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Radiologic Physics: Nuclear Medicine
PET Imaging and Quantification
Suleman Surti
[email protected]
(215) 662-7214
Positron decay
vi) Two 511 keV photons
produced by e+ eannihilation
i) Unstable parent nucleus
~180˚
iii) Positron travels
short distance in tissue
(Neutrino escapes)
ii) Proton decays to neutron
Emits positron and neutrino
11C
t1/2 = 20 minutes
13N t
1/2 = 10 minutes
15O
t1/2 = 2 minutes
18F
t1/2 = 110 minutes
Positron Emission Tomography
A primary goal and usefulness of a tomographic imaging modality such as
PET is to achieve images where the intensity of each voxel in the image is
proportional to the activity concentration present in the corresponding location
in the patient
True, Scatter, Random coincidences
in PET
Trues ~ 2 . A
A = Activity
 = stopping power
Scatter
Scatters ~ k . Trues
k ~ energy threshold
(depends on energy resolution)
True
Random
Randoms ~ 2 . ( . A) 2
2 = coincidence timing window
(depends on decay time/light)
Count-rate Performance
70-cm long
phantom (20-cm
diameter)
NEMA NU2-2001
Noise
Equivalent
Count-rate
NEC =
T/(1+S/T+R/T)
Philips Gemini TF
Univ. of Pennsylvania
Limits on spatial resolution
1. Positron range, R:
18F
11C
82Rb
Rmax (mm)
2.6
3.8
16.5
FWHMp (mm)
0.22
0.28
2.6
2. Photon non-collinearity:
R
~180˚
FWHMNC=0.0022 X scanner diameter
(2-mm for a 90-cm diameter)
3. Detector resolution (FWHMd )
FWHM sys  FWHM 2p  FWHM 2NC  FWHM 2d
PET Instrumentation Design
• Scintillators
stopping power, speed, light output
• Detector configuration
scintillator - photo-sensor coupling
• Scanner geometry
field-of-view (axial)
2-dimensional vs. 3-dimensional
Time-of-flight PET
• Data processing / image reconstruction
scatter, randoms and attenuation correction
iterative reconstruction algorithms
Comparison of Scintillators
350
300
250
Decay time (ns)
200
Light output (% NaI)
150
Stopping power
(100*1/cm)
100
50
0
NaI
Scintillator
 (ns)
 (cm-1)
E/E (%)
Rel. light output (%)
BGO
NaI(Tl)
230
0.35
6.6
100
BGO
300
0.95
10.2
15
GSO
GSO
60
0.70
8.5
25
LSO
LSO
40
0.86
10.0
70
LuAP
18
0.95
~15
30
LPS
30
0.70
~10
73
LaBr
35
0.47
2.9
150
Scintillation Detector
To pre-amplifier
Photo-Multiplier
Tube (PMT)
Vacuum
9
7
Dynode
stages
5
3
1
10
8
6
4
2
Photoelectron paths
Light collection region
Semi-transparent photocathode
Light from
scintillator
CONDUCTION BAND (empty)
e-
ACTIVATOR STATES
ENERGY
GAP (Eg)
Light photon
VALENCE BAND (full)
Scintillator
Small crystals require position
encoding
Block Detector
CTI HR+ (1995)
BGO
8 x 8 array
4 x 4 x 30 mm3
19 mm PMTs (4)
18,432 crystal elements (32 rings)
1,152 PMTs
Block vs. Quadrant Sharing
Standard Block
(Casey-Nutt)
Quadrant Sharing Block
(W.-H. Wong)
Similar spatial resolution with larger PMTs
or
Better spatial resolution with similar size PMTs
Continuous optical coupling
More uniform light output -> better energy resolution
Similar spatial resolution with larger PMTs
Example: Philips Allegro (2001) 17,864 crystal elements (GSO)
420 PMTs
2D (septa) vs. 3D (no septa)
2D Imaging
3D Imaging
Low Scatter fraction ~ 10%
High Scatter fraction ~ 30%
Low geometric sensitivity
High geometric sensitivity
2D
3D
Axial Slice
Axial Slice
Energy threshold reduces scatter &
random coincidences- particularly in 3D
