Download SPECT

Emission Tomography Principles
and Reconstruction
Professor Brian F Hutton
Institute of Nuclear Medicine
University College London
brian.hutton@uclh.nhs.uk
Outline
• imaging in nuclear medicine
• basic principles of SPECT
• basic principles of PET
• factors affecting emission tomography
History
• Anger camera 1958
• Positron counting, Brownell 1966
• Tomo reconstruction; Kuhl & Edwards 1968
• First rotating SPECT camera 1976
• PET: Ter-Pogossian, Phelps 1975
SPECT
Anger gamma camera
Detector:
400x500mm
Energy resn
Intrinsic resn
Detector
To Display &
Computer
X Y Z
Position/Energy Circuits
Photo Multiplier Tubes
~9mm thick
~10%
3-4mm
Radionuclides:
Collimator
Tc-99m
6hr
Designed140keV,
to suit energy
I -123 hole
159keV,
13hr
HR:
size 1.4mm
Ga-68 length
93-296keV,
33mm3.3dy
I-131 septa
360keV,
8dy
0.15mm
Capacitor
NaI (Tl)Crystal
Light
HV Supply
Anode
Collimator
Gamma Ray
Anode
e-
D8
D7
Output
Pulse
D8
D7
D6
D6
D5 D5
(D - Dynode)
D4
D4
D3 D3
e-
D2
D1
Cathode
D1
D2
eCathode
Light from
crystal
Organ-specific options
specialized collimators for standard cameras
parallel
fanbeam
conebeam
pinhole
slit-slat
crossed slit
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
SPECT
Reconstruction
sinogram for each
transaxial slice
Filtered back projection
1 angle
2 angles
4 angles
16 angles
128 angles
Organ-specific systems
specialised system designs, with use limited to a
specific application
Positron Annihilation
Isotope
Emax
(keV)
Max range FWHM
(mm)
(mm)
18F
663
2.6
0.22
11C
960
4.2
0.28
13N
1200
5.4
0.35
15O
1740
8.4
1.22
82Rb
3200
17.1
2.6
Coincidence Detection
detector 1
coincidence
window
detector 2
time (ns)
PET "Block" Detector
Scintillator
array
PMTs
C
BGO
(bismuth
germanate)
Images courtesy of CTI
A
Histogram
B
Attenuation Correction in PET
attenuation for
activity in body
N = N0 e -x. e - (D-x)
= N0 e -D
attenuation for
external source
N = N0 e -D
(D=body thickness)
(for 511 keV  ~ 0.096/cm
attenuation factors: 25-50)
Coincidence Lines of Response (LoR)
sinogram
fanbeam
parallel
PET
Reconstruction
sinogram
1 angle
2 angles
4 angles
• conventional filtered back projection
• iterative reconstruction
16 angles
128 angles
Understanding iterative reconstruction
Objective
Find 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)

X
object

Y2
Y
estimated
projection
Y1
X
System matrix
0
0
0
0
0
0
0
1
0
0
0
0
pixeli
0
0
0
1
0
0
0
0
0
0
1
0
voxelj
0
0
0
0
0
0
BP
patient
update
(x ratio)
original
projections
ML-EM
reconstruction
NO
original
CHANGE
estimate
FP
estimated
projections
current
estimate
Image courtesy of Bettinardi et al, Milan
Noise control
• stop at an early iteration
• use of smoothing between iterations
• post-reconstruction smoothing
• penalise ‘rough’ solutions (MAP)
• use correct and complete system model
Factors affecting quantification
courtesy Ben Tsui, John Hopkins
detector
+
without
attenuation
correction
transmission
with
attenuation
correction
System matrix: with attenuation
0
0
0
0
0
0
0
0.2
0
0
0
0
0
0
0
0.5
0
0
0
0

0
0
0.9
0
0
0
0
0
0
0
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
Modelling resolution
Gamma camera resolution
• depends on distance
SPECT resolution
• need radius of rotation
PET resolution
• position dependent
System matrix: including resolution model
0
0
0
0
0
0
0.1
0.2
0.1
0
0
0
0
0
0.2
0.5
0
0
0
0

0
0.3
0.9
0.3
0
0
0.2
0
0
0
PET resolution
detector
depth of interaction results in
asymmetric point spread function
positron range
colinearity
FWHMtotal2 = FWHMdet2 + FWHMrange2 + FWHM1802
radial int
radial ext
tangential
Modelling resolution
• potentially improves resolution
• requires many iterations
• slow to compute
• stabilises solution
• better noise properties
detector
(projection)

object
w/o resn model
Courtesy: Panin et al
IEEE Trans Med Imaging
2006; 25:907-921
with resn model
Can we consider measurements to be quantitative?
Scatter correction
• multiple energy windows for SPECT; PETCT
standard models
• SPECT local effects; PET more distributed
detector
object
Scatter fraction
• SPECT ~35%
PET 2D ~15%; 3D ~40%
Scatter
• influenced by photon
energy, source location,
scatter medium
• reduces contrast
Monte Carlo
measured
• scatter models
analytical, 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
3D reconstruction
Approaches
• rebin data followed by 2D reconstruction
single slice rebinning (SSRB)
multi-slice rebinning (MSRB)
Fourier rebinning (FORE)
• full 3D reconstruction
3D OSEM
3D RAMLA
limits for
FORE
2D 4min
FORE 2D-OSEM
28subsets 2 iter
3D 4min
2D 2min
3D 2min
VUE Point 3D-OSEM
FORE 2D-OSEM
28subsets 2iter
28subsets 5 iter
Courtesy V Bettinardi, M Gilardi, Milan
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