1. Nuclear Medicine Systems: PET Instrumentation
BMME 560 & BME 590I Medical Imaging: X-ray, CT, and Nuclear Methods
Nuclear Medicine Systems: PET Instrumentation
Guest Lecturer:
Marijana Ivanovic
Office:
Radiology, 2112 Old Clinic
Tel:

2. Positron Emission Tomography (PET):
1973
1980
1994
2000

3. Positron combines with Electron and annihilates
Positron ( ß+) emission
511 keV X Z A Y
Z-1 A
+ b+ + n
Two anti-parallel 511 keV
Photons produced g P N
180 o - e + + + _ _ + + _ _ g e + P N
Unstable parent
nucleus
Positron combines with
Electron and annihilates
511 keV
Proton decays to
Neutron 
positron and
neutrino emitted n

4. PET SPECT Electronic collimation (x1,y1) (x2,y2) Collimated detector
Only photons emitted perpendicularly to
the detector plane can pass through collimator
(x1,y1)
(x2,y2)
Coordinates of a two simultaneously detected photons determine the line of photon emission (line of response – LOR)
Collimator hole determines
the line of photon emission
Collimated detector
Electronic collimation

5. Why positron emitters for tracers?
Many of the positron emitters occur naturally in biological molecules (C, N, O, etc.)
Many have small molecular weights relative to the biological molecules they may be used to label (e.g., 18F) even if they aren’t found there naturally.
Consequently, radioactive isotopes can be attached to biologically interesting molecules with no or minimal impact on the behavior of those molecules in the body.

6. Why Use 18F Labeled FDG? Free space Cell Blood
18FDG competes with glucose for transport and phosphorylation
18FDG is not further metabolized but is trapped with high efficiency within cell since dephosphorylation of fluorodeoxyglucose-6-phosphate is slow and only FDG may be transported across membranes
Free
space
Cell
Blood
Hexokinase
Glucogen
CO2 + H2O
Glucose
Glucose
Hexokinase
FDG-6-PO4
FDG
FDG

7. Selected PET tracer compounds
USE
18F-2-deoxyglucose
metabolic imaging
13NH3 (ammonia)
blood flow
11C-palmitate
myocardial metabolism
11C labeled amino acids
protein synthesis, pancreatic
imaging, tumor metabolism
11C-butanol
myocardial flow/perfusion
15O2 , H2 15O
blood flow
C 15O2
blood volume
82Rb-chloride
myocardial perfusion

8. Positron range
Positron loses its kinetic energy in collisions with atom of the surrounding matter and comes to the rest within few mm from origin in the body tissue ( in ~ 10-9 sec). g
annihilation
photon
The average distance a positron travels before annihilation is determined by the energy with which it leaves the nucleus.
electron/positron
annihilation e- e+
b+ energy dependent
range
annihilation
photon g P N
  F emits positrons which can travel up to 3 mm (1/10 inch) max.

9. Non-colinearity of the annihilation photons
Two photons are not quite
180° apart. Angular deviation
is about ± 0.25° due to
positron motion
If a positron annihilates with an electron before it loses all kinetic energy (not at rest)
Positron
annihilation + + + + + + + + + + _ - e + e e e + _ _ _ _ _ _ _ _ _ _ b +
180º ± 0.25º

10. Ideal PET Real PET True LOR LOR Detected LOR 180º ± 0.25º 180º
Finite
Positron range
180º ± 0.25º
Non-collinearity
Detected LOR
The finite positron range and the non-colinearity of the annihilation photons give rise to an inherent positional inaccuracy not present in SPECT. However, other characteristics of PET more than offset this disadvantage.

