CN102755172B - Nuclear medical imaging method and device - Google Patents
Nuclear medical imaging method and device Download PDFInfo
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- CN102755172B CN102755172B CN201210129209.9A CN201210129209A CN102755172B CN 102755172 B CN102755172 B CN 102755172B CN 201210129209 A CN201210129209 A CN 201210129209A CN 102755172 B CN102755172 B CN 102755172B
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Abstract
The invention provides a nuclear medical imaging method and device for obtaining optimal images. The nuclear medical imaging method comprises a first determining step of determining a line of response set by the positions of a pair detector crystals, a defining step of defining an array of radiation points corresponding to the line of response, a second determining step of determining a solid angle with the bottom surface being surfaces of the pair of detector crystals defining the line of response for each point in the array of the radiation points corresponding to the line of response, a generating step of averaging the solid angle for generating an average solid angle, a third determining step of determining a position factor related to depth-of-interaction, and a calculation step of multiplying the reciprocal of the average solid angle by the position factor for calculating a geometric corrective factor for the line of response.
Description
The application advocates the interests of the priority of the Japanese patent application No. 2012-85541 of the U.S. Patent Application No. application on April 4th, 13/096,672 and 2012 applied on April 28th, 2011, and quotes the full content of above-mentioned patent application in this application.
Technical field
Embodiments of the present invention relate to nuclear medicine (imaging) method and nuclear medical imaging apparatus.
Background technology
PET (positron emission tomography) (Positron Emission Tomography:PET) is 1 branch of nuclear medicine, is imported by the radiopharmaceuticals of positron emission in the body of subject.When radiopharmaceuticals decays, generate positron.Specifically, in the phenomenon be known with electron reaction and as the paired annihilation events (Positron annihilation) of positron respectively by multiple positron, generate along while counting line while roughly contrary direction movement, there is right γ (gamma) photon.The γ photon detected in gate time at the same time carry out record to usually used as paired annihilation events by PET scanner (scanner).In flight time (Time Of Flight:TOF) imaging, also measure the time in the counting interval (interval) while right each γ photon occurs simultaneously in detection.Flight-time information represent the event detected, position simultaneously on counting line.In order to the image of the subject or checking matter of rebuilding or generate scanning (scan) object, use the data (data) of multiple paired annihilation events.
Fig. 1 is the figure of an example of the geometric configuration representing PET imaging device.In FIG, illustrate in the three-centered detector (3D detector) had at PET imaging device, the positron radiated with measure while the cross-section coordinate (transaxial coordinate) of counting line (Line Of Response:LOR) and axial coordinate (axial coordinate).Cross-section coordinate is such as the coordinate set in the axial cross section of body perpendicular to the direction of principal axis of scanner, subject, and axial coordinate is such as the coordinate set in the axial cross section of body along the direction of principal axis of scanner, subject.Coordinate (x
e, y
e, z
e) or (s
e, t
e, z
e) represent the image coordinate of positron " e " radiated.In addition, coordinate (x
a, y
a, z
a) represent the position gamma-ray detector crystal " a " of the side determined as LOR being detected, in addition, coordinate (x
b, y
b, z
b) represent the position gamma-ray detector crystal " b " of the opposing party determined as LOR being detected.The projection coordinate of the LOR determined in non-TOF, can by " (s,
z, θ), in this case z=(z
a+ z
b)/2 " represent.
corresponding with the slope of the line segment ab in cross-section coordinate, " θ " is corresponding with the slope of the line segment ab in axial coordinate.Or the projection coordinate of the LOR determined also can comprise the additional dimension " t " of TOF-LOR." t " is the value corresponding with the time, such as, and " t
a" detect that the time of γ photon is corresponding, " t with " a " in gate time at the same time
b" detect that the time of γ photon is corresponding, " t with " b " in gate time at the same time
e" correspond at the same time in gate time " a " detect that the time of γ photon and " b " detect the time difference of the time of γ photon.By using the information of " t ", can determine " z " of the projection coordinate of determined LOR.In the PET imaging device of these types (non-TOF, TOF), except existing except deviation in the detector efficiency of each crystal, the detection efficiency of scanner entirety is decided by geometric key element, and then, depend on the incident angle of the solid angle formed in the zone, the distance put from detector crystal to radiation and the crystal to LOR.
The angle of the two dimension in the solid angle three dimensions that to be object formed relative to " certain point ".Mathematically, the solid angle Ω formed for bottom surface with face S is represented by following formula (1).
At this, " n subscript (hat) " of formula (1) is the unit vector (vector) from " certain point ", " da " of formula (1) is the elementary area of face portion, and " r " of formula (1) is the distance from initial point to face portion.Solid angle is the size of the size of the outward appearance of the object represented when observing from " certain point ".The solid angle of object and the fan-shaped area equation of the unit sphere (centered by angular vertex) limited by object.This definition can be applied in the arbitrary dimension comprising one dimension and two dimension.Fig. 2 A is the figure of the concept representing two-dimentional solid angle.Fig. 2 A represents the concept that the solid angle used in the detection of event occurs while PET.Solid angle shown in unit circle represents that the crystal measured from point " p " is to the solid angle of (i, j), or LOR
ijsolid angle.That is, in the PET scanner shown in Fig. 1, the solid angle of LOR depends on (s
e, t
e, z
e) position of radiation point that represents.(s
e, t
e) coordinate represents the position of the radiation point in cross section (the left figure of Fig. 1), z
ecoordinate represents the identical position on the direction of principal axis of scanner.Towards the visual field end of cross section, namely " | s| (absolute value of s) " larger, the solid angle of LOR becomes larger.
