WO2004109332A1 - Generating detector efficiency estimates for a oet scanner - Google Patents
Generating detector efficiency estimates for a oet scanner Download PDFInfo
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- WO2004109332A1 WO2004109332A1 PCT/GB2004/002340 GB2004002340W WO2004109332A1 WO 2004109332 A1 WO2004109332 A1 WO 2004109332A1 GB 2004002340 W GB2004002340 W GB 2004002340W WO 2004109332 A1 WO2004109332 A1 WO 2004109332A1
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/29—Measurement performed on radiation beams, e.g. position or section of the beam; Measurement of spatial distribution of radiation
- G01T1/2914—Measurement of spatial distribution of radiation
- G01T1/2985—In depth localisation, e.g. using positron emitters; Tomographic imaging (longitudinal and transverse section imaging; apparatus for radiation diagnosis sequentially in different planes, steroscopic radiation diagnosis)
Definitions
- This invention relates to a method of, and computer software for, generating detector efficiency estimates for a positron emission tomography (PET) scanner including a detector array for generating detection data and a single photon source.
- PET positron emission tomography
- a typical emission scan begins with the injection of a solution including a tracer, which is a pharmaceutical compound including a radio-isotope with a short half-life, into the subject.
- the subject may be human or animal.
- the tracer moves to, and is typically taken up, in one or more organs in the subject according to biological and biochemical processes which occur within the subject.
- the radio-isotope decays, it emits a positron, which travels a short distance before annihilating with an electron. This annihilation produces two high energy photons propagating in opposite directions.
- the PET scanner includes a photon detector array arranged (usually in a ring-shaped pattern) around the scanning area.
- coincidence count data from each sinogram bin is typically processed using tracer uptake models and image processing techniques to obtain volumetric medical images and volumetric tracer uptake rate data for the subject.
- the scanner is provided with one or more positron emitter rod sources, formed of a material such as Ge, which emit dual annihilation photons.
- a blank scan in which the subject being scanned is not present in the scanning area (typically, the scanner is empty except for the presence of the sources) and a transmission scan in which the subject is present in the scanning area.
- the source material is a positron emitter
- the two photons arising from the annihilation of a positron and an electron are acquired in coincidence, in the same manner as with an emission scan.
- the results of the blank scan are then divided by the results of the transmission scan to derive an attenuation sinogram.
- the attenuation sinogram is then used to correct the emission scan for attenuation, although extra processing steps on blank scan data and transmission scan data are sometimes used.
- estimates of the detector efficiency may be calculated by adding together the counts detected between a particular detector and all the detectors that are in coincidence with it. This is known as a fansum and it is based on the following measurement model:
- My S f S j Ay (1)
- My are the measured coincidence counts between detectors i andy, €; and e j are the intrinsic efficiencies of detectors i and / respectively and Ay are the ideal coincidence counts measured if the detectors had an ideal performance.
- the measured counts My of a particular coincident pair of detectors are proportional to the product of the individual detector efficiencies.
- An estimate of each individual detector efficiency e ⁇ is made by summing the measured coincidence counts My over all detectors / in coincidence with detector /:
- a single photon source is provided for the transmission scan.
- the source a Cs pellet
- the transport mechanism having a spiral tube that is placed in front of the detectors when a transmission scan takes place.
- the source is moved at a constant speed through the scanner in a helical motion to provide a full 3D transmission scan.
- Some PET scanners have detector arrangements which rotate during an acquisition. Typically, the detectors are arranged in two or more banks which do not fully surround the subject. Alternatively, the detectors may be arranged in a non-ring-shaped pattern. In these cases, the model given by equation (1) above is no longer appropriate, as the rotation causes the direct inter-relation between detectors and entries in the sinogram bins to break down. Every sinogram bin typically contains counts detected by multiple detector-pairs.
