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Radiologic Physics: Nuclear Medicine. PET Imaging and Quantification. Suleman Surti [email protected] (215) 662-7214. vi) Two 511 keV photons produced by e + e - annihilation ~180˚. i) Unstable parent nucleus. iii) Positron travels short distance in tissue (Neutrino escapes). - PowerPoint PPT Presentation
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PET Imaging and Quantification
Suleman [email protected]
(215) 662-7214
Radiologic Physics: Nuclear Medicine
i) Unstable parent nucleus
ii) Proton decays to neutron Emits positron and neutrino
vi) Two 511 keV photonsproduced by e+ e-
annihilation ~180˚
iii) Positron travelsshort distance in tissue(Neutrino escapes)
11C t1/2 = 20 minutes13N t1/2 = 10 minutes
15O t1/2 = 2 minutes18F t1/2 = 110 minutes
Positron decay
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
Positron Emission Tomography
True
Scatter
Random
Trues ~ 2 . AA = Activity = stopping power
Scatters ~ k . Truesk ~ energy threshold(depends on energy resolution)
Randoms ~ 2 . ( . A) 2
2 = coincidence timing window(depends on decay time/light)
True, Scatter, Random coincidences in PET
70-cm long phantom (20-cm diameter)
NEMA NU2-2001
Philips Gemini TFUniv. of
Pennsylvania
Noise Equivalent Count-rate
NEC = T/(1+S/T+R/T)
Count-rate Performance
3. Detector resolution (FWHMd )
Limits on spatial resolution
R
1. Positron range, R:
18F 11C 82Rb
Rmax (mm) 2.6 3.8 16.5
FWHMp (mm) 0.22 0.28 2.6
~180˚2. Photon non-collinearity:
FWHMNC=0.0022 X scanner diameter(2-mm for a 90-cm diameter)
€
FWHMsys = FWHMp2 + FWHMNC
2 + FWHMd2
• Scintillators stopping power, speed, light output
• Detector configuration scintillator - photo-sensor coupling
• Scanner geometryfield-of-view (axial)
2-dimensional vs. 3-dimensional Time-of-flight PET • Data processing / image reconstruction
scatter, randoms and attenuation correction iterative reconstruction algorithms
PET Instrumentation Design
0
50
100
150
200
250
300
350
NaI BGO GSO LSO
Decay time (ns)
Light output (% NaI)
Stopping power(100*1/cm)
Scintillator NaI(Tl) BGO GSO LSO LuAP LPS LaBr (n )s 230 300 60 40 18 30 35
μ (cm-1) 0.35 0.95 0.70 0.86 0.95 0.70 0.47Δ /E E (%) 6.6 10.2 8.5 10.0 ~15 ~10 2.9
Re .l ligh toutpu (t %) 100 15 25 70 30 73 150
Comparison of Scintillators
VALENCE BAND (full)
CONDUCTION BAND (empty)
e-
ACTIVATOR STATESENERGYGAP (Eg) Light photon
Scintillator
Photo-MultiplierTube (PMT)
Scintillation Detector
CTI HR+ (1995)
BGO8 x 8 array 4 x 4 x 30 mm3
19 mm PMTs (4)
Block Detector
18,432 crystal elements (32 rings)1,152 PMTs
Small crystals require position encoding
Similar spatial resolution with larger PMTsor
Better spatial resolution with similar size PMTs
Standard Block(Casey-Nutt)
Quadrant Sharing Block(W.-H. Wong)
Block vs. Quadrant Sharing
More uniform light output -> better energy resolutionSimilar spatial resolution with larger PMTs
Example: Philips Allegro (2001) 17,864 crystal elements (GSO)420 PMTs
Continuous optical coupling
2D Imaging 3D Imaging
Low Scatter fraction ~ 10% High Scatter fraction ~ 30%
Axial Slice Axial Slice
Low geometric sensitivity High geometric sensitivity
2D 3D
2D (septa) vs. 3D (no septa)
Energy threshold reduces scatter & random coincidences- particularly in 3D
0
5000
10000
100 300 500Energy (keV)
TrueScatter
Scatter/True=k
Scatter/True>k
0
20
40
60
80
100
120
140
160
0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8
Activity Concentration
NEC (kcps)
NECR-3D:380 keV '01
NECR-3D:300 keV '01
NECR-2D:300 keV '01
S.Kohlmeyer and T. LewellenUniversity of Washington
70-cm long phantom NEMA 2001GE Advance
0.14 μCi/cc
2D: 2001