If 2 gamma-rays from radial location get to 2 crystal on LOR respectively, then decided the angle of incidence of LOR by 2 gamma-ray angle of incidence.Fig. 2 B is the figure representing gamma-ray angle of incidence.As shown in Figure 2 B, the gamma-ray each angle of incidence radiated from positron radiation can be described by " α " as polar angle (polar angle) and " β " as azimuth (azimuthal angle).Any one angle in " α " and " β " all depends on the relation of the normal of the plane of crystal of gamma-ray heading and gamma-rays incidence.Become large along with angle " α " or " β ", gamma-ray incident angle also becomes large.To be determined by gamma-ray incident angle the gamma-ray transit dose of directional crystal, that is, the position (Depth-Of-Interaction:DOI) of interactional depth direction.As a result, the normal of 1 crystal of LOR with " | s| " larger or " θ " larger, angle " α " or " β " become larger mode to be changed, and therefore, the impact of DOI changes.This situation tilts to carry out characterization by LOR further.
In addition, geometric key element is also subject to the impact of the crystal positions of the LOR in detector block (block).The crystal distance detector block end of LOR is nearer, and 1 side of detector block applies the impact larger than the crystal of the LOR of the front surface (front face) of detector block to solid angle and angle of incidence.
Geometric correction factor for revising the initial data measured also can be decided by the data obtaining the high plane of counting or the line of rotation.At first, the difference of initial data to radiogenic geometric arrangement, decay and each crystal efficiency about counting is revised.Then, the section (profile) along the radial direction of " s " shown in Fig. 1 is generated, as ring (ring) poor (z to each cross section
b-z
a) function.Further, about the section of these radial directions, get its inverse, and directly apply as geometric correction factor.
Fig. 3 is the figure of the photon radiation line source of the wire representing plane photon radiation line source in non-TOF-PET imaging device and rotation.In assay method in the past, in order to measure correction factor, as shown in the left figure of Fig. 3, plane radiographic source is placed in central authorities, or, as shown in the right figure of Fig. 3, the wire radiographic source through scanning is rotated.By making whole detectors to being exposed to plane radiographic source, collect correction factor data.Further, according to the relative count value about certain specific LOR, about whole LOR determine average while count value ratio calculate correction factor.But, in the radiogenic some assay methods of wire through scanning using plane radiographic source, rotation, there is some restrictions.Fig. 4 is the figure of the photon radiation line source of the wire representing multiple plane photon radiation line source in TOF-PET imaging device and multiple rotation.1st, in order to carry out the normalization of TOF-PET scanner, when employing the determination techniques for catching various radial location, as shown in the left figure of Fig. 4, plane radiographic source needs to be placed in multiple horizontal level.Or, in order to carry out the normalization of TOF-PET scanner, as shown in the right figure of Fig. 4, need wire radiographic source is rotated with various radius.Multiplely obtain scanning (scan) if used, then produce complicated normalization scanning step, multiple wire or multiple plane radiogenic configuration, multiple data obtain and much longer sweep time etc. for TOF-PET scanner normalization and overstate several shortcomings (demerit) large.2nd, generally performing the normalization scanning of high counting in order to reduce noise (noise), in final image, also predicting the effect of unfavorable (minus).Noise profile in normalization data is uneven, that is, have large ring difference (z
b-z
a) LOR in, there is higher noise.In addition, before generation correction factor, need other the correction such as the efficiency of crystal used in a detector.Extra noise may be increased by these corrections.In addition, existing PET rebuild in, in order to maintain the effectiveness of Poisson model (Poisson model), in logic preferred system (system) response in comprise correction factor.Along with in the method in the past of statistical noise, all carry out the reconstruction departing from Poisson model.Therefore, perhaps the PET image that result obtains is not optimum from the viewpoint of logic.
Prior art document
Non-patent literature
Non-patent literature 1:R.D.BADAWI, et al. " DEVELOPMENTS IN COMPONENT-BASED NORMALIZATION FOR 3D PET ", Phys.Med.Biol.44 (1999) 571-594, Printed in the UK.
Summary of the invention
The problem to be solved in the present invention is, provides a kind of nuclear medicine method and the nuclear medical imaging apparatus that can obtain optimum image.
The nuclear medicine method of embodiment comprises the 1st deciding step (step), definition step, the 2nd deciding step, generation step, the 3rd deciding step and calculation procedure.In the 1st deciding step, determine the pair of detectors crystal position separately of the detector had by nuclear medical imaging apparatus specify while counting line.In definition step, define with determined by above-mentioned 1st deciding step above-mentioned while radiation point corresponding to counting line array (array).In the 2nd deciding step, to each point in the array of the above-mentioned radiation point corresponding with counting line while of above-mentioned, determine the solid angle formed as bottom surface on the surface of the above-mentioned pair of detectors crystal of counting line while of above-mentioned for regulation.In generation step, the solid angle determined is averaged, generates mean solid angle by above-mentioned 2nd deciding step.In the 3rd deciding step, determine to depend on gamma-rays in the above-mentioned pair of detectors crystal of above-mentioned nuclear medical imaging apparatus through, relevant to the position of interactional depth direction position parameter.In calculation procedure, to the inverse of above-mentioned mean solid angle be multiplied by by above-mentioned 3rd deciding step determine above-mentioned position parameter, thus calculate for by above-mentioned 1st deciding step decision above-mentioned while counting line geometric correction factor.According to above-mentioned method, optimum image can be obtained.
Accompanying drawing explanation
Fig. 1 is the figure of an example of the geometric configuration representing PET imaging device.
Fig. 2 A is the figure of the concept representing two-dimentional solid angle.
Fig. 2 B is the figure representing gamma-ray angle of incidence.
Fig. 3 is the figure of the photon radiation line source of the wire representing plane photon radiation line source in non-TOF-PET imaging device and rotation.
Fig. 4 is the figure of the photon radiation line source of the wire representing multiple plane photon radiation line source in TOF-PET imaging device and multiple rotation.
Fig. 5 is the figure of the method representing the geometric correction factor calculating non-TOF-PET.
Fig. 6 is the figure of the lattice array represented in nuclear medical imaging apparatus.