- the transmission measurement is again performed using a single photon point source. However, the source is transported only axially relative to the (rotating) detector arrays. The ECAT ART scanner contains two such sources, one located on each of the opposing detector banks.
- the ECAT ART scanner is described in further detail in the article "The ECAT ART Scanner for Positron Emission Tomography: "1. Improvements in Performance Characteristics", David W. Townsend et al., Clinical Positron Imaging, Vol. 2, No. 1, 1999 and the article “Design and Performance of a Single Photon Transmission Measurement for the ECAT ART” C.C. Watson, W.F. Jones, T. Brun, K. Baker, K. Naigneur and J. Young, LEEE Medical Imaging Conference Record M9-02 ; 1998.
- a phantom scan involves positioning a body, referred to as the phantom, containing a positron-emitting radioactive source material in the scanner.
- a phantom scan is either an emission or a transmission scan with a 'phantom' in the scanner instead of a patient.
- the 'phantom' is an object made out of plexiglass or suchlike and filled with water mixed with a radioactive substance.
- the phantom has known shape and attenuation characteristics.
- a phantom emission scan has in the past been used to calculate detector efficiency estimates.
- a problem with having to conduct a phantom scan is that the phantom must be handled by an operator to place the phantom inside the scanner and to subsequently remove the phantom. It is both inconvenient and time-consuming, since in any event an operator needs to be available to handle the phantom at the start and end of the procedure. It is also potentially hazardous, in terms of lifting the phantom and exposure to radio
- a method of generating detector efficiency data for a positron emission tomography scanner including: a detector array for generating detection data; and a single photon source, wherein the method comprises: conducting an acquisition procedure using the single photon source to produce detection data; and processing said detection data using an efficiency estimation algorithm to calculate data representative of the efficiencies of individual detectors in said array.
- detector efficiencies can be generated without significant inconvenience to an operator.
- the detector efficiencies may be derived from a blank scan acquisition conducted at the operator's convenience. Furthermore, the regular need for the use of a phantom scan procedure can be avoided.
- detector efficiency estimates are made using the artificial coincidence counts generated during a blank scan acquisition made using the single photon source.
- artificial coincidence counts are the only suitable detection data made available as an output from a scanner of a type such as an ECAT EXACT3D PET scanner.
- detector efficiency data cannot be accurately estimated from artificial coincidence counts using known techniques, because the known measurement models do not apply.
- the present invention provides a new measurement model and exemplary efficiency estimation algorithms, which can be applied to artificial coincidence counts produced using blank scans. Further features and advantages of the present invention will become apparent from the following description of preferred embodiments of the present invention, made by way of example only with reference to the accompanying drawings.
- Figure 1 is a schematic diagram of a PET scanning facility
- Figure 2 is a schematic diagram of data processing components in a non- rotating PET scanner and an associated data processing terminal
- Figure 3 is a schematic cross section of the arrangement within a non- rotating PET scanner during an emission scan acquisition
- Figure 4 is a schematic cross section of the arrangement within a non- rotating PET scanner during a blank scan acquisition using a single photon source;
- Figure 5 is an illustration of a user interface provided in an embodiment of the invention for presenting detector efficiency estimates to an operator;
- Figure 6 is a schematic cross section of the arrangement within a rotating PET scanner during a blank scan acquisition using a single photon source.
- a PET scanning facility arranged in accordance with an embodiment of the invention, which includes a PET scanner 2, a tracer generator module 4 and an operator computer terminal 6.
- the scanner 2 includes a detector array 8 arranged about a scanning area, in which a subject 10 is located during a transmission scan and during an emission scan.
- the PET scanner 2 includes a control unit 30, detection data processing circuitry 32, one or more single photon source position detectors 34, a count store 36 and an Input/Output (I/O) device 38.
- the computer terminal 6 includes a central processing unit (CPU) 42, memory 44, hard disc drive 46 and I/O device 40, which facilitates interconnection of the computer 6 with the PET scanner 2.