Scatter/True=k
Scatter
Scatter/True>k
True
NEC Count-rates - 2D vs. 3D
GE Advance
70-cm long phantom NEMA 2001
S.Kohlmeyer and T. Lewellen
University of Washington
160
2D: 2001
140
NEC (kcps)
120
100
80
3D: 2001 (380 keV)
(300 keV)
60
40
NECR-3D:380 keV '01
NECR-3D:300 keV '01
20
NECR-2D:300 keV '01
0
0
0.2
0.14 Ci/cc
0.4
0.6
0.8
1
Activity Concentration
1.2
1.4
1.6
1.8
High count-rate capability in 3D PET requires fast,
dense scintillator with good energy resolution
LSO  = 40 ns
BGO  = 300 ns
0.81/cm
0.91/cm
Compare 3 CTI scanners: LSO Accel, BGO EXACT, BGO HR+ (2D)
70000
3D Accel
NEC (cps)
60000
50000
2D HR+
NEC
40000
ACCEL
30000
• Both measurements
assume randoms
smoothing
BGO EXACT
HR+ 2D
20000
10000
5.6 mCi at start of scan
0
0.000
0.100
0.200
0.300
0.400
0.500
0.600
Specific Activity
3D EXACT
0.700
0.800
0.900
• Courtesy of CTI, inc
1.000
Time-of-flight PET
x
• Can localize source along line
of flight - depends on timing
resolution of detectors
t1
D
t = uncertainty in measurement
of t1-t2
x = uncertainty in position along
LOR
= c . t/2
D/x ~ reduction in variance
or gain in sensitivity
Gain in sensitivity over a
non-TOF scanner
• Time of flight information
t2 reduces noise in images weighted back-projection along
LOR
14
12
10
8
6
4
2
0
200
D=40 cm
D=30 cm
D=20 cm
300
400
500
600
Timing resolution (ps)
700
Does Noise-Equivalent Count-rate
(NEC) infer Image Quality?
NEC = Trues / (1 + Scatter/Trues + Randoms/Trues)
NEC1/2 ~ Signal / Noise
NEC includes global effects
• Trues
• Noise from scatter and randoms
NEC does not include local effects
• Spatial resolution - variations within FOV
• Image reconstruction
• Accuracy of scatter and randoms correction
• Attenuation correction
• Deadtime corrections and normalization
Fully 3D
Iterative
Reconstruction
improves
image quality
Philips Allegro
Filtered
Backprojection
3D
Ramla
Positron Emission Tomography
What is needed to achieve quantitative PET images?
1.
2.
3.
4.
5.
Deadtime correction
Data Normalization
Scatter correction
Randoms correction
Attenuation correction
Deadtime correction
• Deadtime — High count-rate effect present in
radiation detectors
• Two manifestations:
• Pulse pileup — Events are collected
but measurements such as energy and
spatial position are affected (reduced
image quality)
• Loss of counts — Due to electronics
deadtime and determined mainly by
scintillator decay time
• Loss of counts corrected by measuring collected
counts vs activity in a uniform cylinder
Data normalization
• Normalization — non-uniformities in event
detection over the full scanner
• Two sources:
• Variation in amount of scintillation light
collection due to crystal nonuniformities and detector design
(detector effect)
• Difference in detection sensitivity due
to angle of incidence
>d
d
Data normalization techniques
Rotating rod source
N i, j
Norm i, j 
N i, j
Ci, j
Norm
Ci, j 
Normi, j

Uniform cylinder
Scatter Correction (SSS)
CT
A
non TOF
• Contribution to LOR AB
from each scatter point —
Activity distribution and
Klein-Nishina equation
• Repeat for all LORs to get
scatter sinogram
B
TOF
P188
Randoms Correction — Delayed
window technique
Signal A
A
Delayed Signal A
Coincidence
Window, 
B
Signal B
time
Why do we need attenuation
correction?