Fig. 7 A is the figure (1) representing the concept calculating solid angle about nuclear medical imaging apparatus.
Fig. 7 B is the figure (2) representing the concept calculating solid angle about nuclear medical imaging apparatus.
Fig. 7 C is the figure (2) representing the concept calculating solid angle about nuclear medical imaging apparatus.
Fig. 8 is the figure (1) of the calculating of the position parameter representing interactional depth direction.
Fig. 9 is the figure (2) of the calculating of the position parameter representing interactional depth direction.
Figure 10 is the figure of the exemplary method representing the geometric correction factor calculating TOF-PET.
Figure 11 is the figure of the lattice array represented in TOF-PET imaging device.
Figure 12 is the figure of the example representing gamma-rays detection system (system).
Symbol description
100,120: scintillation crystal; 105,125: the array be made up of scintillation crystal; 110,135,140,195: photomultiplier tube; 145: display; 150: data acquisition unit; 170CPU; 175: interface; 180: electronic storage device
Detailed description of the invention
In a form of embodiment, nuclear medicine method comprises: (1) the 1st deciding step, determine the LOR specified by the pair of detectors crystal position separately of the detector had as the PET imaging device of nuclear medical imaging apparatus, (2) definition step, define the array of the radiation point corresponding with the LOR of above-mentioned decision, (3) the 2nd deciding step, to each point in the array of the above-mentioned radiation point corresponding with above-mentioned LOR, determine the solid angle formed as bottom surface on the surface of the above-mentioned pair of detectors crystal of the above-mentioned LOR of regulation, (4) generation step, the solid angle of above-mentioned decision is averaged, generate mean solid angle, (5) the 3rd deciding step, determine depend on gamma-rays in the above-mentioned pair of detectors crystal of above-mentioned imaging device through, as the DOI coefficient of " position parameter relevant to the position of interactional depth direction ", (6) calculation procedure, the inverse of above-mentioned mean solid angle is multiplied by the DOI coefficient of above-mentioned decision, thus calculate the geometric correction factor (geometric corrective factor) of the LOR for above-mentioned decision.In a form of another embodiment, if the readable storage medium of the computer of nonvolatile (computer) encode by order and performed by computer, then make computer execution said method.In a form of another embodiment, nuclear medical imaging apparatus comprises the memorizer being configured to the order that storage can be performed by computer and the computer being configured to perform to perform said method the above-mentioned order that can be performed by computer.
Below, with reference to accompanying drawing, describe the embodiment of nuclear medicine method in detail.
(corrections of non-TOF-PET data)
By before process of reconstruction (process) image data generating, need the initial data using a series of geometric correction factor correction by non-TOF-PET detector maturation.Because the quantity of detector and photomultiplier tube is very many, the deviation of the deviation of gain (gain) of photomultiplier tube, the deviation of physical property of detector and the detection efficiency of detector crystal, in initial data, produce heterogencity.In order to revise such scrambling, about the correction factor to the initial data applicating geometric counted from gamma-ray photon.
Solid angle for radiation point is larger, and its sensitivity detected becomes better.In order to be formed evenly initial data, the counting caused from the LOR becoming highly sensitive LOR can be reduced, the counting caused from the LOR of the LOR becoming muting sensitivity can be increased.
The technological progress of present embodiment is, provide revise along " the s direction " and " t direction " shown in Fig. 1 deviation, geometric correction factor after analytical Calculation.The method is not based on mensuration, but based on analytical Calculation.Geometric counting is based on solid angle and angle of incidence.Therefore, in the exemplary embodiment of this analytical Calculation, geometric correction factor " n
g" as shown in following formula (2), be broken down into 2 sub-factors.In addition, " the n of formula (2)
s" inversely proportional with the solid angle that the surface being used for detecting the paired crystal that event occurs is formed as bottom surface simultaneously, " the n of formula (2)
d" represent the impact of DOI.
n
g≡n
s×n
d …(2)
" n
s" and " n
d" be counted as separate." n
s" by the LOR specified with " j " by 2 detector crystal " i "
ijand the position of radiation point specifies, " n
d" by the crystal in the geometric configuration of scintillation material, detector block, detector block relative position and decide about the tiltangleθ on the regarding crystal surface of LOR.
Fig. 5 is the figure of the method representing the geometric correction factor calculating non-TOF-PET.Fig. 5 represents the correction factor " n of computational geometry
g" exemplary method.In the step S501 shown in Fig. 5, each LOR
ijby coordinate shown in Fig. 1 (s,
z, θ) represent.Following described Fig. 6 also represents whole LOR represented in step S501
ijin 1 LOR
ij.While step S501 specifies with the pair of detectors crystal position separately of the detector that decision is had by nuclear medical imaging apparatus, the 1st deciding step of counting line is corresponding.
In the step S502 shown in Fig. 5, for specifically
evaluate whole LOR
ij.In step S502, cyclically repeat each
under evaluation.In step S502, in order to reduce number of repetition, lower the complexity calculated, " n can be utilized
s" and " n
d" "
symmetry ".Thus, in the crystal of the quantity of the whole transverse direction in block detector, " n
s" and " n
d" identical.That is, in a pair crystal of the whole transverse direction in block detector, " n
s" and " n
d" identical.
In the step S503 shown in Fig. 5, for specifically
along dimension " t " and dimension " s ", simulation (simulate) fixed fire source array (point source array).That is, in step S503, for specifically
the change size of " t " or the size of " s " carry out simulation points radiographic source array.Step S503 and definition and while being determined by the 1st deciding step the step of the array of the radiation point that counting line is corresponding corresponding.The array of radiation point is the radiogenic array of point as radiating gamma-ray point, is that hypothesis is configured with the radiogenic array of multiple points after all.Fig. 6 represents the example of this fixed fire source array.Fig. 6 is the figure of the lattice array represented in nuclear medical imaging apparatus.In figure 6, the square 601 of black determines LOR
ijcrystal (i and j define crystal to).Dotted line in Fig. 6 is to distinguish LOR
ijthe mode of Region Of Interest (Region Of Interest:ROI) illustrate.Filled circles and open circles represent radiation point, namely annihilation position.Open circles is LOR
ijrOI in the point of array.Lattice array (fixed fire source array) also can extend to the below of s axle, or, also can utilize symmetry to reduce calculating number.In addition, as shown in Figure 6, it should be noted that array towards with
one changes.
