Nuclear Medicine Physics Positron Emission Tomography (PET) Jerry Allison, Ph.D. Department of Radiology Medical College of Georgia

Medical College of Georgia And Sameer Tipnis, Ph.D. A note of thanks to Z. J. Cao, Ph.D. Medical College of Georgia And Sameer Tipnis, Ph.D. G. Donald Frey, Ph.D. Medical University of South Carolina for Sharing nuclear medicine presentation content

SPECT vs PET PET SPECT (Simultaneous acquisition) (Step-and-shoot acquisition) 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

SPECT vs PET imaging Attribute SPECT PET Detection Single s Coincident s Radionuclides 99mTc, 67Ga, 111In 18F, 82Rb, 13N, E 70 – 300 keV 511 keV Spatial res.  10 – 12 mm  5 - 6 mm Atten.Correction No / Yes* Yes * Possible with SPECT/CT or transmission source systems 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Three steps in PET imaging 1. Emission of positron by radionuclide EC ~3.3% 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Three steps in PET imaging 2. Annihilation of positron & emission of photon pair 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Three steps in PET imaging 3. Coincidence detection of photon pair 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Lines of response (LOR) PET image formation t1   t = t1 – t2  t < 6 (to 12) ns ? Yes Register as a “coincident” event  t2 Why 5 & 12ns? Lines of response (LOR) Positional information is gained LOR is assigned by electronic coincidence circuitry 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

LORs combined to form image Reconstructions - typically OSEM (iterative) 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

+ emitters used in PET p  n + e+ +  Proton-rich nuclei: positron emission p  n + e+ +  18F9  18O8 + e+ +  T1/2 = 110 min 15O8  15N7 + e+ +  T1/2 = 2 min 13N7  13C6 + e+ +  T1/2 = 10 min 11C6  11B5 + e+ +  T1/2 = 20 min 82Rb37  82Kr36 + e+ +  T1/2 = 73 sec

Two Scientists Awarded Nobel Prize In Physics For Neutrinos Discoveries OCTOBER 06, 2015 NPR: Arthur McDonald of Canada and Takaaki Kajita of Japan were awarded Nobel Prize in Physics Tuesday for discovering that subatomic particles called neutrinos can switch from one kind to another. NPR has more about the win and how it could change physics in a big way. … Today the Nobel Prize in physics was awarded to two researchers. Takaaki Kajita, of Japan, and Arthur McDonald, of Canada, won for showing that particles called neutrinos have mass. For neutrinos to change type, the laws of physics say they have to have mass. But the best theories at the time predicted neutrinos were weightless. In other words, these experiments showed the best theories were wrong.

Cosmic Gall by John Updike Neutrinos Neutrinos, they are very small. They have no charge and have no mass And do not interact at all. The earth is just a silly ball To them, through which they simply pass, Like dustmaids down a drafty hall Or photons through a sheet of glass. They snub the most exquisite gas, Ignore the most substantial wall, Cold-shoulder steel and sounding brass, Insult the stallion in his stall, And, scorning barriers of class, Infiltrate you and me! Like tall And painless guillotines, they fall Down through our heads into the grass. At night, they enter at Nepal And pierce the lover and his lass From underneath the bed—you call It wonderful; I call it crass.

Annihilation  Energy conservation: e- + e+ = 2   Energy conservation: me- = me+ = 511 keV  Eg + Eg = 2 x 511 keV Momentum ( = m ) conservation: pe- = pe+ = 0  pg + pg = 0 but pg  0  2 g always in opposite directions The two coincident g’s are detected in PET.

Uncertainties in annihilation 511 keV angle divergence 180o±0.3o (18F) positron scatters in tissue to lose energy + 511 keV 18F maximum positron range: 2.3 mm  0.22 mm FWHM resolution limitation

Annihilation location  Ejection location The distance depends on the e+ initial kinetic energy and medium. Isotope Max E Max d FWHM F-18 0.64 MeV 2.3 mm .22 mm C-11 0.96 MeV 3.9 mm .28 mm O-15 1.72 MeV 6.6 mm 1.1 mm Rb-82 3.35 MeV 16.5 mm 2.6 mm Shorter distance in a medium with higher density or higher Z