- Operating system programs 48 are stored on the hard disc drive 46, and control, in a known manner, low level operation of the computer terminal 6.
- Program files and data 50 are also stored on the hard disc drive 46, and control, in a known manner, outputs to an operator via associated devices.
- the associated devices include a display, a pointing device and keyboard (not shown), which receive input from, and output information to, the operator via further I/O devices (not shown).
- Included in the program files 50 stored on the hard drive 46 are a detector efficiency calculating application 52 and a detector efficiency user interface application 54.
- a database 56 is used to store the detection data transferred from the PET scanner 2.
- the detection data processing circuitry 32 of the PET scanner operates in two different modes. In a positron source mode, the detection data processing circuitry 32 processes all events detected in the detector array 8, and by using a coincidence timing window, detects coincidences between the events which are recorded as real coincidence counts My. These real coincidence counts are then output to sinogram bins in the count data store 36 for subsequent processing.
- the detection data processing circuitry 32 In a single photon source mode, the detection data processing circuitry 32 generates artificial coincidence counts when a single photon source is exposed in the scanner, as will be described below in further detail.
- the one or more single photon source position detectors 34 provide a position detection output to control unit 30 which generates a dynamic position estimate during the acquisition, which is sent to the detection data processing circuitry 32.
- the detection data processing circuitry 32 filters the detector outputs so as to discard all events detected on the side of the detector array nearest the current source position.
- the detection data processing circuitry 32 generates an artificial coincidence count M'y, and transmits the artificial coincidence count data to sinogram bins in the count data store 36 for subsequent processing.
- One implementation of the invention which is described in further detail below, relates to a non-rotating PET scanner.
- a PET scanner of the ECAT EXACT3DTM type is used.
- the detectors in the detector array 8 are arranged in square detector blocks, each containing 64 detector elements.
- the detectors are arranged in six rings of detector blocks. There are 48 detector blocks in each ring.
- the total number of lines of response (LORs) that can be acquired is thus, in this embodiment, about 190 million.
- Figure 3 illustrates features of operation of the non-rotating PET scanner in a positron source mode during an emission scan.
- Figure 3 shows a positron emission event being registered in the detector array 8.
- the positron 60 annihilates and generates a first photon 62 travelling in one direction and a second photon 64 travelling in an opposite direction.
- the first photon is detected by a detector element 70 on one side of the detector array 8, and the second photon 64 is detected in a different detector element 72 on the other side of the detector array 8.
- the detection data processing circuitry 32 registers the two as a coincidence along the LOR defined between the two different detectors 70, 72.
- FIG 4 illustrates features of operation of the non-rotating PET scanner in a single photon source mode during a blank scan.
- a single photon source 80 is transported around the periphery of the scanning area using a single photon source transport mechanism 82.
- the single photon source 80 referred to as a "point" source due to its small diameter, consisting of a pellet of 137 Cs is driven in a fluid-filled steel tube wound into a helix and positioned just inside the detector array 8.
- the tube diameter is 4mm or smaller and the 137 Cs point source, with a radioactivity typically between 1 to 10 mCi, has a diameter of 2mm or less.
- the source is transported in a liquid carrier and is buoyancy matched to the specific gravity of the liquid. The arrangement is designed to increase consistency of the speed of travel of the source.
- the source carrier is driven by a pump (not shown) at a speed of about 1ms "1 .
- the source 80 is shown in Figure 4 emitting a single photon 90 which is detected at a detector 94 on the opposite side of the detector array 8 to the source 80.
- Figure 4 shows a virtual second photon 92, representing the artificial coincidence registered in the detector data processing circuitry 32 as described above.
- An event along an LOR is registered between the two respective detectors involved, the detector 94 registering the event and the detector 96 immediately adjacent the current source position 80.
- the detection data processing circuitry 32 generates an artificial coincidence count which is transmitted to the count data store 36.
- a measurement model for an acquisition using the single photon source is provided, which takes into account the method whereby the artificial coincidence counts are produced.