In the step S504 shown in Fig. 5, to LOR
ijrOI in each some radiographic source, try to achieve solid angle.Further, in the step S504 shown in Fig. 5, for same or roughly the same
to whole LOR
ijcalculate mean solid angle.Step S504 with to while radiation point corresponding to counting line array in each point, determine that the 2nd deciding step of the solid angle formed as bottom surface on the surface of the regulation pair of detectors crystal of counting line is simultaneously corresponding.In addition, step S504 with the solid angle determined by the 2nd deciding step is averaged, generate mean solid angle generation step correspondence.The point calculating solid angle is the point of the open circles shown in Fig. 6.N
sthe LOR in the embodiment of non-TOF-PET scanner
ijrOI in the inverse of mean solid angle of point of whole open circles.Owing to arranging response by photoelectricity effect, therefore, for LOR
ijrOI in whole point calculate n
s.To LOR
ijrOI in each point, according to a position and regulation LOR
ijthe front surface of 2 crystal, carry out the calculating of solid angle.When 1 crystal is positioned at the border of detector, in the computational process of solid angle, consider boundary effect.That is, if crystal is positioned at border, then sometimes solid angle is increased by the side of crystal.By LOR
ijrOI in the value of solid angle of whole points synthetically average, the inverse of its meansigma methods is used as n
s.About the calculating of solid angle, below, be described in more detail.
Fig. 7 A, Fig. 7 B and Fig. 7 C are the figure representing the concept calculating solid angle about nuclear medical imaging apparatus.Fig. 7 A represents the crystal pair for just right, can as where tried to achieve the figure of solid angle.In fig. 7, show at the front surface of crystal parallel with y direction, the just right crystal centering that the boundary face of crystal is parallel with x direction, when from radiation point " p " to " D > > W " that the distance " D " in the x direction of crystal front surface is more much larger than the thickness " W " in the y direction of this crystal, solid angle δ roughly tries to achieve the situation into " W/D ".Fig. 7 B represents the crystal pair for tilting, can as where tried to achieve the figure of solid angle.In figure 7b, show the crystal pair that the front surface of crystal and the relative y direction of boundary face and x direction are tilted, when distance from radiation point " p " to " by the intersection point of crystal front surface and boundary face; the cross section orthogonal with xy plane " " D ' " compares " D ' > > W ' " from the length of radiating point " p " and observe the y direction in this cross section the range of observation of crystal " W ' " is much larger, solid angle δ roughly tries to achieve the situation into " W '/D ' ".The analytical Calculation of W ' and D ' need to consider the border that the crystal in space is right and towards, very complicated.Fig. 7 C shows for roughly proportional relative to the ratio of the quantity of whole lines (solid line+dotted line) with the quantity of the solid line situation of the simultaneous solid angle δ that the crystal tilted is right.
γ photon must move " certain distance " that statistics determines by detector material and photon energy (energy) in detector.The displacement of this restriction is interpreted as " through ", be " n to this through relevant sensitivity
d".Fig. 8 and Fig. 9 is the figure of the calculating of the position parameter representing interactional depth direction.Therefore, in the step S505 shown in Fig. 5, with through LOR
ijmode, together with the tiltangleθ of LOR, use parallel rays, calculate DOI impact and try to achieve n
d.Step S505 and determine depend on gamma-rays in the pair of detectors crystal of nuclear medical imaging apparatus through " position parameter (the n relevant with the position of interactional depth direction
d) " the 3rd deciding step corresponding.In step S505, as shown in Figure 8, in order to use angle of inclination to be " as LOR
ijthe θ at angle of inclination " " quite thin parallel rays (that is, the adjacent wire spacing in parallel rays is such as the level (level) of below 1mm) ", calculate parallel rays and 2 crystal intersection point separately, and be set to through LOR
ij.Each intersection point use simple Physical Attenuation model (model) and with sensitivity n
dbe associated.As a result, n
dcalculating based on the intersection point of whole rays and detector crystal.To every bar ray of light, length between the intersection point of the not only crystal of the LOR of calculating object, also calculates the whole displacement in other crystalline material.As shown in Figure 9, the crystal definition of 2 Dark greys is for calculating n
dlOR
ij.Arrow shown in Fig. 9 represents the γ photon of 2 511keV.Arrive the crystal of object at photon before, photon have passed some distances in other crystal.Further, n
dbe calculated as " e
-μ L1(1-e
-μ L0) e
-μ l1(1-e
-μ l0) ".At this, μ represents the γ photon absorbing 511keV, and to the linear attenuation coefficient that each flicker (scintillation) crystal provides.In addition, as shown in Figure 9, " L
1" be arrive the displacement of detector to the crystal of the object of arrival one side, " L from photon
0" be the distance of photon movement in the crystal of the object of a side.In addition, as shown in Figure 9, " L
1" be arrive the displacement of detector to the crystal of the object of arrival the opposing party, " L from photon
0" be the distance of photon movement in the crystal of the object of the opposing party.
