PET Imaging and Quantification

Suleman Surtisurti@mail.med.upenn.edu

(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