- the measurement model has the form:
- M' ⁇ ⁇ t A'y+ ⁇ j A' j i M'y are the measured artificial coincidence counts between detectors i and j, € t and € j are the intrinsic efficiencies of detectors i and j respectively and A 'y are the ideal artificial coincidence counts measured if the detector efficiency of detector i had an ideal performance, while the detection efficiency of detector j was 0.
- This measurement model takes into account that an event M'y recorded between detectors i and / will have originated either from detector i ovj. These will have each have a respective inherent efficiency e,- or e,.
- a ',y A ' ⁇ .
- the measurement model used is additive, in that the measured artificial coincidence counts M'y of a particular artificially coincident pair of detectors are equal to a weighted sum of their individual efficiencies.
- an algorithm is provided for estimating detector efficiencies.
- Equation (4) can then be rearranged and solved as an efficiency estimation algorithm using the following iterative formula:
- the Maximum Likelihood methods applicable to image reconstruction in PET such as the MLEM algorithm as described in the article "Maximum likelihood reconstruction for emission tomography", Y.Nardi, L.A Shepp, IEEE Trans. Med. Imag., Vol. 1, pp.l 13-121,1982, may be used.
- the MLEM algorithm we introduce some notation as follows. We use the index z to run over all sinograms bins. We define a matrix P to concisely write the linear measurement model as:
- M z ⁇ P ⁇ J ⁇ j (6) j or in matrix notation:
- the detector efficiency calculating component 52 implements the efficiency calculating algorithm as provided hereby.
- a PET scanner can perform a transmission scan in either 2D or 3D mode.
- both 3D blank and 3D transmission acquisitions can be conducted. Data from several LORs are then combined together in a process called spanning to save data storage space. This means that the above measurement model (4) is no longer appropriate.
- a blank scan acquisition is preferably conducted in a 2D mode (in which no spanning occurs), and from this individual detector efficiencies can be calculated.
- the 3D blank scan data may be used to calculate detector efficiencies in accordance with the invention.
- FIG. 5 illustrates a display part of a user interface generated on the computer terminal 6 by the detector efficiency user interface application 54 using the detector efficiency data calculated by data efficiency calculating application 52 after a blank scan procedure as described above.
- the display includes a detector representation 100, showing each of the detectors in the detector array 8 on a flat grid 100.
- the flat grid 100 is divided into detector block squares 102, representing each detector within the block as a separate pixel for the different brightness intensity.
- the brightness intensity used in each pixel represents the calculated efficiency estimate for the respective detector.
- a relatively dark pixel 104 is used to illustrate a detector having a relatively poor efficiency.
- a relatively bright pixel 106 is used to illustrate a detector having a good, fully operative efficiency.
- An intermediate brightness intensity pixel 108 is used to represent a detector having an intermediate efficiency.
- the detector efficiency user interface application 54 is provided with one or more automated block impairment check routines which are used to highlight blocks which are determined to be below a given threshold of acceptable operation.
- a first check routine an average efficiency is taken across each block. If the average detector efficiency within any block is below a preset threshold, the relevant block square 102 is highlighted, for example by the addition of a coloured border or a flashing display of the block square 102.
- the coefficient of variation of efficiencies within a block is calculated.
- the detector efficiency user interface application 54 then highlights, for example by the addition of a coloured border or a flashing display, any detector square 102 representing a block having a coefficient of variation which is higher than a preset threshold.
- Other implementations of the invention relate to a rotating PET scanner.
- a scanner of the ECAT ARTTM type is used.
- this implementation of the invention applies also to other rotating scanners which include a single photon source.
- the detectors in the detector array are arranged in square detector blocks, each containing 64 detector elements. The detectors are arranged in two banks of detector blocks.
- FIG. 6 illustrates features of operation of the rotating PET scanner in a single photon source mode during a blank scan.