In the step S506 shown in Fig. 5, by n
swith n
dbe multiplied, then obtain geometric correction factor n
g.Step S506 with the position parameter determined by the 3rd deciding step is multiplied by the inverse of mean solid angle, thus to calculate for corresponding by the calculation procedure of geometric correction factor of counting line while the 1st deciding step decision.As illustrated in step S502, re-define step, the 2nd deciding step, generation step, the 3rd deciding step and calculation procedure for counting line while of multiple respectively.And, to the data (initial data) that the scanning of the object by employing nuclear medical imaging apparatus (non-time-of-flight type PET device) obtains in the photography of reality, the correction factor of applicating geometric, thus data (initial data) are normalized.Specifically, in order to revise initial data, geometric correction factor n
gget inverse, and by defining LOR
ijdetector crystal i and detector crystal j detect counting be multiplied.Technological progress according to the present embodiment, solid angle becomes larger, geometric correction factor n
gbecome less, and then, by geometric correction factor n
gthe quantity counted when being applied to initial data diminishes.Therefore, technological progress according to the present embodiment, the analytic method of the correction factor of the counting of the LOR of the counting providing a kind of calculating fully can reduce highly sensitive LOR, increase muting sensitivity.Thus, in the present embodiment, use non-TOF-PET device, optimum image can be obtained.
Above-mentioned geometric correction factor is based on the non-TOF-PET device (at this, whole LOR is the LOR of the detector of same ring) of two dimension.But above-mentioned geometric correction factor also dimensionally can expand range of application (also can be set to the LOR of the inclination of the combination using different rings).So-called two dimension refers to the situation of the PET scanner of single ring.Even single ring use, detector crystal and geometric configuration are also three-dimensional.
When designing PET device, determine geometric correction factor.Therefore, comprise at unconfined embodiment and only calculate once geometric correction factor in advance, and during normalized after can being stored in, Computer can access the method for the geometric correction factor of (access).
Analytic method described above also can combine with other the physical model such as positron scope, intracrystalline scattering.Substantially, this method is the method for execution point as the calculating of distribution function (point spread function:PSF).Therefore, it is possible to PSF this method enrolled in reconstruction calculates.
(corrections of TOF-PET data)
The technological progress of present embodiment also can be applied to TOF-PET data.In non-TOF-PET scanner, in the past, in normalization step, the deviation of the solid angle for " t " was ignored.That is, suppose that solid angle does not change with the size one of " t ".But when radiating point and changing with the size one of " t ", the solid angle for radiation point also changes.Especially, for TOF-PET scanner, need to revise the deviation for the solid angle of " t ", otherwise the image that result obtains will be uneven.
N
dcalculating identical with the situation of non-TOF-PET.In TOF-PET, solid angle n
sby calculating with non-TOF-PET diverse ways.
Figure 10 represents to calculate the geometric correction factor (n of TOF-PET
g) the figure of exemplary method.Method shown in the method for Figure 10 and Fig. 5 is similar, below, only explains its difference.
In the step S1001 shown in Figure 10, each LOR
ijby coordinate shown in Fig. 1 (s,
z, θ) represent.Step S1001 is corresponding with the 1st above-mentioned deciding step.
In step 1002, for specifically
evaluate whole LOR
ij.In order to reduce the complexity of calculating, n can be utilized
swith n
d's
symmetry.In step S1002, in order to reduce number of repetition and reduce the complexity calculated, " n can be utilized
s" and " n
d" "
symmetry ".
In the step S1003 shown in Figure 10, for specifically
along dimension " t " (and dimension " s "), simulation points radiographic source array (point source array).That is, in step S503, for specifically
the change size of " t ", the size of " s " carry out simulation points radiographic source array.Step S1003 is corresponding with above-mentioned definition step.In addition, the array of radiation point is same as described above, corresponding with the array radiating gamma-ray point.Array towards as shown in Figure 6, with
one changes.
In the step S1004 shown in Figure 10, use the timing resolution of TOF detector by LOR
ijrOI and fixed fire source array, along LOR
ijtowards being divided into different groups (group) (subgroup (subgroup)).In step S1004, by LOR
ijrOI in some radiographic source be divided into subgroup, to be defined as flight-time information.Figure 11 is the figure of the lattice array represented in TOF-PET imaging device.Figure 11 represents and LOR
ijfixed fire source array corresponding to ROI in the example of subgroup of point.These subgroups are 1101,1102,1103 and 1104.
Further, in the step S1004 shown in Figure 10, try to achieve and LOR
ijfixed fire source array corresponding to ROI in the solid angle of each point.Further, in step S1004, for LOR
ijrOI in the radiogenic subgroup of point try to achieve solid angle respectively, and calculate the mean solid angle of each subgroup.Further, in S1004, about same or roughly the same
to whole LOR
ij, calculate mean solid angle according to the radiogenic each subgroup of point.That is, step S1004 comprises above-mentioned the 2nd deciding step (determining the some solid angle separately in the array of specific LOR); By with while radiation point corresponding to counting line array stripe become by multiple segmentation step of radiating the subgroup a little formed; And to each subgroup, the solid angle that the point in subgroup determines respectively is averaged, generate the generation step of mean solid angle.In addition, as in the step S1004 of segmentation step, according to the timing resolution of time-of-flight type PET device and by with while radiation point corresponding to counting line array stripe become and radiate by multiple the subgroup a little formed.That is, in segmentation step, for the whole LOR by the detector crystal i slope identical with the LOR that detector crystal j can be formed, according to the accuracy of detection of differential time of flight, the Range-partition of the position that each LOR can estimate radiation point is become multiple subgroup.The point calculating solid angle is the point of the open circles of Figure 11.In step S1004, by subgroup separately in solid angle be averaged, calculate multiple n
s.That is, in step S1004,1 n is calculated to each subgroup
svalue.
In the step 1005 shown in Figure 10, to obtain n
sfor the purpose of, in order to calculate the impact of DOI, with through LOR
ijmode use 1 parallel rays of the tiltangleθ with LOR.In step 1005, with through LOR
ijmode, use parallel rays together with the tiltangleθ of LOR, calculate DOI impact and try to achieve n
d.Step S1005 is corresponding with the 3rd above-mentioned deciding step.
