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Professor Brian F Hutton Institute of Nuclear Medicine University College London

Emission Tomography Principles and Reconstruction. Professor Brian F Hutton Institute of Nuclear Medicine University College London brian.hutton@uclh.nhs.uk. Outline. imaging in nuclear medicine basic principles of SPECT basic principles of PET factors affecting emission tomography.

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Professor Brian F Hutton Institute of Nuclear Medicine University College London

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  1. Emission Tomography Principles and Reconstruction Professor Brian F Hutton Institute of Nuclear Medicine University College London brian.hutton@uclh.nhs.uk

  2. Outline • imaging in nuclear medicine • basic principles of SPECT • basic principles of PET • factors affecting emission tomography

  3. SPECT History • Anger camera 1958 • Positron counting, Brownell 1966 • Tomo reconstruction; Kuhl & Edwards 1968 • First rotating SPECT camera 1976 • PET: Ter-Pogossian, Phelps 1975

  4. Anger gamma camera Detector: 400x500mm ~9mm thick Energy resn ~10% Intrinsic resn 3-4mm Radionuclides: Tc-99m 140keV, 6hr I -123 159keV, 13hr Ga-68 93-296keV, 3.3dy I-131 360keV, 8dy Collimator Designed to suit energy HR: hole size 1.4mm length 33mm septa 0.15mm

  5. Organ-specific options specialized collimators for standard cameras fanbeam parallel conebeam slit-slat crossed slit pinhole

  6. Single Photon Emission Computed Tomography (SPECT) Single Photon Emission Computed Tomography (SPECT) • relatively low resolution; long acquisition time (movement) • noisy images due to random nature of radioactive decay • tracer remains in body for ~24hrs: radiation dose ~ standard x-ray • function rather than anatomy

  7. SPECT Reconstruction sinogram for each transaxial slice Filtered back projection 1 angle 2 angles 4 angles 128 angles 16 angles

  8. Organ-specific systems specialised system designs, with use limitedto a specific application

  9. Isotope Emax (keV) Max range (mm) FWHM (mm) 18F 663 2.6 0.22 11C 960 4.2 0.28 1200 5.4 0.35 13N 1740 8.4 1.22 15O 3200 17.1 2.6 82Rb Positron Annihilation

  10. Coincidence Detection No Collimator detector 1 coincidence window detector 2 time (ns)

  11. PET "Block" Detector Scintillator array PMTs C BGO (bismuth germanate) A B Histogram Images courtesy of CTI

  12. Attenuation Correction in PET attenuation for activity in body N = N0 e -x. e - (D-x) = N0e -D attenuation for external source N = N0 e -D (D=body thickness) 'Exact' attenuation correction (for 511 keV  ~ 0.096/cm attenuation factors: 25-50)

  13. Coincidence Lines of Response (LoR) sinogram fanbeam parallel

  14. PET Reconstruction sinogram 1 angle 2 angles 4 angles 128 angles 16 angles • conventional filtered back projection • iterative reconstruction

  15. Understanding iterative reconstruction detector (measurement) Y • Objective • Find the activity distribution whose estimated projections match the measurements. • Modelling the system (system matrix) • What is the probability that a photon emitted from location X will be detected at detector location Y. • detector geometry, collimators • attenuation • scatter, randoms m X estimated projection object Y1 m X Y2

  16. System matrix pixeli voxelj

  17. ML-EM reconstruction BP NO CHANGE original estimate update (x ratio) patient original projections FP current estimate estimated projections

  18. Image courtesy of Bettinardi et al, Milan

  19. Noise control • stop at an early iteration • use of smoothing between iterations • post-reconstruction smoothing • penalise ‘rough’ solutions (MAP) • use correct and complete system model

  20. Factors affecting quantification courtesy Ben Tsui, John Hopkins

  21. detector - + without attenuation correction transmission with attenuation correction

  22. System matrix: with attenuation m

  23. Partial volume effects • effect of resolution and/or motion • problems for both PET and SPECT • similar approaches to correction • scale of problem different due to resolution • some different motion effects due to timing: • ring versus rotating planar detector

  24. Modelling resolution • Gamma camera resolution • depends on distance • SPECT resolution • need radius of rotation • PET resolution • position dependent

  25. System matrix: including resolution model m

  26. detector colinearity positron range FWHMtotal2 = FWHMdet2 + FWHMrange2 + FWHM1802 PET resolution depth of interaction results in asymmetric point spread function radial int radial ext tangential

  27. Modelling resolution detector (projection) • potentially improves resolution • requires many iterations • slow to compute m • stabilises solution • better noise properties object w/o resn model Courtesy: Panin et al IEEE Trans Med Imaging 2006; 25:907-921 with resn model

  28. Can we consider measurements to be quantitative? • Scatter correction • multiple energy windows for SPECT;PETCT standard models • SPECT local effects; PET more distributed detector object • Scatter fraction • SPECT ~35% PET 2D ~15%; 3D ~40%

  29. Scatter Monte Carlo • influenced by photon energy, source location, scatter medium • reduces contrast measured • scatter models • analytical, Monte Carlo, approximate models • measurement • triple energy window (TEW), multi-energy • subtract from projections: • measured proj – TEW • or combine with projector in reconstruction: • compare (forward proj + TEW) with measured proj

  30. 3D reconstruction • Approaches • rebin data followed by 2D reconstruction • single slice rebinning (SSRB) • multi-slice rebinning (MSRB) • Fourier rebinning (FORE) • full 3D reconstruction • 3D OSEM • 3D RAMLA limits for FORE

  31. 2D 4min 3D 4min 2D 2min 3D 2min VUE Point 3D-OSEM 28subsets 2iter FORE 2D-OSEM 28subsets 5 iter FORE 2D-OSEM 28subsets 2 iter Courtesy V Bettinardi, M Gilardi, Milan

  32. Summary • Emission tomography • functional rather than anatomical • single photon versus dual photon (PET) • main difference is ‘collimation’ • Iterative reconstruction • very similar approach for SPECT and PET • currently most popular is OSEM (or similar) • the better the system model the better the reconstruction

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