- the scanner has first and second detector banks 120, 122 arranged on different sides of the patient.
- the two detector banks are arranged on a rotating mechanism which allows the banks to rotate about the scanning area.
- a single photon source 124, 130 is held in a single photon source transport mechanism 126, 132 on one side of each of the respective detector banks 120, 122.
- the single photon sources 124, 130 are thus mounted to, and rotate with, the detector array.
- Each single photon source transport mechanism is arranged to transport the source axially of the rotating detector array during a blank scan.
- a single photon source shield 128, 134 is provided for each single photon source. During a blank scan, one or each of the single photon source shields 128, 134 may be retracted to expose the single photon source in the interior of the scanner.
- Figure 6 shows an arrangement in which the first single photon source 124, on the first detector bank 120, is exposed by the retraction of the first single photon source shield 128, and in which the second single photon source 130, on the second detector bank 122, is exposed by the retraction of the second single photon source shield 130. In this arrangement, a photon 138 emitted by the first single photon source 124 is detected at a detector element 140 on the second detector bank 122.
- a count in a given sinogram bin corresponding to a single LOR may correspond to different detector pairs, because when the detector array is rotating, the LOR between two detectors at time 0 would later be covered by another detector pair.
- detector efficiency estimates are obtained by conducting a blank scan acquisition when the rotatable detector array is held stationery. In the case that no LORs are combined (i.e. no spanning), there is a one-to-many correspondence between the (non-zero) bins in the sinogram and the detectors.
- the measurement model is:
- the set Z. contains all sinogram bins that correspond to the detector j, i.e.
- Equation (10) is the Maximum Likelihood solution for the efficiencies when the measurement model (9) applies.
- the sum in (10) over all sinogram bins that correspond to the same detector reduces the noise in the estimated efficiencies.
- detector efficiency estimates are generated when the detector array is rotating.
- the detector efficiencies may depend on temperature, and hence rotation speed, thus detector efficiency estimates may be generated which more accurately reflect the operative efficiencies.
- one single photon source is exposed to the (rotating) detector array, and the other single photon source is shielded, as shown in Figure 6.
- a measurement model that can be used has the same form as equation (9) above, albeit with different A' z values, where there again is a one-to-many correspondence between j z and z.
- a prescription similar to (10) can also be used.
- the measurement is then to be repeated with the other source viewable such that the other detector bank is irradiated.
- both single photon sources are exposed in the (rotating) detector array.
- At least some of the sinogram bin values will correspond to either two or four overlapping but different source-detector arrangements.
- the scanner hardware After a full rotation, the scanner hardware will add counts to the same set of gates again, such that N different sinograms, each with uniquely identifiable correspondence between the LOR and detector, are produced.
- a measurement model, and analysis function, such as that described above using a single exposed source may then be employed to generated detector efficiencies.
- the present invention may be implemented to generate detector efficiencies from a blank scan that could then be used as a daily check on the performance of a PET scanner.
- the invention may also be used to provide detector efficiency factors for a scanner normalisation procedure. Since detector efficiencies change over time, it is desirable to be able to ascertain detector efficiency factors on a regular basis and without the need for complex procedures.
- the above-mentioned efficiency estimation algorithms, or alternatives thereto, may be used in such a procedure.
- the above embodiments are to be understood as illustrative examples of the invention.