In the step 1006 shown in Figure 10, by n
swith n
dbe multiplied, obtain LOR
ijthe geometric correction factor n of the subgroup of interior correspondence
g.Step 1006 be multiplied by position parameter for by the multiple inverse of each subgroup to mean solid angle a little formed that radiate, thus calculating for while counting line geometric correction factor calculation procedure correspondence.Illustrated by step S1002, re-define step, the 2nd deciding step, generation step, segmentation step, the 3rd deciding step and calculation procedure respectively for counting line while of multiple.Further, in the normalized in time-of-flight type PET (TOF-PET) device, to the scanning of the object by employing TOF-PET device and the geometric correction factor of market demand obtained, thus data are normalized.At this, in normalized, for each paired annihilation events that data comprise, select corresponding geometric correction factor according to counting line while the timing of this paired annihilation events and this annihilation events.That is, in normalized, in order to be specified to the subgroup belonging to pair annihilation event, the timing of this paired annihilation events is used.First, in order to carry out the correction of enumeration data, the TOF information " Δ t " of use case calculates the radiation point position of 1 event.Then, use this position, be used for 1 n of this TOF event from the some middle taking-up of subgroup 1101,1102,1103 and 1104
s.
2 look-up tables are used in the both sides of the embodiment of above-mentioned non-TOF-PET and TOF-PET.That is, for counting line while of each, the table of the solid angle of each of each point of the array comprising radiation point is generated.In other words, about whole
for each fixed fire source array each point each, by each LOR
ijsolid angle be stored into the 1st look-up table (Lookup table) (Tab-I).Its result, in Tab-I, 1 LOR
ijthere are the multiple solid angles changed with the size one of " t ".In addition, each LOR
ijn
dbe stored in the 2nd table (Tab-1I).N
ddo not depend on fixed fire source array.In non-TOF-PET, by 1 LOR
ij(have coordinate (s,
z, θ)) use together with Tab-I, as with this LOR
ijcorresponding solid angle, obtains all possible solid angle, is averaged and gets inverse, obtains ns.In addition, similarly, by LOR
ijuse together with Tab-II and obtain n
d.By n
swith n
dbe multiplied, obtain for the geometric normalized n of non-TOF
g.
On the other hand, in TOF-PET, pass through LOR
ij1 the paired annihilation events detected except coordinate (s,
z, θ) outside, also comprise TOF information (Δ t).Further, contain the size of whole " t ", obtained and LOR by Tab-I
ijcorresponding whole solid angle.Δ t, for calculating the position of radiation point t, thus, judges which subgroup (shown in Figure 11) is corresponding with this paired annihilation events.If it is determined that subgroup, then the solid angle that subgroup comprises be averaged and get inverse, obtain n
s.LOR
ijn
dobtained from Tab-II by the method identical with non-TOF-PET.Further, by n
swith n
dbe multiplied and obtain for the geometric normalized n of TOF-PET
g.Non-TOF-PET and TOF-PET respective in, to the look-up table of both sides or geometric correction factor be stored in the memorizer of PET device with the data that counting line is associated simultaneously, or be stored into the computer-readable storage medium (hard disk (hard disk), nonvolatile memory (memory), volatile memory, CD-ROM, DVD etc.) of nonvolatile.
(hardware (hardware))
Figure 12 represents that the exemplary hardware that jointly can use with the technological progress of present embodiment is formed.Figure 12 is the figure of the example representing gamma-rays detection system.In fig. 12, on fiber waveguide (light guide) 130, be configured with photomultiplier tube (Photomultiplier Tube:PMT) 135 and photomultiplier tube 140, under fiber waveguide 130, be configured with the array 105 be made up of scintillation crystal.The array 125 be made up of the 2nd scintillation crystal is opposed with scintillation crystal 105, overlaps in photomultiplier tube 195,110 with fiber waveguide 115.
In fig. 12, if radiate gamma-rays from subject (not shown), then gamma-rays advances on the rightabout approximately differing 180 °.Detect gamma-rays by scintillation crystal 100 and 120 roughly simultaneously.Further, when detecting gamma-rays by scintillation crystal 100 and 120 in the binding hours specified, scintillation event is determined.Like this, gamma-rays timing detection system detects a pair gamma-rays by scintillation crystal 100 and 120 roughly simultaneously.But, in order to simplify, at this, gamma-ray detection by scintillation crystal 100 is described.Self-evident for a person skilled in the art, about the explanation of scintillation crystal 100, be equally applicable to the gamma-ray detection by scintillation crystal 120.
Each photomultiplier tube 110,135,140 and 195 is connected with data acquisition unit 150 respectively.Data acquisition unit 150 comprises to process the hardware that the mode from the signal of photomultiplier tube is formed.Data acquisition unit 150 measures the gamma-ray time of advent.Data acquisition unit 150 generates two output (is the combination use of PMT135/140, and is the combination use of the T110/195) time of the differentiation pulse relative to system clock (not shown) being carried out encoding.In time-of-flight type PET system, transacter 150 is typically with the precision rise time labelling (time stamp) of 15 ~ 20 skins (pico) second.Data acquisition unit 150 measures the size of the signal (4 signals to data acquisition unit 150) of each PMT.
The output of data acquisition unit 150 is sent to CPU170 to carry out processing.This process comprises according to the output presumption energy of data acquisition unit 150, the process of position and the process estimating the time of advent for each paired annihilation events according to exported time mark, in order to improve the precision of energy, position and the presumed value of time, also can comprise according to correction in advance, apply a large amount of correction steps.Self-evident for a person skilled in the art, CPU170 can be installed as the individual logical doors such as application-specific IC (Application Specific Integrated Circuit:ASIC), field programmable gate array (Field-Programmable Gate Array:FPGA) or CPLD (Complex Programmable Logic Device:CPLD).The installation of FPGA or CPLD can describe language by VHDL (VHSIC Hardware Description Language:VHSIC hardware description language), Verilog or other hardware and carry out (code) change of encoding, in addition, also this coding directly can be stored in the electronic memory in FPGA or CPLD, or be stored as other electronic memory.In addition, electronic memory also can be non-volatile memorizer such as ROM, EPROM (Electrically programmable read only memory: EPROM), EEPROM (Electrically Erasable Programmable Read Only Memory: EEPROM), flash memory (flash memory).In addition, electronic memory also can be the memorizer of the volatibility such as static (static) or dynamic (dynamic) RAM.In addition, not only in order to carry out the dialogue between FPGA or CPLD and electronic memory, also in order to managing electronic memorizer, the blood processor of microcontroller (microcontroller) or microprocessor (microprocessor) etc. also can be set.
