Professor Brian F Hutton Institute of Nuclear Medicine University College London
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Transcript of Professor Brian F Hutton Institute of Nuclear Medicine University College London
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Professor Brian F HuttonInstitute of Nuclear MedicineUniversity College London
Emission Tomography Principles and Reconstruction
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Outline
• imaging in nuclear medicine
• basic principles of SPECT
• basic principles of PET
• factors affecting emission tomography
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SPECT
History• Anger camera 1958• Positron counting, Brownell 1966• Tomo reconstruction; Kuhl & Edwards 1968• First rotating SPECT camera 1976• PET: Ter-Pogossian, Phelps 1975
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Gamma Ray
Light
Position/Energy CircuitsX Y Z
Collimator
NaI (Tl)Crystal
Photo Multiplier Tubes
Detector
To Display &Computer
D1
D3
D5
D7D8
D6
D4
D2
Anode
Cathode
Anode
D8
D7
D6
D5
D4
D3
D2
D1
Cathode
HV Supply
Capacitor
OutputPulse
Light fromcrystal
(D - Dynode)
e-
e-
e-
Anger gamma camera Detector: 400x500mm ~9mm thickEnergy resn ~10%Intrinsic resn 3-4mm
Radionuclides: Tc-99m 140keV, 6hr I -123 159keV, 13hr Ga-68 93-296keV, 3.3dy I-131 360keV, 8dy
CollimatorDesigned to suit energy HR: hole size 1.4mm
length 33mmsepta 0.15mm
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parallel fanbeam conebeam
pinhole slit-slat crossed slit
Organ-specific options specialized collimators for standard cameras
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Single Photon Emission Computed Tomography (SPECT)
Single Photon Emission Computed Tomography (SPECT)
• relatively low resolution; long acquisition time (movement)• noisy images due to random nature of radioactive decay• tracer remains in body for ~24hrs: radiation dose ~ standard x-ray• function rather than anatomy
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SPECT Reconstruction
1 angle 2 angles 4 angles 16 angles 128 angles
Filtered back projection
sinogram for each transaxial slice
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Organ-specific systemsspecialised system designs, with use limited to a specific application
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Positron Annihilation
Isotope Emax
(keV)Max range
(mm)FWHM(mm)
18F11C13N15O
82Rb
663 2.6 0.22
960 4.2 0.28
1200 5.4 0.35
1740 8.4 1.22
3200 17.1 2.6
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Coincidence Detection
detector 1
detector 2
coincidence window
time (ns)
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PET "Block" Detector
Scintillatorarray
PMTs
Histogram
A B
C
Images courtesy of CTI
BGO(bismuth
germanate)
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Nox
µ
D
e-µx e-µ(D-x)
No
µ
D
e-µ0 e-µD
Attenuation Correction in PET
attenuation foractivity in bodyN = N0 e -x. e - (D-x)
= N0 e -D
attenuation for external sourceN = N0 e -D
(D=body thickness)
(for 511 keV ~ 0.096/cmattenuation factors: 25-50)
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Coincidence Lines of Response (LoR)
parallelfanbeamsinogram
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PET Reconstruction
1 angle 2 angles 4 angles 16 angles 128 angles
• conventional filtered back projection• iterative reconstruction
sinogram
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Understanding iterative reconstruction
ObjectiveFind the activity distribution whose estimated projections match the measurements.
Modelling the system (system matrix)What is the probability that a photon emitted from location X will be detected at detector location Y.
- detector geometry, collimators- attenuation- scatter, randoms
detector(measurement)
object
estimated projection
X
Y
X
Y1
Y2
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0
0
0
0
0
0
1
0
0
0
0 0 0 0 0 0 0 1 0 0
0 0 0 0 0 1 0 0 0 0
System matrix
voxelj
pixeli
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ML-EM reconstruction
originalprojections
estimatedprojections
currentestimate
original estimate
update(x ratio)
FP
BP NOCHANGE
patient
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Image courtesy of Bettinardi et al, Milan
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• stop at an early iteration• use of smoothing between iterations• post-reconstruction smoothing• penalise ‘rough’ solutions (MAP)• use correct and complete system model
Noise control
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Factors affecting quantification
courtesy Ben Tsui, John Hopkins
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+ -
transmission
withoutattenuationcorrection
withattenuationcorrection
detector
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0
0
0
0
0
0
0.9
0
0
0
0 0 0 0 0 0 0 0.2 0 0
0 0 0 0 0 0.5 0 0 0 0
System matrix: with attenuation
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Partial volume effects
• effect of resolution and/or motion
• problems for both PET and SPECT
• similar approaches to correction
• scale of problem different due to resolution
• some different motion effects due to timing:
ring versus rotating planar detector
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Modelling resolution
Gamma camera resolution• depends on distance
SPECT resolution• need radius of rotation
PET resolution • position dependent
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0
0
0
0
0
0.3
0.9
0.3
0
0
0 0 0 0 0 0 0.1 0.2 0.1 0
0 0 0 0 0.2 0.5 0.2 0 0 0
System matrix: including resolution model
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FWHMtotal2 = FWHMdet
2 + FWHMrange2 + FWHM180
2
positron range colinearity
detector
PET resolution
depth of interaction results in asymmetric point spread function
tangential
radial int radial ext
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detector(projection)
object
Courtesy: Panin et al IEEE Trans Med Imaging 2006; 25:907-921
• potentially improves resolution• requires many iterations• slow to compute
Modelling resolution
w/o resn model
with resn model
• stabilises solution• better noise properties
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detector
object
Scatter correction
• multiple energy windows for SPECT; PETCT standard models
• SPECT local effects; PET more distributed
Can we consider measurements to be quantitative?
Scatter fraction
• SPECT ~35% PET 2D ~15%; 3D ~40%
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• scatter modelsanalytical, Monte Carlo, approximate models
• measurement triple energy window (TEW), multi-energy
subtract from projections:measured proj – TEW
or combine with projector in reconstruction:compare (forward proj + TEW) with measured proj
Scatter• influenced by photon energy, source location, scatter medium• reduces contrast
measured
Monte Carlo
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3D reconstructionApproaches• rebin data followed by 2D reconstruction
single slice rebinning (SSRB)multi-slice rebinning (MSRB)Fourier rebinning (FORE)
• full 3D reconstruction3D OSEM3D RAMLA
limits for FORE
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Courtesy V Bettinardi, M Gilardi, Milan
2D 4min 3D 4min 2D 2min 3D 2min
FORE 2D-OSEM 28subsets 5 iter
VUE Point 3D-OSEM28subsets 2iter
FORE 2D-OSEM 28subsets 2 iter
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Summary
Emission tomography• functional rather than anatomical• single photon versus dual photon (PET)• main difference is ‘collimation’
Iterative reconstruction• very similar approach for SPECT and PET• currently most popular is OSEM (or similar)• the better the system model the better the reconstruction