Aside: + endpoint energy & spatial resolution 82Rb Emax. = 3.35 Mev Rangerms  2.6 mm Emax. = 0.64 Mev Rangerms  0.2 mm 18F 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Residual momentum of e+ and e- Neither positron nor electron are at complete rest when annihilation occurs. The residual momentum causes a small angular deviation from 180. h  0.0022 × ring diameter For D = 80 cm, h ~ 2 mm g LOR

Ultimate spatial resolution in PET The uncertainties in annihilation (location & residual particle momentum) determine the ultimate (limiting) spatial resolution (~ 2 mm)

Coincidence Detection 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Types of coincidences (correct LOR assigned) (incorrect LOR assigned) True Scatter Random True coincidences form a “true” distribution of radioactivity Scatter & random coincidences distort the distribution of radioactivity, add to image noise, degrade image quality 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Count rates P = T + R + S T = P - S - R R = t R1 R2 P = “Prompts” (count rate measured by detector pair) T = Trues S = Scatter R = Randoms R1, R2 = singles count rates in detectors 1 and 2 P = T + R + S T = P - S - R R = t R1 R2 Typically, t ~ 6 – 12 ns t = coincidence timing window 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

A = injected activity T  A, R1  A, R2  A But, R = t R1 R2 Increasing injected activity to compensate for the size of an obese patient not very effective! (randoms rate increases faster than trues rate) To improve image quality, increase the acquisition time (therefore increasing the signal-to-noise ratio) 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Time window and random coincidence Coincidence time window: 6 - 12 ns Larger window: higher count rate but more random coincidence counts More injection activity  higher single count rate  more random coincidence counts c o i n c i d e n c e c i r c u i t

Estimating random coincidence Delay the coincidence time window to acquire pure random count rate Two coincidence time windows separated by 64 ns (0-12 ns and 64-76 ns) No true events in delayed window No scatter events in delayed window Randoms count rate same in both windows

Major components of a PET scanner Scintillation detector rings To convert 511 keV g photons to light photons PMTs To convert the light photons to electrons and to greatly multiply the electron number PMTs being eliminated in some cameras: PET/MR (avalanche photodiodes) Electronic circuits To amplify, shape, manipulate and discriminate the electrical signals Computers To acquire, process, display, and store images

No collimators in a PET scanner Photon direction determined by LOR  no collimators Absence of lead improves: detection efficiency (count rate) spatial resolution

Modes of PET imaging 2-D (with septa): Coincidences between detectors in the same or nearby ring permitted 3-D (without septa): Coincidences between detectors in any ring permitted 2-D imaging: high resolution (reduces randoms + scatter) 3-D imaging: high sensitivity (increased randoms + scatter) 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

PET systems Multiple rings ~5 rings ~8 slices/ring Ring diameter  80 to 92 cm Transverse FOV  60 cm Axial FOV  15 - 20 cm Attenuation correction X-ray CT 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

2015. Nuclear Medicine Physics for Radiology Residents 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Detector materials BGO (Bi4Ge3O12) used by GE LSO (Lu2SiO5) used by Siemens GSO (Gd2SiO5 ) used by Philips LYSO (Lu2YSiO5, 9(L):1(Y)) used by all

Stopping power (attenuation coefficients) 511 keV 140 keV NaI 0.32/cm 2.44/cm BGO 0.86/cm 11.76/cm LSO 0.81/cm 9.26/cm GSO 0.67/cm 6.37/cm LYSO 0.50/cm _____________________________________________________________________________ N(x) = N e-mx For x = 3 cm, NBGO = 7.5% (for a 3 cm thick BGO detector, 92.5% of 511 keV photons absorbed)

Scintillation decay time If time interval between detecting 2 g’s too short  pile-up effect  reduced count rate and artifacts NaI BGO LSO GSO LYSO 230 ns 300 ns 40 ns 60 ns 53 ns

NaI BGO LSO GSO LYSO Energy resolution Depending on both fluctuation of blue photon number and light output NaI BGO LSO GSO LYSO 10% 17% 12-18% 10% 12-18%

Detector blocks PET 20 – 30 mm PET spatial resolution is primarily dictated by dimensions of individual detector elements (width ~ 4 – 6 mm) Each detector element optically isolated by reflective material in crystal cuts PMT array can determine which detector element(s) absorbed 511 keV photon 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Detector blocks ~300 optically independent blocks in a PET scanner partial cuts of the detector -> 8 x 8 small crystals 4 PMTs per block. ~20,000 detection elements z R T R T A detector optical coupling PMT

