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Microsoft PowerPoint - resident NM


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  • 1. Physics of nuclear medicine introduction historic and current NM technologies principle of gamma camera image quality and gamma camera performance characteristics gamma camera QC data acquisition and processing methods SPECT and SPECT/CT other devices Physics of nuclear medicine Cherry SR, Sorenson JA, Phelps ME, “Physics in Nuclear Medicine” 3rd ed (2003) Chapters 12, 13, 14, 15, and 17 Introduction of nuclear medicine radiopharmaceutical (a radionuclide attached to a chemical compound) administered to patient, then (hopefully) concentrated to the abnormal sites through interaction between the pharmaceutical and cells or molecules decay of the radionuclide in the sites: emitting of single or annihilation photons detection of the emitted photons using gamma camera, PET scanner or other devices Introduction of nuclear medicine sensitive to functional changes earlier detection of diseases and exclusive diagnostic capability, e.g. perfusion for heart, brain, kidney and lungs, and metabolism for cancers interaction at cellular or molecular levels bound directly to a target molecule (111In- monoclonal antibody), low sensitivity accumulated by molecular or cellular activities of the target (18F-FDG, 99mTc-sestamibi, 131I−), high sensitivity
  • 2. Introduction of nuclear medicine emitted photon energy: 70 to 511 keV most low-energy photons absorbed by tissues most high-energy photons penetrating the detector charged particles penetrating only mm of tissue pixel value of the image: concentration of radioactivity may need post-acquisition data processing poorer image quality due to limited photon number and poor spatial resolution History of nuclear medicine 1895 discovery of x-ray by Roentgen 1896 discovery of radioactivity by Bequerel 1898 production of radium by Curie 1927 use of radon to measure the blood transit time 1930s invention of cyclotron by Lawrence 1945 invention of nuclear reactor 1951 rectilinear scanner to acquire images History of nuclear medicine 1958 invention of Anger camera 1964 use of Tc-99m (I-131 only prior to 1964) Tc-99m: metastable (T1/2 = 6.01 hr) pure γ decay (E = 140 keV), flexible for labeling I-131: electrons and 364 keV photons, thyroid disorders only 1970 derivation of image reconstruction algorithm for tomography (CT, SPECT, PET) 1998 rapid spread of PET and PET/CT The most often used radionuclide: Tc-99m metastable state of 99Tc43: T1/2 = 6.01 hr long enough for imaging but short for reduced radiation dose to patient pure γ decay: less radiation dose E = 140 keV: enough photons to escape from the patient body but most stopped by the detector flexible for labeling (attached to a pharmaceutical): wide clinical applications
  • 3. The most often used radionuclide: Tc-99m The first rectilinear scanner (1951) The first Anger camera (1958) Dual-head gamma camera
  • 4. SPECT gamma camera Two detectors mounted on a rotation gantry with different angles (180°, 90°) for tomography Mobile semiconductor gamma camera 15×20×10 cm CZT detector breast imaging sports medicine ER and OR imaging 3000 0.3×0.3 cm discrete crystals 48 PSPMT Planar imaging Tc-99m sestamibi 3 mm lesion detectable Tc-99m MDP Dynamical imaging a series of images with time
  • 5. SPECT imaging transaxial coronal sagittal SPECT imaging short-axis vertical long-axis horizontal long-axis Pros and cons of nuclear medicine inherent molecular imaging high sensitivity low concentration of radionuclide ~ pmol/liter biodistribution depends not only on the specificity of the carrier but also on the route of administration. noisy and suboptimal resolution Molecular imaging ACR definition: Spatially localized and/or temporally resolved sensing of molecular and cellular processes in vivo. SNM definition: Visualization, characterization, and measurement of biological processes at the molecular and cellular levels in human and other living systems. 2-D or 3-D imaging and variation over time Including NM, PET, MRI, MRS, optical, US and CT with contrast