Or CPU170 also can as a series of computer-readable order for being stored in the both sides of storage medium that above-mentioned electronic memory and hard disk, CD, DVD, USB flash disk (flash drive) etc. both known or the some of a side.In addition, computer-readable order is as application program (utility application), the element of background process (background daemon) or operating system (operating system), or their combination provides, the Xeon processor (registered trade mark) manufactured with American I ntel company, or the blood processor such as the Opteron processor (registered trade mark) that AMD of the U.S. manufactures, and Microsoft VISTA (registered trade mark), UNIX (registered trade mark), Solaris (registered trade mark), the operating systems known by those skilled in the art such as LINUX (registered trade mark) and Apple MAC-OS (registered trade mark) perform in linkage.
Once by CPU170 process, then the signal after process carries out storage to electronic storage device 180 and the both sides to the display of display 145 or a side.Self-evident for a person skilled in the art, electronic storage device 180 also can be that hard disk drive, CD-ROM drive, DVD driver, USB flash disk, RAM, ROM etc. are at the known electronic storage device of this technical field.Display (display) 145 also can be installed as liquid crystal display (Liquid Crystal Display:LCD), cathode ray tube display (Cathod ray tube:CRT), plasma (plasma) display, Organic Light Emitting Diode (Organic Light Emitting Diode:OLED), Light-Emitting Diode (LED) etc. at the known display of this technical field.Like this, the record of electronic storage device 180 described herein and display 145 is only example, never for limiting the scope of the progress of present embodiment.
In addition, Figure 12 comprises the interface (interface) 175 gamma-rays detection system being connected to other external device (ED) and the both sides of user (user) or a side.Such as, interface 175 can be USB (universal serial bus) (Universal Serial Bus:USB) interface, PCMCIA (personal computer memory card international association) (the Personal Computer Memory Card International Association:PCMCIA) interface known by the art such as interface, Ethernet (Ethernet) (registered trade mark) interface.In addition, interface 175 both can be wired or wireless, also can comprise the man-machine interface (human interface) that the user for both sides or a side etc. with keyboard (keyboard) and mouse (mouse) engages in the dialogue.
Above, as described, according to the present embodiment, optimum image can be obtained.
Although the description of several embodiment of the present invention, but these embodiments are pointed out as an example, is not intended to limit scope of the present invention.These embodiments can be implemented with other various forms, in the scope of main idea not departing from invention, can carry out various omissions, displacement, change.These embodiments or its distortion are contained in scope of invention and main idea, and are contained in the invention of claims record and the scope of equalization thereof.
Claims (16)
1. a nuclear medicine method, is characterized in that, comprises:
1st deciding step, determine the pair of detectors crystal position separately of the detector had by nuclear medical imaging apparatus specify while counting line;
Definition step, define with determined by above-mentioned 1st deciding step above-mentioned while radiation point corresponding to counting line array;
2nd deciding step, to each point in the array of the above-mentioned radiation point corresponding with counting line while of above-mentioned, determines the solid angle formed as bottom surface on the surface of the above-mentioned pair of detectors crystal of counting line while of above-mentioned for regulation;
Generation step, is averaged the solid angle determined by above-mentioned 2nd deciding step, generates mean solid angle;
3rd deciding step, determine to depend on gamma-rays in the above-mentioned pair of detectors crystal of above-mentioned nuclear medical imaging apparatus through, relevant to the position of interactional depth direction position parameter;
Calculation procedure, by the inverse of above-mentioned mean solid angle is multiplied by by above-mentioned 3rd deciding step determine above-mentioned position parameter, calculate for by above-mentioned 1st deciding step determine above-mentioned while counting line geometric correction factor; And
The step of above-mentioned definition step, above-mentioned 2nd deciding step, above-mentioned generation step, above-mentioned 3rd deciding step and above-mentioned calculation procedure is repeated for each of counting line while of multiple.
2. nuclear medicine method according to claim 1, is characterized in that,
Also comprise normalization step, to the scanning of the object by employing above-mentioned nuclear medical imaging apparatus and the data obtained, applying the above-mentioned geometric correction factor calculated by above-mentioned calculation procedure, thus above-mentioned data are normalized.
3. nuclear medicine method according to claim 1, is characterized in that,
The array of above-mentioned radiation point with radiate gamma-ray corresponding.
4. nuclear medicine method according to claim 1, is characterized in that,
Above-mentioned nuclear medical imaging apparatus is non-time-of-flight type PET device.