Detector blocks Each detection element: 4 mm (T) × 4 mm (A) × 30 (R) mm Small tangential and axial sizes  good spatial resolution Large radial size  high stopping power  high count rate All blocks acquire data simultaneously  significantly increasing count rate

Assembly of detector blocks 1 block has 8 x 8 = 64 detectors 1 bucket has 4 blocks = 256 detectors 1 ring has 16 buckets = 4096 detectors 5 rings block 1 block 2 bucket 1 block 3 block 4

October 7, 2015 -- Researchers at the University of California, Davis (UC Davis) have received a five-year, $15.5 million grant to develop what they are calling the world's first total-body PET scanner. National Cancer Institute and will fund the Explorer project, led by Simon Cherry, PhD, distinguished professor of biomedical engineering and Ramsey Badawi, PhD, a professor of radiology. The total-body PET scanner would image an entire body all at once, and it would acquire images much faster or at a much lower radiation dose by capturing almost all of the available signal from radiopharmaceuticals. … the design would line the entire inside of the PET camera bore with multiple rings of PET detectors. … such a total-body PET design could reduce radiation dose by a factor of 40 or decrease scanning time from 20 minutes to 30 seconds http://www.auntminnie.com/index.aspx?sec=sup&sub=mol&pag=dis&ItemID=112051

Advantages of PET imaging No collimators  higher detection efficiency and better spatial resolution Ring detectors  higher detection efficiency Block detectors  higher detection efficiency and better spatial resolution

2014 PET image of the ACR phantom 7.9 mm

2014 SPECT image of the ACR phantom as 9.5 mm 31.8 mm 15.9 mm Sphere diameters: 9.5, 12.7, 15.9, 19.1, 25.4, and 31.8 mm

Time-of-flight PET Theoretically it is possible to determine the annihilation location from the difference in arrival times of two g photons: d = c∙Dt/2. Because of fast speed of light (c = 30 cm/ns), fast time resolution of detection is required for spatial accuracy. e.g. 0.067 ns  1 cm accuracy No such fast scintillator yet. The currently used LYSO for ToF PET has a time resolution of 0.585 ns which leads to 8.8 cm accuracy. d LOR t1 t2

Time-of-flight PET ToF is used to improve SNR. The improved SNR is used either for better image quality or for shorter scan time. No additional hardware is needed except more CPU due to the heavy computation. 1 cm t1 t2 0.067 ns 9 cm t1 t2 0.585 ns

Time of flight PET image of a big patient Better resolution with ToF!

Iterative reconstruction © Physics in Nuclear Medicine: Cherry, Sorenson and Phelps

LORs combined to form image Reconstructions - typically OSEM (iterative) 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Iterative reconstruction algorithms One iteration at one view: Forward project the image Compare to acquired data Backproject P - P0 Update the image N1 = N0 + bpj (P - P0) N0 P0 P N1 P-P0

Iterative reconstruction algorithms Maximum likelihood - expectation maximization (ML-EM) accurate but slow convergence Updates pixel values once after comparisons for all views Ordered subset - expectation maximization (OSEM) Updates pixel values after comparison for a subset of views Number of views in subset increases as convergence occurs Less iterations needed to achieve same accuracy (faster convergence)

Iterative reconstruction algorithms © Physics in Nuclear Medicine: Cherry, Sorenson and Phelps

PET Data Corrections Attenuation (MOST IMPORTANT CORRECTION) CT based Normalization Correction for variation in performance of ~20,000 individual detectors Random coincidences Delayed coincidence time window (~64 ns) Scattered radiation Modeling from transmission & emission data Extrapolation from tails of projections Dead time Empirical models

Photon attenuation within patient Every PET study is compensated for attenuation. Correction of attenuation in PET reconstruction needs attenuation map from CT  values must be extrapolated from CT energies (< 120 keV) to 511 keV w/o compensated

Biograph TruePoint PETCT Discovery PET/CT 710 (GE) Biograph TruePoint PETCT (Siemens) Ingenuity TF PET/CT (Philips) 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Attenuation Correction in PET/CT 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Attenuation Correction Without AC With AC 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Photon attenuation in PET High uptake in lungs Low uptake in others High uptake in skin w/o attenuation compensation attenuation compensation

