of 2.1 mm. Each module 46 has four analog position outputs and one fast energy output. The position output is recorded and digitized with a DAQ system to calculate the position of interacting 511 keV gamma rays. The fast sum signal is used to create a simultaneous event trigger at the trigger electronics and then provided to the DAQ system to record event data from the detector module. The optimal on-board readout design includes PMT gain uniformity correction that ensures high energy, spatial resolution, and high rate performance with a minimum number of readout channels.

  In the breast imager according to the present invention, silicon PMT is the preferred material for the photoelectric detector and can be used in place of position sensitive PMT. Typically, silicon PMT modules are provided in small units of about 3 mm in size. Thus, an array of these devices is necessary to cover the same active field. FIGS. 9-11 show a silicon PMT-based, approximately 5 cm × 5 cm effective field photoelectric detector 90 using a nominal 12.5 mm silicon PMT basic imaging module 92 (FIG. 9), each consisting of 16 3 mm silicon PMT units 94. The example of the achievement method of (FIG. 11) is shown. As shown in FIG. 9, the basic first imaging module 92 can have an array of sixteen 3 mm × 3 mm readout pixels or pads 94 arranged in a 4 × 4 array. The 3 mm pads or units 94 have four pads 94 connected to one readout channel 95 (roughly 96) (16 in total) as conceptually shown in the center and bottom of FIG. Can be read with a single pad 94 connected to one read channel 96. As shown in FIG. 10, a 2 × 2 array of these basic 4-side buttable modules (consisting of 8 × 8 = 64 3 mm silicon PMT units) forms an approximately 1 inch square photoelectric detector 97. However, this is equivalent to the commercially available Hamamatsu R8520-C12 PSPMT. With a dead space of approximately 1 mm at the end, the basic imaging module 92 can be matched on four sides. As shown in FIG. 11, the photodetector module is preferably arranged in a large imaging module 90 composed of 4 × 4 basic modules 92, but at that time, the application range and the necessity of the readout device are As described above, it is equivalent to the H8500 or H9500 flat panel PMT. The large imaging module 90 can be used as a plug-in replacement module for an H8500 / H9500 flat panel PMT having, for example, an active surface of approximately 5 cm × 5 cm. The dead area 98 between the basic modules is about 1-2 mm wide, which allows for a scintillation signal appropriate for uniform energy and spatial response.

  FIG. 12 shows a conceptual diagram of a detection module made in accordance with the present invention having a small silicon photomultiplier tube sensor. The detector used in this example is a 3 × 3 mm-MPPC-SMD package from Hamamatsu. FIG. 12 is a conceptual diagram depicting the formation of a miniature detector module 100 using a Hamamatsu S 10362-33 silicon multi-pixel photon counter (MPPC) 102. The detection module includes a scintillator array 104, a tilted fiber optic light guide 106, and a silicon PMT array 108. Each detection element 102 occupies an area of about 5 mm square in the array 108. In the example depicted in FIG. 12, an effective field of view exceeding 2.5 cm (width) and 5 cm (height) is equivalent to one set of two R8520-C12 PSPMTs or a flat panel PMT of H8500 / H9500 half , 6 × 11 silicon PMT array 108. By coupling the scintillation array 104 with the photodetector array 108 via a tilted fiber optic light guide 106, a scintillation array readout device is obtained. With a sampling pitch of 5 mm and an effective size of each photoelectric detector element of approximately 3 mm × 3 mm, the average active surface coverage is about 35%.

  An important and integral part of the breast imager of the present invention receives and digitizes the signal from the readout device and then transfers the digitized data to a computer system for data processing, data analysis and tomographic image reconstruction. It is a data collection system. The readout device depends on the particular choice of technical solution selected for the imager module.

  In a first embodiment of the data collection and processing system according to the present invention, the breast PET imager places 12 detector modules in the ring. Each module utilizes a Hamamatsu H8500 flat panel PMT combined with a pixelated LYSO scintillator array. Each PMT amplifier board supplies four position-encoded analog signals and one analog sum signal to a total of 60 channels per detector. The trigger is formed in a separate trigger hardware module. In the simplest form of trigger circuit, if any two of the twelve modules in the ring are accidentally synchronized, the logic simultaneous circuit provides a single simultaneous trigger to the data acquisition (DAQ) unit. Such a function is obtained, for example, by Mesytec MSCF-16 available from Mesytec GmbH, Putzbrunn, Germany, which has a multiplicity trigger function.