  • 6. Molecular imaging modalities resolution ____ modality sensitivity spatial temporal contrast MRI + 10-100µm msec +++ MRS + 1 cm min-h + PET +++ 3-4 mm min ++ SPECT ++ 8-12 mm min + optical +++ 1-2 mm msec +++ US +++ 1 mm msec ++ +++: high, ++: medium, +: low Probes for molecular imaging bound directly to a target molecule accumulated by molecular or cellular activities of the target activated by the target enzyme in vivo Smart NIR agents With specific enzyme cleavage, fluorophores are separated from the backbone and each other so as to markedly increase their fluorescence. Gamma camera pat ient collimat or det ect or PMT pre- amp amp & sum posit ionPHA comput er display X Y Z
  • 7. Major components collimator to establish position relationship between γ photon source and detector (projection imaging) scintillation detector (NaI(Tl)) to convert x or γ photons to blue light photons photomultiplier tube (PMT) to convert blue photons to electrons and to increase the number of electrons electronics to amplify and discriminate electrical signals display to display the image acquired by gamma camera Collimator to establish position relationship between the source and detector poor spatial resolution (ability to see details) and low detection efficiency (ability to count photons) The weak link of a gamma camera: The collimator determines the resolution and sensitivity of a gamma camera. Collimator design principle: to optimize the trade-off between resolution R and sensitivity η hole size d and hole length l smaller (or longer) holes higher R but lower η septal thickness t penetration < 5% hole orientation: parallel-hole, converging, diverging pinhole: single hole d t Parallel-hole collimator increasing source-to-detector distance leads to same sensitivity same FOV same image size lower R
  • 8. Converging beam collimator increasing source-to-detector distance leads to decreasing FOV increasing image size lower R higher sensitivity FOV F fan cone Pinhole collimator increasing source-to-detector distance leads to increasing FOV decreasing image size lower R lower sensitivity FOV NaI (Tl) detector energy spectrum scintillation process to convert γ photons to blue photons (E ≈ 3 ev or λ ≈ 415 nm) theoretical deposited energy spectrum in detector photopeak: completely absorbed compton edge: Ee = E0 – Es (at 180º) above the edge: multiple scatter below the edge: single + multiple NaI (Tl) detector energy spectrum photopeak single scatter double scatter p.e p.e c.s c.s Compton edge c.s c.s c.s c.s
  • 9. NaI (Tl) detector energy spectrum actual deposited energy spectrum in detector spread photopeak caused by imperfect energy resolution (random fluctuation of blue photon number in detector) backscatter peak due to photon penetrating the detector, backscattered by surrounding material, reentering detector, and absorbed by the detector: Eb + Ee = E0 iodine escape peak 30 keV K-shell x-rays following p.e. absorption of iodine: Ee ≅ E0 – 30 keV lead K-shell x-ray (80 – 90 keV) following p.e. in lead NaI (Tl) detector energy spectrum backscatter iodine escape lead x-rays p.e c.s p.e p.e NaI x-ray x-rayp.e e x γ NaI (Tl) detector energy spectrum Hg-197 w.o. scatter I-131 w/w.o. scatter Advantages of NaI (Tl) detector good stopping power for low-energy γ (ρ = 3.67 g/cm3, Zeff = 50, PE dominant) µ = 16.58 cm-1 @ 69 keV, t = 0.95 cm, T ≅ 0% µ = 2.57 cm-1 @ 140 keV, t = 0.95 cm, T = 7.7% µ = 0.72 cm-1 @ 247 keV, t = 0.95 cm, T = 48.5% good detector linearity over 20 - 2000 keV good conversion efficiency: ~ 26 eV/blue photon good transparent to blue photons blue photons matched with the performance of PM tube easy to manufacture