• More accurate activity distribution
uniform liver, ‘cold lungs’
• Improved lesion detectability
deep lesions
• Reduce image artifacts and streaking
reconstruct using consistent data
• Improved image quality with iterative reconstruction
include attenuation into model
But…attenuation correction must be
FAST
- compared to emission scan
ACCURATE - e.g. near lung boundary
LOW NOISE - minimize noise propagation
Attenuation correction
can be calculated directly in PET
patient
d1
d2
PET: High energy photons with
small , but pair of photons must
traverse entire body width.
I/I0 = e -d1 e-d2 = e-(d1+d2)
Total path length,
D=d1+d2
D can be independently measured
and allows an accurate correction
(511kev) = 0.095/cm
I/I0= e-d1 e-d2 = 0.06
for D=30cm
Transmission sources for attenuation
measurements
1.
PET transmission source (68Ge/68Ga) - source of coincident annihilation photons
(mono energetic @ 511 keV), 265 day half life
2.
Single photon source (137Cs) - source of single -rays (mono energetic @ 662 keV),
20 yr half-life
3.
X-ray CT scan - source of X-rays with a distribution of energies from ~30 to 120 keV.
We can assume an ‘effective’ energy of ~ 75 keV
spectra
positron source -ray source
X-ray source
Intensity
I0(E)
0
30
120
E (keV)
511
662
(Recall that the PET emission data is attenuated at 511 keV)
Transmission Scan
d1
d2
d1 + d2 = D
Emission
I / I0 = e-d1 . e-d2
= e-D
Transmission
I / I0 = e-D
137Cs
point source
662 keV, t1/2 = 30 yr
Post-injection transmission scan
Philips Allegro
• 20 mCi 137Cs pt src
• 40 sec Tx
acquisition
• Energy scaling
• EC subtraction
• Segmentation
• Interleaved Em-Tx
7 Em frames
9 Tx frames
University of Pennsylvania PET Center
CT-based attenuation correction:
threshold method
STEP 1: Separate bone and soft tissue using threshold of 300 H.U.
0.5
2
linear attenuation/density (cm /g)
STEP 2: Scale to PET energy 511 keV.
Scale factors (511:~70 keV):
bone 0.41, soft tissue: 0.50
0.4
0.3
soft tissue / water
bone
0.2
0.1
0
0
100
200
300
400
energy (keV)
STEP 3: Forward project to obtain attenuation correction factors.
Kinahan PE, Townsend DW, Beyer T, et al. Med Phys. 1998; 25(10): 2046-2053.
500
Potential problems for CT-based
attenuation correction
• Difference in CT and PET respiratory patterns
Can lead to artifacts near the dome of the liver
• Use of contrast agent
Can cause incorrect values in PET image
• Truncation of CT image due to keeping arms down in the
field of view to match the PET scan
Can cause artifacts in corresponding regions in PET image
• Bias in the CT image due to beam-hardening and scatter
from the arms in the field of view
Attenuation correction for PET
Types of transmission images
Coincident photon Ge68/Ga-68
(511 keV)
Single photon Cs137
(662 keV)
X-ray
(~30-130 keV)
high noise
15-30 min scan time
low bias
low contrast
lower noise
5-10 min scan time
some bias
lower contrast
no noise
1 min scan time
potential for bias
high contrast
Alessio AM, Kinahan PE, Cheng PM, et al. Radiol. Clin. N. America 2004; 42(6): 1017-1032.
Attenuation correction - increased
confidence of liver lesion
No AC
University of Pennsylvania PET Center
AC
Philips Allegro
Attenuation correction - better comparison
of relative activity of deep (mediastinum)
vs. superficial (axilla) lesions
No AC
University of Pennsylvania PET Center
AC
Philips Allegro
Image quality degrades with heavy
patients
Slim 58 kg
“Normal” 89 kg
Increasing attenuation (less counts)
Increasing scatter (more noise)
Increasing volume (lower count density)
Heavy 127 kg
How can we improve image quality?