  A preferred data collection system is an FPGA-based USB data collection system, 2005 IEEE Nuclear Science Symposium Conference Record, Perto Rico, October 23-29, 2005, pp. 228. In 2971-2975, article of "A flexible high-rate USB2 data acquisition system for PET and SPECT imaging" described by Proffitt other, and, 2006 IEEE Nuclear Science Symposium Conference Record, San Diego, California, October 29 -November 1, 2006, pp. 3063-3067 is a system described in the article “Implementation of a High-Rate USB Data Acquisition System for PET and SPECT Imaging” described by Profitt et al., Which is incorporated herein by reference in its entirety. Include in the description. The FPGA-based USB DAQ system has a modular and scalable architecture that can scale up to 64 channels of simultaneous sampling ADCs per unit, with a sustained trigger rate of over 150 kHz for all 64 channels. In the case of a multi-ring (2-3 rings) system, each unit corresponds to one ring module. Each DAQ unit sends its raw event data on a high speed USB to its own collection computer. Each collection computer performs centroid and energy calculations on all received data and sends the time stamped processing data over Gigabit Ethernet to the event builder / reconfiguration computer. The event builder uses the time stamp to merge separate detector events into a single concurrent event. It can perform image reconstruction or send the data to another computer for image reconstruction. Send the reconstructed tomographic image set to the user interface computer. Other readout devices with similar functionality are commercially available and can be used to achieve the same goal of recording and processing data from the proposed breast imager.

  Many variations of commercially available tomography reconstruction software packages can be used with the ring imagers described herein. The primary function of detecting cancer lesions across the entire breast volume, up to the base of the breast, is achieved by using any of these available reconstruction algorithms described in the literature and / or available from software vendors. it can.

  As an example, to improve the detection efficiency of lesions located near the patient's chest wall, a high resolution PET imaging module with a tilted fiber optic light guide was made in accordance with the present invention. A basic module was used to experimentally demonstrate the idea of implementing a fiber optic tilted light guide with a detector module. The basic (prototype) light guide was made of a regular array of loose arrays bonded with 2 mm square double-coated polystyrene fibers available from Saint-Gobain Crystals, Hiram, Ohio. The light guide parameters were 60.5 mm x 62 mm surface, 56.5 mm long, 10 degree tilt angle, and the resulting 12 mm side shift. The fiber optic light guide light was efficiently transmitted while maintaining the geometric relationship between the pixels in the radiation sensor or scintillator. Other parts of this basic module for high-resolution PET breast imagers were a 50 mm × 50 mm LYSO array of 2 × 2 × 10 mm pixels available from Proteus of Chagrin Falls, Ohio, an H9500 flat panel PMT from Hamamatsu, and others. Dry optical coupling was used for all optical planes. The high resolution PET breast imager module achieved a 12 mm shift that is sufficient to bypass or eliminate the dead edge of the detector module. This demonstrated the ability of the PET breast imager of the present invention for improved detection efficiency for lesions located near the chest wall.

  In the original image obtained with the LYSO scintillation array, each scintillation pixel was very well separated. When irradiated with a Na22 source, an energy spectrum was obtained from a single LYSO pixel. In the energy spectrum shown in the graph, the peaks at 511 keV and 1274 keV were clearly visible. The energy resolution at 511 keV was 16% FWHM.

  A comparison was made between the response without the tilted light guide and the LYSO scintillation array placed directly on top of the H9500 flat panel PMT. In the original image without the light guide, as expected, the scintillation pixels separated better than the image obtained with the light guide. In the energy spectrum obtained from a single LYSO pixel irradiated with a Na22 source, the average energy signal was about twice as high as when the light guide was inserted. The energy resolution at 511 keV was 13% FWHM.