11. Three basic types of coincidence imaging systems…
Dual camera
Coincidence
“Low” cost
Dedicated PET
“High” end
Dedicated PET
Lower cost
Higher sensitivity and count rate

12. Types of Coincident Events
True Coincidence
Scattered Coincidence
Random Coincidence

13. True Coincidence Event
Line of response
(LOR)

14. Scattered Coincidence Event
Scattered events
recorded
simultaneously
True LOR if
unscattered
Recorded LOR

15. Random Coincidence Event
Unrelated events
recorded simultaneously
Incorrect LOR
The random rate for the a particular LOR, Rij, corresponding to a crystal pair is:
Rij = ri x rj x 2t
where ri and rj are the event rates of crystal i and crystal j, and t is the coincidence window width. As the activity in the subject increases, the event rate in each detector increases, but the random event rate will increase as the square of the activity.

16. Scatter correction – Model based Attenuation Correction
Uncorrected emission image
Transmission image
This technique allows scatter correction to be adapted to the morphology and size of each patient
Scatter
estimation
Attenuation Correction
Reconstruction - =
Uncorrected
Scatter
Corrected

17. Random Estimation
Background subtraction
Singles Rate Calculation
R = 2 t N1N2
Delayed Window method

18. Noise Equivalent Count or NEC
The scanner sees Prompt coincidences: P=T+S+R
But, we want True coincidences: T=P-S-R (adds noise) = T2
T + S + r
• f • R
Activity Concentration (MBq/mL)
50000
100000
150000
200000
250000
300000
0.005
0.01
0.015
0.02
0.025
0.03
0.035
0.04
TRUES
RANDOMS
NEC
Counts/s
SNR2  NEC
where
T - Trues,
S - Scatter,
R - Randoms,
r - factor to account for the randoms correction method
f - is the fraction of randoms in the center 24 cm of the field of view [NEMA]
NEC is the number of counts detected as a function of the activity, after correcting for the effects of random and scattered events and taking into account deadtime losses.
NEC is determined partly by detector type, detector and scanner geometry, acquisition mode, and front-end electronic. NEC comparison is major arguing point for some vendors.

19. Scintillator block cut into crystals Crystal cuts form light guide
Detector design : Block
4 PMTs
Scintillator block cut into crystals
Crystal cuts form light guide
Blocks are stacked to form a ring detector. PET systems have more than one ring.

20. CTI – HR+

21. Event positioning A C B D x = y = PMTs A + B - C - D A + B + C+D
The multiple crystals are formed by making a series of cuts into crystal at varying depths, none of which will extend through the entire crystal. If the event occurs in crystal #1, approximately 100% of the light produced in the crystal will be received by PMT A while approximately 0% is received by PMT B. If the event occurs in crystal #2, PMT A will receive approximately 93% of the light and B, 7% of the light. If the event occurs in crystal #3, PMT A receives approximately 78% of the light and B, 22%. Light produced in crystal #4 results in PMT A receiving approximately 59% and B, 41 %. Etc.. 2 3 1 4 5
PMTs A C
Event locations are determined taking by weighting the amplitudes of the signals from the four photomultiplier tubes:
A + B - C - D B D
x = y
A + B + C+D
A + C - B - D
y =
A + B + C + D x

22. Phillips uses an Optically Continuous Lightpipe
39 mm
PMT
Lightguide
1.9 cm
.5 cm
2 cm
GSO crystal array
Light from one crystal reaches 7 nearest PMTs
4 x 4 x 20mm3 GSO crystals
7 PMT local cluster
39-mm PMTs

23. Timing Coincidence Record event Dt < 10 ns ?
In a PET camera, each detector generates a timed pulse when it registers an incident photon. These pulses are then combined in coincidence circuitry, and if pulses fall within a short time-window, they are deemed to be coincident.