3D: 2001 (380 keV) (300 keV)
NEC Count-rates - 2D vs. 3D
Compare 3 CTI scanners: LSO Accel, BGO EXACT, BGO HR+ (2D)
• Both measurements assume randoms smoothing
• Courtesy of CTI, inc
NEC (cps) 3D Accel
2D HR+
3D EXACT
LSO = 40 ns 0.81/cmBGO = 300 ns 0.91/cm
High count-rate capability in 3D PET requires fast, dense scintillator with good energy resolution
• Can localize source along line of flight - depends on timing resolution of detectors
• Time of flight information reduces noise in images - weighted back-projection along LOR
02468
101214
200 300 400 500 600 700
Timing resolution (ps)
Gain in sensitivity over a
non-TOF scanner
D=40 cm
D=30 cm
D=20 cm
Time-of-flight PET
Δx = uncertainty in position along LOR = c . Δt/2
Δt = uncertainty in measurement of t1-t2
D/Δx ~ reduction in variance or gain in sensitivity
D
Δx
t2t1
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
NEC = Trues / (1 + Scatter/Trues + Randoms/Trues)NEC1/2 ~ Signal / Noise
Does Noise-Equivalent Count-rate (NEC) infer Image Quality?
Fully 3D Iterative
Reconstruction improves
image quality
Filtered Backprojection
3DRamla
Philips Allegro
Positron Emission Tomography
What is needed to achieve quantitative PET images?
1. Deadtime correction2. Data Normalization3. Scatter correction4. Randoms correction5. 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 non-uniformities and detector design (detector effect)
• Difference in detection sensitivity due to angle of incidenced > d
Data normalization techniquesRotating rod source Uniform cylinder
€
Normi, j =N i, j
N i, j
€
Ci, jNorm =
Ci, j
Normi, j
TOF
P188
non TOF
Scatter Correction (SSS)
AA
BB
CT•Contribution to LOR ABfrom each scatter point
— Activity distribution andKlein-Nishina equation• Repeat for all LORs to
get scatter sinogram
Randoms Correction — Delayed window technique
AA
BB
Delayed Signal ADelayed Signal A
Signal BSignal B
Coincidence Coincidence Window, Window,
Signal ASignal A timetime
• More accurate activity distributionuniform liver, ‘cold lungs’
• Improved lesion detectabilitydeep lesions
• Reduce image artifacts and streakingreconstruct using consistent data
• Improved image quality with iterative reconstructioninclude attenuation into model
But…attenuation correction must be FAST - compared to emission scanACCURATE - e.g. near lung boundaryLOW NOISE - minimize noise propagation
Why do we need attenuation correction?
Total path length, D=d1+d2
D can be independently measured and allows an accurate correction
PET: High energy photons with small μ, but pair of photons must traverse entire body width.
d1
d2
patient
I/I0= e-μd1 e-μd2 = 0.06 for D=30cm
μ(511kev) = 0.095/cm
Attenuation correction can be calculated directly in PET
I/I0 = e -μd1 e-μd2 = e-μ(d1+d2)
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
E (keV)30 120 511 662
Intensity
I0(E)
X-ray source positron source -ray source
0
spectra
(Recall that the PET emission data is attenuated at 511 keV)
Transmission sources for attenuation measurements
137Cs point source662 keV, t1/2 = 30 yr
I / I0 = e-μd1 . e-μd2
= e-μD
I / I0 = e-μD
d1d2
d1 + d2 = D
Emission
Transmission
Transmission Scan
University of Pennsylvania PET Center
• 20 mCi 137Cs pt src• 40 sec Tx acquisition
• Energy scaling• EC subtraction• Segmentation
• Interleaved Em-Tx 7 Em frames 9 Tx frames
Philips Allegro
Post-injection transmission scan
CT-based attenuation correction: threshold method
0
0.1
0.2
0.3
0.4
0.5
0 100 200 300 400 500
energy (keV)
linear attenuation/density (cm
2 /g)
soft tissue / water
bone
Scale factors (511:~70 keV): bone 0.41, soft tissue: 0.50
STEP 1: Separate bone and soft tissue using threshold of 300 H.U.