2D - counts limited by septa and maximum allowed dose
3D - counts limited by dead-time and randoms
Scintillator
High stopping power - higher coincidence fraction
Fast decay - lower dead-time and randoms
Energy resolution - lower scatter and randoms
Geometry
Sensitivity ~ (Axial FOV)2
(increased scintillator and PMT cost)
Time-of-flight
Requires very fast scintillator with excellent timing
resolution
Time-of-flight PET
x
• Can localize source along line
of flight - depends on timing
resolution of detectors
t1
D
t = uncertainty in measurement
of t1-t2
x = uncertainty in position along
LOR
= c . t/2
D/x ~ reduction in variance
or gain in sensitivity
Gain in sensitivity over a
non-TOF scanner
• Time of flight information
t2 reduces noise in images weighted back-projection along
LOR
14
12
10
8
6
4
2
0
200
D=40 cm
D=30 cm
D=20 cm
300
400
500
600
Timing resolution (ps)
700
Time-of-flight PET
PET scanner
70-cm bore
18-cm axial FOV
CT scanner
Brilliance 16-slice
Philips Gemini TF Univ. of Pennsylvania
PET shows increased FDG uptake in region of porta hepatis
CT demonstrates that this uptake corresponds to the gallbladder
representing acute cholecystitis, not bowel activity
Phantom measurements
non TOF
TOF
non TOF
3 min
5 min
1 min
3 min
4-to-1 contrast; IEC phantom
2.2 mCi in IEC, 5.4 mCi in line source cylinder
TOF
6-to-1 contrast; 35-cm diameter
7.0 mCi in all phantoms
Gemini TF
Heavy-weight patient study
13 mCi
2 hr post-inj
3 min/bed
Colon cancer
119 kg
BMI = 46.5
MIP
LDCT
non-TOF
TOF
Improvement in lesion detectability with TOF
Clinical 18F-FDG imaging
• Clinical 18F-FDG imaging essentially involves two
tasks:
• Identifying regions with abnormal uptake (lesion detection)
• Deriving a measure of glucose metabolism in these regions (lesion
estimation task)
Factors affecting lesion detection and
activity estimation
• Accuracy of scanner normalization and corrections
for deadtime, scatter, randoms, & attenuation
• Remove biases with minimal noise propagation
• Spatial resolution
• Lesion size and partial volume effects
• Lesion activity uptake relative to background
• Scan time
• Reduced noise
• Patient habitus
• Determines amount of Sc, R, and attenuation
• Reconstruction
• Determines amount of noise in image and for iterative algorithms
plays off contrast recovery with noise
Summary
• PET scanner design is still an evolving area of research with
new scintillators and photo-detectors being developed
• Current generation of clinical scanners achieve spatial
resolution of 4-5 mm
• Fully-3D imaging is imaging mode of choice
• PET is still count limited
• TOF PET can help improve the statistical quality of PET
images
• PET/CT as a multi-modality imaging device has increased
the confidence in interpreting PET images
• Future direction - PET/MRI scanners
18F-Fluoro-Deoxy-Glucose
OH
(FDG)
Ido et al. 1978
O
H
OH
OH
OH
18F
Glucose
Blood -> tissue -> cell
phosphorylation - glycogen
FDG
Blood -> tissue
phosphorylation
Patient injected activity: 10 mCi = 3.7 x 108 dps
Tracer kinetics:
6 pico-mole ~ 1 nano-gram
Lesion detectability
2 min
3 min
4 min
5 min
Non-TOF
TOF
• Improved lesion detectability with TOF achieved with short scan time
and reduced reconstruction time (# of iterations)
• Spheres are just barely visible with a 5 minute scan in non-TOF
• After a 2-3 minute scan in TOF the spheres become visible
6-to-1 contrast; 35-cm diam. cyl.; 10-mm diam. spheres
6.4mCi in all phantoms
Time-of-flight scanners need investigation
of new data processing and image
reconstruction methods
• Scatter correction
- can incorporate timing information
- energy based methods - statistical weighting
• Image reconstruction - list-mode ML-EM
- optimize use of TOF
- include data corrections in system model
- spatial recovery
• Data quantification - SUV estimation
- convergence of lesion contrast improves with TOF
• Image evaluation - lesion detectability measures
- how does TOF improve SNR in image?