24. Scintillators used in PET scanners
NaI
BaF2
BGO
LSO
GSO
Effective atomic no. (Z) 51 54 74 66 59
Lin. atten. coef. (cm-1)
0.34
0.44
0.92
0.87
0.62
Index of refraction
1.85
2.15
1.82
1.85
Light yield [%NaI:Tl]
100 5 15 75 41
Peak Wavelength (nm)
410
220
480
420
430
Decay const. (nS)
230
0.8
300 40 56
Fragile
Yes
Slight
No  No No
Hygroscopic
Yes No No No No

25. PET imaging : Projection images are not available during acquisition.
"List mode"Acquisition
[S0, f0, E1,E2, recorded for each
detected coincidence event (LOR)]
Sinograms
"Rebbining"
Reconstruction

26. Events along a particular row in the sinogram represent those associated with parallel LORs. However, due to the curve nature of the detector ring, LORs become closer together towards the edge of the FOV.
0[deg]
0[det]
144[det]
S (xr,f)
p(xr,f)
180[deg]
To correct for this, the data must be re-sampled (arc corrected) prior to reconstruction. Arc correction is typically applied before reconstruction.

27. Depth of Interaction:
When a photon strikes a crystal, it travels a certain distance before its energy is converted into light. If the photon comes from the center of the field of view (FOV), the line of response (LOR) is likely to be correctly localized in the crystal in which the photon entered. The further away from the center of the FOV, the less likely the LOR will be calculated correctly because the photon will hit the crystal on an angle and continue traveling to another crystal before it lights up.
From

30. What is the difference between “2D” and “3D” modes of operation?
Scatter stopped
by septa
3D mode
Scatter assigned
to wrong LOR
From "Aspects of Optimisation and Quantification in Three-Dimensional Positron Emission Tomography", by RD Badawi.

31. The number of possible LORs increases when the septa are removed
2D – coincidences between detectors in the same or neighboring rings permitted.
3D – coincidences between any pair of rings permitted.
From "Aspects of Optimisation and Quantification in Three-Dimensional Positron Emission Tomography", by RD Badawi.

32. Sensitivity 2D 3D Sensitivity : 0.5 % SF: 12%
From "Aspects of Optimisation and Quantification in Three-Dimensional Positron Emission Tomography", by RD Badawi.

34. Time of Flight (TOF) PET

35. PET -Time of Flight (TOF)
Uses difference in photon detection times to guess at tracer emission point
• without timing info, emission point could be anywhere along line
• c = 3x1010cm/s, so Dt = 600 ps ~ Dd = 10 cm in resolution t1 t2
Time-of-Flight Reconstruction. With conventional reconstruction (shown on the left), all pixels along the chord are incremented by the same amount. With time-of-flight reconstruction (shown on the right), each pixel on the chord is incremented by the probability (as determined by the time-of-flight measurement) that the source is located at that pixel. amount.
From: W. W. Moses, IEE, TNS

38. The factors affecting quantitation:
Attenuation
Scatter
Randoms (PET only)
Normalization
Dead Time (more in PET than SPECT)
Partial volume
Noise

39. Is attenuation effect more pronounced in PET or SPECT ?

40. SPECT PET N = N0 e –m x1 N = N0 e –m x1 e –m x2 N = N0 e –m (x1 + x2)
Attenuation changes depending
on emission site.
Attenuation of all photons traveling along
LOR is the same for position along the LOR.

41. No Attenuation correction Attenuation correction
Example 3
Example 2
No Attenuation correction
Attenuation correction

42. Options in collecting and processing transmission scan data
PET Attenuation:
Options in collecting and processing transmission scan data
Transmission scanning with an external photon source
Combined PET/CT scanners

43. Attenuation correction - Transmission scanning with an external source
The fraction absorbed in a transmission scan, along the same line of response (LOR) can be used to correct the emission scan data
PET transmission (TX) scans are done at or near 511 keV
Transmission scans are noisy and slow (~ 3 min bed  addition min for whole body PET scan)

44. CT based attenuation correction

45. CT based attenuation correction – converting CT numbers to attenuation values
Bi-linear scaling methods apply different scale factors for bone and non—bone materials
Should be calibrated for every kVp