  • 10. Disadvantages of NaI (Tl) detector poor stopping power at Eγ > 200 keV slow scintillation decay (230 ns) low counting rate Compton scatter dominated at Eγ > 250 keV poor spatial resolution fragile must keep dry Photomultiplier tube to create and amplify e-pulse photocathode (CsSb): blue light to electrons 9 - 12 dynodes: each increasing electrons 3 – 6 times anode: collect electrons: 610 ≅ 6 × 107 NaI(Tl) 0 + 3 0 0 v + 4 0 0 v + 5 0 0 v + 6 0 0 v + 7 0 0 v cathod anode dynode To preamplifier Photomultiplier tube stable high voltage 1200 V needed for 10 dynodes 1% increase of high voltage 10 % increase of current at anode sealed in glass and evacuated wrapped in ‘Mu-metal’ (alloy of Fe, Ni, Cu) to shield magnetic field magnetic field affecting focusing of electron beam Photomultiplier tube 40 to 100 PM tubes (d = 5 cm) in a modern gamma camera photocathod directly coupled to detector or connected using plastic light guides anode connected to electronics in the tube base ultrasensitive to magnetic field
  • 11. Electronics preamplifier to amplify pulses from the PM tube to match impedances between the detector and subsequent components to shape pulses for subsequent processing voltage- and charge-sensitive circuits amplifier to amplify pulses from mV to V to reshape slow decay pulses to narrow ones using resistor-capacitor circuit baseline restoration circuit Electronics pulse height analyzer: selecting the pulses of certain voltage amplitudes (channel) to discriminate against unwanted γ photon lower-level discriminator upper-level discriminator anticoincidence circuit 1 2 3 V2 (154 keV) V1 (126 keV) Electronics position circuit x y Z YY ky Z XX kx YYXXZ i i i i y i i i i x i i i i i i i i ∑∑ ∑∑ ∑∑∑∑ −+ −+ −+−+ − = − = +++= Display cathode ray tube (CRT) linearity dynamic range contrast brightness LCD: thin film transistor (TFT) plasma display e- source anode def lect ion plates screen z y x
  • 12. Detection of a γ-photon 1 γ-photon 1 electrical pulse (1 count) The photon may experience p.e in the detector (A), c.s in the detector (B), or c.s in the patient (C). energy deposited on the detector # blue photons pulse height entire energy maximum pulse height (A) partial energy reduced pulse height (B, C) A B C BA C Image quality main factors of image quality: 1. contrast: the difference in count density between two objects (or background) C = (Imax-Imin)/(Imax+Imin), MTF (k) = Cout(k)/Cin(k) 2. resolution: ability to distinguish between two objects in close distance, measured by full width at half maximum (FWHM) of PSF image sharpness and details 3. artifacts Factors determining image quality camera performance characteristics detection efficiency count rate image noise contrast, resolution collimator performance resolution patient-to-detector distance resolution energy resolution width of energy window scatter counts contrast non-uniform FOV artifacts dead time artifacts or count loss at high count rates Factors determining image quality patient motion contrast, resolution, artifacts photon attenuation and scatter contrast low-pass filter in the reconstruction resolution wrong energy window contrast, artifacts
  • 13. Non-uniform FOV collimator defect defected PMTs Image noise and off-peak effects 50,000 500,000 1,000,000 2,000,000 Collimator performance low-energy all purpose (LEAP) collimator better efficiency but worse resolution low-energy high resolution (LEHR) better resolution but worse efficiency low-energy fan-beam (LEFB) collimator low-energy cone-beam (LECB) collimator medium-energy all purpose (MEAP) high-energy all purpose (HEAP) collimator Patient-to-detector distance system resolution Rsys intrinsic resolution Rint collimator resolution Rcol at d > 5 cm, Rcol >> Rint larger d poorer Rcol poorer Rsys RRR 2 col 2 intsys +=