Impact of Misregistration on Attenuation Correction Lateral walls of myocardium in the PET data, corrected with the lower attenuation of lung tissue Proper registration Lateral walls of myocardium corrected with the higher attenuation of heart tissue (Ref: Attenuation correction of PET cardiac data with low-dose average CT in PET/CT, Tinsu Pan et al, Med. Phys. 33, October 2006) 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Typical PET / CT imaging (1) (2) 512 x512 128 x128 120 kVp  511 keV 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Role of FDG (Fluorodeoxyglucose) 18F-FDG is a glucose analog Actively transported into cells by GLUT (glucose transport proteins) Both glucose and FDG phosphorylated by hexokinase Glucose-6-phosphate undergoes further metabolism in the glucose pathway 18F-FDG-6-phosphate does not, and remains trapped in the cell 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Role of FDG Tumor cells  increased level of GLUT-1 and GLUT-3  higher levels of hexokinase  highly metabolically active (high mitotic rates)  favor the more inefficient anaerobic pathway adding to the already increased glucose demands These combined mechanisms allow for tumor cells to absorb and retain higher levels of FDG compared to normal tissues 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Role of FDG NOTE - FDG is NOT cancer specific and will accumulate in areas with high levels of metabolism and glycolysis  increased uptake can be expected in sites of: (1) hyperactivity (muscular, nervous) (2) active inflammation (infection, sarcoid, arthritis, etc.) (3) tissue repair, etc. 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Hyperactivity High FDG uptake in pectoralis major after strenuous exercise 24 hours prior to study 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Inadequate fasting 45 min fasting Overnight fasting 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Semiquantitative PET: Standard Uptake Value (SUV) Defined as the ratio of activity concentrations SUV = conc. in vol. of tissue / conc. in whole body SUV = (MBq/kg) / (MBq/kg) Usually, SUV ~ 2.5 taken as cut-off between malignant and non-malignant pathology 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

SUV in clinical studies Numerator: highest pixel value (SUVmax) from an ROI Or SUVmean Denominator: Activity administered/ body mass Or lean body mass Or body surface area SUV will depend on – physiologic condition, uptake time, fasting state, etc. Image noise, resolution, ROI definition Small changes in SUV need to be interpreted carefully 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Requirement for reproducible SUV 18FDG uptake period, scan length, scanning range, scanning direction (e.g. head to toe) Patient preparation: fasting, medication Reconstruction parameters: slice thickness, filters Region-of-interest definition (SUVmax / SUVmean/ body mass/ lean body mass/ body surface area) Consistency is the most important factor! 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Effect of uptake time 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Clinical Use of PET Oncology (~ 90%) Cardiac & Neuro (~ 10%) 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Typical oncology protocol Administered dose – 10 to 20 mCi FDG ~ 60 mins. in “quiet” room” to allow adequate uptake and trapping, clearance from blood Scanning – typically “eyes-to-thighs” 6 to 7 “bed” positions (each ~ 15 cm FOV) Total scan time is ~ 30 mins. (3 mins/bed) With time, SUVtumor  & SUVbkg.  2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Patient dose (FDG) Effective dose to pt. Organ of max. dose: bladder 10 mCi (370 MBq) injection  ~ 7 mSv Organ of max. dose: bladder Equivalent dose 10 mCi (370 MBq) injection  ~ 63 mGy CT (for AC) ~ 5 mSv CT (Diag. / contr.) ~ 15 – 18 mSv 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Personnel dose (FDG) 2.4 Sv / hr per mCi @ 1 m, 1 hr uptake + void Effective dose (whole body exposure) 10 mCi (370 MBq) injection  ~ 24 Sv/hr @ 1 m Doses lower towards pt. feet compared to torso Minimize time of pt. contact at injecting, escorting to rest room, positioning for scan Increase distance from the pt when communicating 2.4 mRem/hr @ 1 m 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

SPECT & PET SPECT – 2 views from opposite sides Res.  collimator res., which degrades rapidly with increasing distance from collimator face PET – Simultaneous acquisition Res.  detector width; is max in center of ring SPECT sensitivity ~ 0.02% Huge losses due to absorptive collimators PET sensitivity- 3D ~ 2% or higher High sensitivity due to coincidence detection (electronic collimation) 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR

Superior spatial resolution Higher sensitivity Attenuation Correction Advantages of PET over SPECT Superior spatial resolution Higher sensitivity Attenuation Correction 2015 Nuclear Medicine Physics for Radiology Residents Sameer Tipnis, PhD, DABR