46. D. Townsend , SNM'04

47. PET/CT scanners
PET CT
PET CT
Siemens “Biograph”
Philips “Gemini”

49. Example : Good quality scan
Imaging time 3 min/bed on PET/CT LSO system

50. Extremely Large Patient – poor scan quality
Imaging time was increased to 5min/bed – probably needs min

51. Potential problems for CT-based attenuation correction
Artifacts in the CT image propagate into PET image
Difference in CT and PET respiratory patterns can lead to artifacts near the dome of the liver
Use of contrast agents or implants can cause incorrect PET values
Bias in Ct image due to beam hardening and scatter from the arms in the field of view
Truncation of CT image due to keeping arms in down in the field of view to match the PET scan

52. Attenuation Artifacts :
Hip-implant  overcorrection:
Same artifacts with any metal object (Pacemaker, dental fillings, clips, …)
Solution : Look at Non-attenuation corrected PET

53. Attenuation Artifact:
Breathing

54. Attenuation Artifact:
Dental Artifact

57. Requirements for TOF Fast Scintillator (currently LSO)
Fast, Stable PMT
Fast, Stable electronics
Image reconstruction algorithms for TOF data

58. Brain PET-MRI
Comparative brain images with diffusion echoplanar imaging sequence applied during PET acquisition. Images courtesy of Bernd Pichler, Ph.D., and colleagues at the University of Tübingen and David Townsend, Ph.D., and colleagues at the University of Tennessee.
The prototype dedicated brain PET scanner uses next-generation avalanche photodiode detector (APD) technology to render the PET scanner impervious to magnetic fields and, in turn, not interfere with the imaging capabilities of MRI. The PET scanner is integrated into a standard Siemens 3-tesla Magnetom Trio MRI unit. The combined system has a spatial resolution of 3 mm, axial field-of-view of 19 cm, and transaxial field-of-view of 30 cm.

59. Whole-Body PET-MRI MRI PET
The Philips PET-MRI scanner layout, showing the separation of the two scanner types by a revolving bed. This layout prevents interference of the PET imaging by the strong magnetic field of the MRI scanner and the bed enables easy transfer of the patient from one scanner to the other and accurate positioning of the patient in each scanner.
MRI
PET
Some of the first clinical images from Philips' PET/MR system.

60. Standard Uptake Value (SUV):
Counts *Weight
SUV =
*1000
Dose
Where:
Counts (Bq/cc) = Pixel value calibrated to Bq/cc and decay corrected to scan start time
Weight (kg) = Patient weight in kg
Dose (Bq) = Dose in Bq at injection time (This value is decay corrected to scan start time. )
1000 = Number of cc/kg for water (an approximate conversion of patient weight to distribution volume)
Tissue Typical FDG SUV
Lung
Bone marrow
Breast
Liver
Tumor >3-4

61. SUV Variability: Effect of Object Size (Partial Volume Effect)
Problem:
Finite Spatial Resolution Blurs Object of Size Close to Spatial Resolution
True Max
Measured Max
Activity
Profile
Smaller objects (< 2 FWHM) will decrease SUV value due to partial volume effect. This is system dependent and varies with filter and slice thickness.
Correction: Using phantom with different sphere sizes to determine partial volume effect (and recovery coefficients) for different sizes on each system.

62. SUV Variability: Effect of Object Size1 2 3 8 41 51 43 42 52 61 63 62 9 7
16.0
12.0
Sphere # Diameter Sphere/Bkg
(mm) Activity
5 14.2
SUV-max
8.0
4.0 1 2 3 41 42 43 51 52 61 62 63 7
Bkg
Sphere #

63. Change in SUV due to change in Imaging parameters:
WB protocol
Imaging: 7 beds; 3 min bed
Reconstruction: OSEM (2i,8s),
Zoom 1(5.15mm/pixel), Filter- 5mm
Head& neck protocol
Imaging: 2 beds; 10 min bed
Reconstruction: OSEM (6i,16s),
Zoom 1.5(3.43 mm/pixel), Filter- 5mm
Head Protocol:
Imaging: 1 beds; 10 min bed
Reconstruction: OSEM (6i,16s),
Zoom 2.5(2.15 mm/pixel), Filter- 2mm