  • 14. Detection efficiency Energy resolution and energy window energy spread due mainly to fluctuation of the blue photon number in the detector and of electric signal in the subsequent electronics energy resolution: 8 – 10% for NaI ~ 20% for BGO energy window: ±10% for NaI ±30% for BGO better energy resolution smaller energy window fewer scatter counts Multiple energy window summing images to increase count rate Tl-201: 70±10% keV + 167±10% keV In-111: 171 keV + 245 keV Ga-67: 93 keV + 185 keV + 300 keV dual energy window simultaneous acquisition to accelerate study e.g. cardiac perfusion: Tc-99m and Tl-201 140±10% keV and 70 keV + 167 keV Down scatter contamination must be corrected. Performance at high count rates pulse pile-up effects Two events acquired at different locations but same time are recorded as a single event with summed energy at a location between them. 2 scatter counts possibly accepted as 1 event image quality degradation rejected if both events in photopeak count loss
  • 15. Performance at high count rates typical dead time in clinic: 4 – 8 µs 5 µs dead time 20% count loss at 40,000 cps e.g. first-pass cardiac study: 100,000 cps very high count rate may paralyze camera. Camera quality control uniformity: daily, 256×256, > 4M counts resolution: weekly, 512×512, > 4M counts energy and COR: monthly acquisition of new uniformity maps and possible energy map: quarterly, > 30M counts Uniformity of detector integral unif = max. count – min. count < 5% max. count + min. count differential unif = max. count diff. – min. count diff. < 5% max. count diff. + min. count diff. Bar phantoms
  • 16. Data acquisition collimator: LEAP, LEHR, LEFB, LECB, MEAP, HEAP energy window: match the radioisotope and energy resolution pixel size: 1/3 ~ 1/2 of spatial resolution 64×64, 128×128 or 256×256, 2 bytes in pixel depth patient close to the detector, steady, in FOV sizepixel sizedetector sizematrix = Matrix size 64×64 128×128 Data acquisition static acquisition: recording x and y in a matrix dynamic acquisition: recording a sequence of static images at different time, each image corresponding a certain time period list mode acquisition: recording x, y, t (and R-wave trigger for gated list mode), no frames during acquisition and later reframing needed Data processing windowing in display: 2 byte image displayed on a 256 gray color monitor 2-D filtering the image: reducing noise temporal filtering for dynamic images: reducing noise ROI: maximum, minimum, mean counts, s.d. time-activity curve from a dynamic image: renogram, first-pass count profile: often used in camera QC
  • 17. Time-activity curve SPECT eliminate overlaying and underlying activity of a slice better contrast more accurate lesion localization more demanding technically and longer data acquisition more severe image noise Data acquisition a sequence of 2-D static images at different angular positions (views) detector rotation range 180º with 2 perpendicular detectors or 360º with 2 opposite detectors 45º RAO 45º LPO Data acquisition circular or elliptical orbit closer to the patient better spatial resolution step-shoot or continuous acquisition
  • 18. Data acquisition energy window acquisition time or counts per view matrix size for each view depending on the spatial resolution (64×64 or 128×128) number of views = matrix size for 360º SPECT (64 or 128) ECG gated for cardiac SPECT View number 128 views 64 views 43 views 32 views SPECT camera performance mechanical center coinciding with COR using software, calibration and testing all detectors aligned accurately in axial direction to acquire same slice data uniformity < 1% ~ 41 M for 64×64 image SPECT reconstruction filtered backprojection algorithm (FBP) iterative algorithms (OSEM, MLEM) compensation techniques attenuation scatter patient motion spatially variant blurring
  • 19. Filtered backprojection ramp filter required even for noise-free data to remove 1/r blurring low-pass filter to suppress noise Filtered backprojection Hann filter: 0.5 k (1 + cos(πk/kc)) Butterworth filter: 0 k 1 + k / k c 2n 4.25 4.15 8.15 Ramp Hann BW 4.25 BW 4.15 BW 8.15 frequency filter 1 Iterative algorithm to assume an initial image and to update the image iteratively Steps of one iteration: 1. to project the image 2. to compare to the data 3. to backproject P - P0 4. to update the image I1 = I0 + bpj (P - P0) I0 P0 P I1 P-P0 Photon attenuation and scatter attenuation decreased photon number on AB due to absorption and scatter, half of 140 keV photons absorbed over ~ 5 cm in water scatter and downscatter misplaced source position C instead of A det ect or A B C D pat ient θ