  Comparison of response without tilted light guide and LYSO scintillation array placed directly on top of H9500 flat panel PMT. The comparison showed that: The scintillation signal when the light guide is in place is about half that of the direct signal, but the performance of the energy resolution remains very good, and with the spatial resolution, in a bright scintillator with a pixel size of 1-3 mm Operation became possible. The results of this assembled first module for a high resolution PET breast imager fully confirmed that the imager is capable of improving the detection efficiency of lesions located near the patient's chest wall. The scintillation signal detected at the PMT level was strong and provided very good spatial resolution and excellent energy performance.

  While the above description includes a number of specific descriptions, materials, and dimensions, these should not be construed as limiting the scope of the invention, and some of the presently preferred embodiments of the invention It should be construed as merely providing an explanation. Accordingly, the scope of the invention should be determined not by the embodiments illustrated, but by the appended claims and their legal equivalents.

20 Breast Imager 22 Photodetector Module 24 Breast 26 Chest Wall 28 Patient Bed 30 Scintillator Sensor 32 Position Sensitive Photomultiplier Tube (PSPMT)
34 optical light guide 36 surface defining the upper end of the photodetector module 38 upper end 39 of the photodetector module readout 40 breast imager 42 scanner ring 44 part of patient bed 46 photodetector module or imager module 47 patient bed 48 Inner tube 49 Directional arrow 50 Breast imaging system, first embodiment 60 Breast imaging system, second embodiment 62 Upper imaging ring 64 Lower imaging ring 66 PMT perimeter 68 Gap 70 Breast imaging system, second implementation Form 71 Imaging Module Ring 72 Imaging Module 73 Thin-walled Tube 74 Scintillator Sensor 75 Fiber Optic Light Conductor 76 Optical Window 77 PMT Enclosure 78 Breast Imaging System, Third Embodiment 79A First Imaging Ring Sector 7 B 2 th imaging ring sectors 80 side opening 81 the scintillator module 82 round PMT
84 scintillator block 86 fiber optic light guide 88 scintillator block 90 large silicon imaging module or photoelectric detector 92 basic silicon imaging module 94 3 mm basic silicon unit or pad 95 readout channel 96 composed of 4 pads 96 readout consisting of 16 pads Channel 97 1 inch square silicon photo detector 98 dead area 100 small detector module 102 using silicon MPPC silicon MPPC
104 scintillator array 106 inclined optical fiber light guide 108 silicon PMT array

Claims (20)