5. a nuclear medicine method, is characterized in that, comprises:
1st deciding step, determine the pair of detectors crystal position separately of the detector had by nuclear medical imaging apparatus specify while counting line;
Definition step, define with determined by above-mentioned 1st deciding step above-mentioned while radiation point corresponding to counting line array;
2nd deciding step, to each point in the array of the above-mentioned radiation point corresponding with counting line while of above-mentioned, determines the solid angle formed as bottom surface on the surface of the above-mentioned pair of detectors crystal of counting line while of above-mentioned for regulation;
Segmentation step, becomes the array stripe of the above-mentioned radiation point corresponding with counting line while of above-mentioned and radiates by multiple the subgroup a little formed;
Generation step, to each subgroup be made up of multiple radiation point be partitioned into by above-mentioned segmentation step, is averaged the solid angle that each point in subgroup determines, generates mean solid angle;
3rd deciding step, determine to depend on gamma-rays in the above-mentioned pair of detectors crystal of above-mentioned nuclear medical imaging apparatus through, relevant to the position of interactional depth direction position parameter;
Calculation procedure, about each subgroup be made up of above-mentioned multiple radiation point, to the inverse of above-mentioned mean solid angle be multiplied by by above-mentioned 3rd deciding step determine above-mentioned position parameter, thus calculate for by above-mentioned 1st deciding step decision above-mentioned while counting line geometric correction factor; And
The step of above-mentioned definition step, above-mentioned 2nd deciding step, above-mentioned segmentation step, above-mentioned generation step, above-mentioned 3rd deciding step and above-mentioned calculation procedure is repeated for each of counting line while of multiple.
6. nuclear medicine method according to claim 5, is characterized in that,
Also comprise normalization step, to the scanning of the object by employing above-mentioned nuclear medical imaging apparatus and the data obtained, applying the above-mentioned geometric correction factor calculated by above-mentioned calculation procedure, thus above-mentioned data are normalized.
7. nuclear medicine method according to claim 6, is characterized in that,
Above-mentioned normalization step comprises following steps: each paired annihilation events comprised above-mentioned data, selects corresponding geometric correction factor according to counting line while the timing of this paired annihilation events and this annihilation events.
8. nuclear medicine method according to claim 7, is characterized in that,
Also comprising determining step, in order to determine the subgroup belonging to above-mentioned paired annihilation events, using the timing of this paired annihilation events.
9. nuclear medicine method according to claim 5, is characterized in that,
The array of above-mentioned radiation point with radiate gamma-ray corresponding.
10. nuclear medicine method according to claim 5, is characterized in that,
Above-mentioned nuclear medical imaging apparatus is time-of-flight type PET device.
11. nuclear medicine methods according to claim 10, is characterized in that,
The array stripe of the above-mentioned radiation point corresponding with counting line while of above-mentioned, according to the timing resolution of above-mentioned time-of-flight type PET device, becomes and radiates by multiple the subgroup a little formed by above-mentioned segmentation step.
12. nuclear medicine methods according to claim 5, is characterized in that,
Also comprise table generation step, to counting line while of each, generate the table of the solid angle of each point of each point comprising the array of above-mentioned radiation point.
13. 1 kinds of nuclear medical imaging apparatus, is characterized in that,
Possess the processor being configured to perform the order that computer can perform, for:
Counting line while determining to be specified by pair of detectors crystal position separately,
Define the array of the radiation point corresponding with counting line while of above-mentioned,
To each point in the array of the above-mentioned radiation point corresponding with counting line while of above-mentioned, determine the solid angle formed as bottom surface on the surface of the above-mentioned pair of detectors crystal of counting line while of above-mentioned for regulation,
The solid angle of above-mentioned decision is averaged and generates mean solid angle,
Determine to depend on gamma-rays in above-mentioned pair of detectors crystal through, relevant to the position of interactional depth direction position parameter,
By being multiplied by above-mentioned position parameter to the inverse of above-mentioned mean solid angle, calculate the geometric correction factor for counting line while above-mentioned determined.
14. nuclear medical imaging apparatus according to claim 13, is characterized in that,
Also possess memorizer, store above-mentioned geometric correction factor explicitly and above-mentioned determined while counting line.
15. 1 kinds of nuclear medical imaging apparatus, is characterized in that,
Possess the processor being configured to perform the order that computer can perform, for:
Counting line while determining to be specified by pair of detectors crystal position separately,
Define the array of the radiation point corresponding with counting line while of above-mentioned,
To each point in the array of the above-mentioned radiation point corresponding with counting line while of above-mentioned, determine the solid angle formed as bottom surface on the surface of the above-mentioned pair of detectors crystal of counting line while of above-mentioned for regulation,
The array stripe of the above-mentioned radiation point corresponding with counting line while of above-mentioned is become and radiates by multiple the subgroup a little formed,
To each subgroup be made up of above-mentioned multiple radiation point, the solid angle that each point in subgroup determines is averaged, generates mean solid angle,
Determine to depend on gamma-rays in above-mentioned pair of detectors crystal through, relevant to the position of interactional depth direction position parameter,
About each subgroup be made up of above-mentioned multiple radiation point, above-mentioned position parameter is multiplied by the inverse of above-mentioned mean solid angle, thus calculates the geometric correction factor for counting line while of above-mentioned.
16. nuclear medical imaging apparatus according to claim 15, is characterized in that,
Also possess memorizer, store above-mentioned geometric correction factor explicitly and above-mentioned determined while counting line.
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| JP2012085541A JP6054050B2 (en) | 2011-04-28 | 2012-04-04 | Nuclear medicine imaging method, nuclear medicine imaging apparatus and storage medium |
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| CN105769229B (en) * | 2014-12-24 | 2018-08-03 | 沈阳东软医疗系统有限公司 | A kind of method, image rebuilding method and device calculating PET system increasing radius value |
| CN105212957B (en) * | 2015-08-25 | 2018-01-16 | 浙江大学 | A kind of crystal level PET system time modification method based on TV Merge |
| CN105496436B (en) * | 2015-11-28 | 2018-06-19 | 上海联影医疗科技有限公司 | For the time-correcting method and device of PET device |
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| CN107970037B (en) * | 2016-10-24 | 2021-01-05 | 上海东软医疗科技有限公司 | Imaging method and imaging system |
| JP2019007825A (en) * | 2017-06-23 | 2019-01-17 | 株式会社島津製作所 | Nuclear medicine diagnostic device and routine maintenance and checkup method of nuclear medicine diagnostic device |
| CN109934278B (en) * | 2019-03-06 | 2023-06-27 | 宁夏医科大学 | High-dimensionality feature selection method for information gain mixed neighborhood rough set |
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