  • 20. Photon attenuation effect Attenuation compensation geometric mean P = (p1× p2)1/2 exact compensation for a point source in uniform medium analytical method: uniform attenuation built in FBP, magnifying image noise Chang’s method, for uniform µ (brain SPECT) transmission images attenuation map used in iterative algorithm, most accurate and best noise control p1 p2 Transmission image Gd-153 (97-103 keV, 8 mo) moving line source for parallel-hole collimators stationary line source for fan-beam collimators stationary point source for cone-beam collimators x-ray source and detector (SPECT/CT) p = p0 exp(-Σµi∆xi): Σµi∆xi= ln(p0/p) µi Transmission image scaling µ to the photon energy of emission image downscatter contamination
  • 21. Photon scatter and compensation reduced contrast spill of counts from a hot spot scatter model built in iterative algorithm deconvolution dual energy window method prior to image reconstruction data P acquired from 126 - 154 keV data S acquired from 91 - 125 keV compensated data = P - S/W, e.g. W = 2 Photon scatter and compensation Partial volume effects occurring for small sources Vs resolution volume VR = π.FWHMT 2.FWHMA when Vs < VR, pixel value < concentration Partial volume effects reducing contrast and error in quantitaion, ‘spillover’ recovery coefficient RC = Capparent/Ctrue RC used to correct PV with known Vs and VR , but not feasible in clinic
  • 22. Compensation for movement patient motion 1. a Tc point source with Tl patient 2. fast, repeated acquisition 3. software correction physiological organ movement gated cardiac imaging SPECT/CT scanner A gamma camera and a multi-slice spiral CT scanner on the same gantry with a single patient table SPECT/CT scanner CT: to create attenuation map for SPECT attenuation correction with any radioisotopes Image fusion for SPECT and CT to better localize the disease SPECT/CT advantage over PET/CT: possible to label the imaging agent with a therapeutic isotope to highly-specifically treat the disease SPECT/CT scanner GE Infinia Hawkeye helical CT, 140 keV, 2.5 mA, 4 rows × 384 Elements, 16 slice/min, in-plane res = 4 lp/cm, sw = 0.5 mm Siemens Symbia T, T2, T6, T16 Philips Precedence 6, 16 slice
  • 23. SPECT/CT image fusion Cardiology SPECT/CT image fusion Oncology Gas-filled detectors to measure activity only ionization chamber: dose calibrator and survey meter Geiger-Muller counter (quenching gas): sensitive survey meter, area monitor γ h e + _ . Dose calibrator high pressure (12 a.p.) Ar-filled ion chamber to assay activity only sample volume effect linearity of response versus sample activity
  • 24. Dose calibrator quality control constancy: daily, Cs-137 (660 keV, 30 y) and Co-57 (122 keV, 9 mo): ±10% linearity: quarterly, 10 µCi - 300 mCi Tc-99m, long-term decay or lineator: ±10% accuracy: yearly, Cs-137 and Co-57: ±5% geometry: upon installation, Tc-99m: ±10% Well counter detecting in-vitro x- and γ-rays main components single NaI crystal (4.5×5 cm or 1.6×3.8 cm) with a hole for sample a PM tube preamplifier amplifier SCA or MCA readout device Well counter detection efficiency intrinsic: 100% for Eγ < 150 keV geometry: for < 1 mL sample at bottom: 93% absolute activity: Asam= Astd× [Csam/Cstd] shielding energy calibration dead time ~ 4 µs A < 10 kBq for 50 kBq, 18% loss Thyroid probe measuring thyroid uptake of I-131 in-vivo 5×5 cm NaI(Tl) with 15 cm long conical collimator pointing to neck, thigh bkg calibration phantom with known activity for calculating uptake 1 – 2 cm diff. in depth 10 – 40% diff. in count rate
  • 25. Miniature γ probe used in surgery detecting sentinel lymph nodes with Tc-colloid detecting radioactive monoclonal antibodies of In-111, I-131, I-125 5×10 mm, high directional sensitivity, light, easy to use, no hazard