A breast imager for efficient detection of lesions located near a patient's chest wall,
A table for receiving a patient in an abdominal position, the table including an open tube extending downward to receive the breast of the patient;
One or more rings comprised of a gamma ray detection module surrounding the breast in the tube; and
A breast imager having a data collection system for receiving imaging data from the readout device of the gamma ray detection module,
Each of the gamma ray detection modules is
A scintillator sensor adjacent to said open tube;
An inclined imaging light guide extending from the scintillator sensor;
A photomultiplier tube extending from the light guide;
A readout device connected to the photomultiplier;
Including the breast imager. The breast imager of claim 1,
The breast imager detects scintillation light generated by 511 keV gamma rays absorbed by a scintillator sensor;
And the breast imager has a spatial resolution of better than 2 mm,
Above breast imager. The breast imager of claim 1,
The detection module includes a tip surface and a dead edge whose detection efficiency drops sharply at the tip surface;
The tilted imaging light guide has a tilt angle of 5 to 45 degrees;
And the tilted imaging light guide allows for side shifting of the image, the shifting of the image allows for the exclusion of the edge on one side of the detection module,
A breast imager as described above.   The breast imager of claim 1, wherein the one or more rings of the gamma ray detection module can be scanned vertically toward and away from the chest wall of the patient so that the entire breast is imaged. The breast imager described above is possible. The breast imager of claim 1,
The breast imager comprises three or more of the rings of gamma ray detection modules;
The plurality of rings of the gamma ray detection module are stationary with respect to the table; and
The plurality of stationary rings enables imaging of the entire breast with the gamma detection module fixed at one position relative to the table.
A breast imager as described above. The breast imager of claim 1,
The plurality of rings of gamma ray detection modules are divided into two or more ring sectors, each of the ring sectors having a gamma ray detection module on its own;
And the ring sector containing the gamma ray detection module is rotated around the breast to provide a more angled field of view for improved 3D tomography reconstruction;
A breast imager characterized by that. The breast imager of claim 1, wherein the gamma ray detection module is
An scintillation module having four round PMTs in an array;
And a tilted optical fiber light guide coupled in a direct coupling mode between the scintillation module and the round PMT;
A breast imager characterized by comprising:   2. The breast imager of claim 1, wherein the scintillator sensor is selected from the group comprising a high stopping power plate crystal scintillator material and a pixelated high stopping power crystal scintillator material. The breast imager of claim 1,
The scintillator sensor is made of pixelated high stopping power crystalline scintillator material;
And said pixellated high stopping power crystal scintillator material, LSO, LYSO, LuAP, LFS , GSO, GYSO, BGO, NaI (Tl), and breast imager, characterized in that it is selected from the group comprising LaBr _3. The breast imager of claim 1,
The ring is embedded in the table, allowing the end of the gamma ray detection module to be positioned directly on the chest wall;
Geometrical arrangement of the ring and the gamma ray detection module significantly reduces the dead area of the breast imager;
A breast imager characterized by that.   The breast imager of claim 1, wherein the table is made of a material selected from the group comprising a carbon fiber matrix, plastic, and lightweight metal.   The breast imager of claim 1, wherein the photomultiplier tube is a standard square photomultiplier tube, a standard round photomultiplier tube, a multi-element photomultiplier tube, a position detection type flat panel photomultiplier tube, A breast imager selected from the group comprising a position sensitive microchannel plate-based photomultiplier tube, a large avalanche photodiode with a resistive readout, and a silicon photomultiplier array.   2. The breast image of claim 1, wherein the tilted imaging light guide material is selected from the group comprising glass, quartz, polystyrene, acrylic, polyvinyl toluene, and a liquid core light guide housed in a small diameter TEFLONR tube. A breast imager characterized by The breast imager of claim 1,
The tilted imaging light guide includes an array of individual light guide elements;
Breast imager characterized in that the light guide element comprises a diameter or square profile (profile) between 10 microns and 3 mm.   2. The breast imager of claim 1 wherein the photomultiplier tube is a compact silicon photomultiplier tube array. The breast imager of claim 1,
Having an optical window between the scintillation sensor and the light guide;
The optical window enhances the transmission of light through the light guide;
A breast imager characterized by that.   The breast imager according to claim 1, further comprising an optical window between the light guide and the photomultiplier tube. The breast imager of claim 1,
The photomultiplier tube is an array of 3 mm × 3 mm silicon photodetector units;
The silicon units are arranged in a 16 unit 4x4 array to form a basic butt silicon PMT module;
Four of the basic buttable silicon PMT modules are arranged in a 2x2 array to form a nominal 1 inch square photodetector;
A breast imager characterized by that. A breast imager for efficient detection of lesions located near a patient's chest wall,
A table for receiving a patient in an abdominal position, the table including an open tube extending downward to receive the breast of the patient;
One or more rings comprised of gamma ray detection modules surrounding the open tube; and
A breast imager having a data collection system for receiving imaging data from the readout device of the gamma ray detection module,
Each of the gamma ray detection modules is
A scintillator sensor adjacent to said open tube;
An inclined imaging light guide extending from the scintillator sensor;
A photomultiplier tube having an array of compact silicon photodetector units in the form of a photomultiplier tube extending from the photoconductor;
A readout device connected to the photomultiplier;
Including
The one or more rings of the gamma ray detection module can be scanned vertically toward and away from the patient's chest wall;
The one or more rings of the gamma ray detection module are slightly rotatable around the breast in the tube;
A breast imager as described above. A breast imager for efficient detection of lesions located near a patient's chest wall,
A table for receiving a patient in an abdominal position, the table including an open tube extending downward to receive the breast of the patient;
Three rings composed of gamma ray detection modules surrounding the breast; and
A breast imager having a data collection system for receiving imaging data from the readout device of the gamma ray detection module,
Each of the gamma ray detection modules is
A scintillator sensor adjacent to said open tube;
A tilted imaging optical fiber light guide extending from the scintillator sensor;
A photomultiplier tube extending from the light guide;
A readout device connected to the photomultiplier;
Including the breast imager.

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