STEP 3: Forward project to obtain attenuation correction factors.
STEP 2: Scale to PET energy 511 keV.
Kinahan PE, Townsend DW, Beyer T, et al. Med Phys. 1998; 25(10): 2046-2053.
Potential problems for CT-based attenuation correction
• Difference in CT and PET respiratory patternsCan lead to artifacts near the dome of the liver
• Use of contrast agentCan cause incorrect values in PET image
• Truncation of CT image due to keeping arms down in the field of view to match the PET scanCan 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
Types of transmission images
Coincident photon Ge-68/Ga-68(511 keV)
high noise15-30 min scan time
low biaslow contrast
Single photon Cs-137
(662 keV)
lower noise5-10 min scan time
some biaslower contrast
X-ray(~30-130 keV)
no noise1 min scan timepotential for bias
high contrast
Attenuation correction for PET
Alessio AM, Kinahan PE, Cheng PM, et al. Radiol. Clin. N. America 2004; 42(6): 1017-1032.
University of Pennsylvania PET Center
No AC AC
Philips Allegro
Attenuation correction - increased confidence of liver lesion
No AC AC
University of Pennsylvania PET Center Philips Allegro
Attenuation correction - better comparison of relative activity of deep (mediastinum)
vs. superficial (axilla) lesions
Slim 58 kg “Normal” 89 kg Heavy 127 kgIncreasing attenuation (less counts)
Increasing scatter (more noise)
Increasing volume (lower count density)
Image quality degrades with heavy patients
ScintillatorHigh stopping power - higher coincidence fractionFast decay - lower dead-time and randomsEnergy resolution - lower scatter and randoms
GeometrySensitivity ~ (Axial FOV)2 (increased scintillator and PMT cost)
Time-of-flight Requires very fast scintillator with excellent timing resolution
2D - counts limited by septa and maximum allowed dose3D - counts limited by dead-time and randoms
How can we improve image quality?
• Can localize source along line of flight - depends on timing resolution of detectors
• Time of flight information reduces noise in images - weighted back-projection along LOR
02468
101214
200 300 400 500 600 700
Timing resolution (ps)
Gain in sensitivity over a
non-TOF scanner
D=40 cm
D=30 cm
D=20 cm
Time-of-flight PET
Δx = uncertainty in position along LOR = c . Δt/2
Δt = uncertainty in measurement of t1-t2
D/Δx ~ reduction in variance or gain in sensitivity
D
Δx
t2t1
PET scanner70-cm bore18-cm axial FOV
CT scannerBrilliance 16-slice
PET shows increased FDG uptake in region of porta hepatisCT demonstrates that this uptake corresponds to the gallbladder representing acute cholecystitis, not bowel activity
Philips Gemini TF Univ. of Pennsylvania
Time-of-flight PET
1 min
non TOF TOF
3 min
4-to-1 contrast; IEC phantom2.2 mCi in IEC, 5.4 mCi in line source cylinder
3 min
non TOF TOF
5 min
6-to-1 contrast; 35-cm diameter7.0 mCi in all phantoms
Phantom measurements
Heavy-weight patient study13 mCi 2 hr post-inj3 min/bed
MIP
Colon cancer
119 kgBMI = 46.5
non-TOF
Gemini TF
Improvement in lesion detectability with TOF
TOF LDCT
• 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)
Clinical 18F-FDG imaging
• 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
Factors affecting lesion detection and activity estimation
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
OH
H
OH
OHOH
18F
O
Patient injected activity: 10 mCi = 3.7 x 108 dpsTracer kinetics: 6 pico-mole ~ 1 nano-gram
Ido et al. 1978
GlucoseBlood -> tissue -> cellphosphorylation - glycogen
FDGBlood -> tissuephosphorylation
18F-Fluoro-Deoxy-Glucose (FDG)
Lesion detectability
• 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
2 min 3 min 4 min 5 min
Non-TOF
TOF
6-to-1 contrast; 35-cm diam. cyl.; 10-mm diam. spheres 6.4mCi in all phantoms
• 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?
Time-of-flight scanners need investigation of new data processing and image
reconstruction methods