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Application Number | Priority Date | Filing Date | Title |
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AT04735921T ATE545875T1 (en) | 2003-06-04 | 2004-06-03 | METHOD FOR ESTIMATING DETECTOR EFFICIENCY FOR A PET SCANNER |
US10/559,145 US7564035B2 (en) | 2003-06-04 | 2004-06-03 | Generating detector efficiency estimates for a pet scanner |
EP04735921A EP1631844B8 (en) | 2003-06-04 | 2004-06-03 | Generating detector efficiency estimates for a pet scanner |
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GBGB0312776.8A GB0312776D0 (en) | 2003-06-04 | 2003-06-04 | Generating detector efficiency estimates for a pet scanner |
GB0312776.8 | 2003-06-04 |
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EP (1) | EP1631844B8 (en) |
AT (1) | ATE545875T1 (en) |
GB (1) | GB0312776D0 (en) |
WO (1) | WO2004109332A1 (en) |
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JP2006524328A (en) * | 2003-04-22 | 2006-10-26 | コーニンクレッカ フィリップス エレクトロニクス エヌ ヴィ | Attenuation map creation from PET scan |
US10143437B2 (en) * | 2015-06-30 | 2018-12-04 | General Electric Company | Systems and methods for dynamic scanning with multi-head camera |
US10481285B1 (en) | 2018-08-13 | 2019-11-19 | General Electric Company | Systems and methods for determination of depth of interaction |
US10976452B2 (en) | 2018-08-13 | 2021-04-13 | General Electric Medical Systems Israel, Ltd. (Il) | Systems and methods for improved medical imaging |
US10591619B2 (en) | 2018-08-15 | 2020-03-17 | GE Precision Healthcare LLC | Anodes for improved detection of non-collected adjacent signals |
US10247834B1 (en) | 2018-08-15 | 2019-04-02 | General Electric Company | Anodes for improved detection of non-collected adjacent signal |
CN110215227B (en) * | 2019-06-05 | 2022-10-14 | 上海联影医疗科技股份有限公司 | Time window setting method and device, computer equipment and storage medium |
US11092701B1 (en) | 2020-07-07 | 2021-08-17 | GE Precision Healthcare LLC | Systems and methods for improved medical imaging |
US11320545B2 (en) | 2020-07-07 | 2022-05-03 | GE Precision Healthcare LLC | Systems and methods for improved medical imaging |
CN113069138B (en) * | 2021-03-23 | 2023-06-30 | 上海联影医疗科技股份有限公司 | Positron emission tomography device, coincidence efficiency detection method and normalization method |
KR20230064665A (en) * | 2021-11-03 | 2023-05-11 | (주) 인비즈 | Pet customized helping rehabilitation equipment manufacturing apparatus and method using 3d printing technic |
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US6040580A (en) * | 1993-03-26 | 2000-03-21 | Cti Pet Systems, Inc. | Method and apparatus for forming multi-dimensional attenuation correction data in tomography applications |
US6008493A (en) * | 1997-05-30 | 1999-12-28 | Adac Laboratories | Method and apparatus for performing correction of emission contamination and deadtime loss in a medical imaging system |
US6201247B1 (en) * | 1998-04-02 | 2001-03-13 | Picker International, Inc. | Line source for gamma camera |
US6100531A (en) * | 1998-04-17 | 2000-08-08 | Adac Laboratories | Dual-purpose radiation transmission source for nuclear medicine imaging system |
US6198104B1 (en) * | 1998-10-23 | 2001-03-06 | Adac Laboratories | Randoms correction using artificial trigger pulses in a gamma camera system |
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- 2003-06-04 GB GBGB0312776.8A patent/GB0312776D0/en not_active Ceased
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2004
- 2004-06-03 US US10/559,145 patent/US7564035B2/en not_active Expired - Fee Related
- 2004-06-03 AT AT04735921T patent/ATE545875T1/en active
- 2004-06-03 EP EP04735921A patent/EP1631844B8/en not_active Expired - Lifetime
- 2004-06-03 WO PCT/GB2004/002340 patent/WO2004109332A1/en active Search and Examination
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EP1631844B1 (en) | 2012-02-15 |
US7564035B2 (en) | 2009-07-21 |
EP1631844A1 (en) | 2006-03-08 |
GB0312776D0 (en) | 2003-07-09 |
EP1631844B8 (en) | 2012-03-21 |
US20060249682A1 (en) | 2006-11-09 |
ATE545875T1 (en) | 2012-03-15 |
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