What do PET clinicians/researchers want? What can the PET ...

What do PET clinicians/researchers want? What can the PET ...

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

C t1/2 = 20 minutes 13 N t1/2 = 10 minutes 11 O t1/2 = 2 minutes 18 F t1/2 = 110 minutes 15 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

True Scatters ~ k . Trues k ~ energy threshold (depends on energy resolution) 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: F C Rb 18 11 82 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)) (c -1) E/E/EE/E (%) R l..l.igh oupu % () BGO NaI(Tl) 230

0.35 6.6 100 GSO BGO 300 0.95 10.2 15 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 Topre-amplifier

Photo-Multiplier Tube (PMT) Vacuum Dynode stages 9 10 7 8 5 6 3

4 1 2 Photoelectronpaths Semi-transparent photocathode Light collectionregion Light from scintillator CONDUCTION BAND (empty) e- ACTIVATOR STATES ENERGY GAP (Eg) Light photon

Scintillator VALENCE BAND (full) 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 10000 Scatter Scatter/True>k 5000 0 100

300 500 Energy (keV) 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 140 120 100 80

3D: 2001 (380 keV) (300 keV) 60(kcps) NEC 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 1.2 Activity Concentration 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 NEC (cps) 3D Accel 60000 50000 2D HR+ 40000 NEC 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 Time of flight information t2 reduces noise in images weighted back-projection along

LOR t1 D 14 12 = uncertainty in measurement 10 8 of t1-t2 uncertainty in position along 64 non-TOF scanner 2 LOR 0 sensitivity over a Gain in = c . t/2 200 300 400 500 x ~ reduction in variance

or gain in sensitivity D=40 cm D=30 cm D=20 cm 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 Normi, j = N i, j Ci, j Norm Ci, j = Normi, j

Uniform cylinder Scatter Correction (SSS) CT A non TOF B Contribution to LOR AB from each scatter point Activity distribution and Klein-Nishina equation Repeat for all LORs to get scatter sinogram 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 Butattenuation 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 g-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 g-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 + d 2 = D Emission I / I0 = e-d1 . e-d2 = e-D Transmission I / I0 = e-D Cs point source 662 keV, t1/2 = 30 yr 137

Post-injection transmission scan iversity of Pennsylvania PET Center 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 CT-based attenuation correction: threshold method STEP 1: Separate bone and soft tissue using threshold of 300 H.U. STEP 2: Scale to PET energy 511 keV. 2 0.5 /g)

0.4 Scale factors (511:~70 keV): bone 0.41, soft tissue: 0.50 0.3 soft tissue / water bone 0.2 0.1 linear 0 attenuation/density (cm 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 Time of flight information t2 reduces noise in images weighted back-projection along LOR t1 D 14 12 = uncertainty in measurement 10

8 of t1-t2 uncertainty in position along 64 non-TOF scanner 2 LOR 0 sensitivity over a Gain in = c . t/2 200 300 400 500 x ~ reduction in variance or gain in sensitivity D=40 cm D=30 cm D=20 cm 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 F-Fluoro-Deoxy-Glucose (FDG) 18 OH O Ido et al. 1978 H OH OH

OH 18 F 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?

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