Positron Emission Tomography Physics & Instrumentation Dimitra G. Darambara, Ph.D Multimodality Molecular Imaging Joint Department of Physics RMH/ICR Outline Introduction PET Physics overview Types of events Acquisition modes Detectors Performance parameters Data reconstruction and corrections Detector developments New development in PET dimitra.darambara@icr.ac.uk Medical Imaging Techniques Anatomical CT MRI US Functional or metabolic SPECT PET (f)mri(s) (probe) optical (SAI) High resolution morphological capabilities with physiological info Real-time physiological info Biological processes at molecular level PET in Medical Imaging Powerful and sensitive means to non-invasively investigate biological processes in-vivo Radiotracer-imagingimaging technique Inject tracer compounds labelled with positron-emitting radionuclides Images of the static or dynamic distribution of tracer Info on blood flow, metabolism, proliferation, angiogenesis, hypoxia, gene expression In clinic: Diagnosis, treatment follow-up, therapy assessment Oncology,cardiology,neurology, psychiatry,treatment planning Drug development PET: unique molecular imaging modality 3 main factors: No need for heavy absorptive collimators higher sensitivity can detect and image nano- and pico- molar levels of tracers in-vivo Absolute quantification capability through accurate attenuation correction Availability of β -emitting radionuclides that are organic atoms can be substituted in molecules without modifying their biological activity PET radiotracers to investigate any biological process without interfering with normal biochemistry Positron Emission and Annihilation Proton decays to n, e, ν e e combines with e - and annihilates ν 511 kev γ e e - γ 511 kev 1
Annihilation Coincidence Detection A molecular probe labelled with e -emitting radionuclides annihilation -- simultaneous 2 x 511 kev gamma-rays -- 180 -- Line of Response (LOR) -- electronic collimation Isotope Positron-emitting Radionuclides Half-life life Max E Range (MeV) (mm) Production C-11 20.4 mins 0.96 0.4 Cyclotron N-13 9.96 mins 1.20 0.7 Cyclotron Scanner: rings of position-sensitive photodetectors -- millions of coincident pairs detected -- coincidence time window (6-12 ns) -- coincidence logic electronics a time stamp to the record of detected events -- scan time: 5-60 mins sensitivity, acquisition mode, size of ROI, amount of injected activity Reconstructed images of radiotracer distribution e e - O-15 123 secs 1.74 1.1 Cyclotron F-18 110 mins 0.63 0.3 Cyclotron Ga-68 68.3 mins 1.83 1.2 Generator Rb-82 78 secs 3.15 2.8 Generator Cu-62 9.74 mins 2.93 2.7 Generator Cu-64 12.7 hrs 0.65 0.3 Cyclotron e -emitting radionuclide Positron Range --- Non-Colinearity Effective e range e range 511 kev e e- Non-colinearity Error due to noncolinearity 511 kev Positron Range Non-Colinearity 2 fundamental limits of spatial resolution: 2-2.5mm for clinical whole-body scanners Positron Range: --- its distributions have exponential shape with long tails not well described by Gaussian function rms effective range better indicator than FWHM --- radionuclide-specific, depends on positron emission E e 511 kev 511 kev Non-Colinearity: --- can be described as an approximate Gaussian angular distribution around 180 with FWHM~0.5 --- linearly dependent on the distance D between coincident detectors: R = 0.0022xD Detected Events in PET A detected event in PET is valid if: 2 photons are detected within a predefined electronic time window coincidence window Types of Coincidence Events The LOR between the 2 photons is within a valid acceptance angle of the scanner The E deposited in the detector by both photons is within the selected E window E-gating technique around the photopeak in the E spectrum need for detectors with good E resolution v narrow E gate Such coincidence events prompts TRUE RANDOM SCATTERED 2
Types of Coincidence Events Single: a single photon recorded by a detector True coincidence : detection of 2 singles from the same positron annihilation within the coincidence window True Count Rate: T = Aε 2 Ω coin Where A = activity concentration of the source ε = detector detection efficiency for 511 kev Ω coin = solid angle for detection of coincidence annihilation photons Types of Coincidence Events Random or accidental coincidence: detection of 2 singles from 2 different annihilations within the coincidence window Count Rate: R = 2τN x N z where 2τ the width of time window N single event rate incident upon detector x and z High Random count rate impact on noise, dead time and pulse pile-up limit sensitivity Types of Coincidence Events Scattered Events: one or both of the photons from the same annihilation have undergone a Compton interaction (body, detector, scanner gantry) Count Rate: S = Aε 2 Ω d where Ω d the solid angle for the FOV for coincidence scatter events Scatter and Random events undesirable source of background counts loss of contrast resolution and quantitative accuracy in final reconstructed image Can be reduced by narrowing E and coincidence window and limiting FOV activity Prompt Count Rate: Sum {TSR} Noise Equivalent Count Rate The performance of a PET scanner characterised by a figure of merit a trade-off between undesired contributions and scanner sensitivity NECR, defined as: T 2 NECR = T S kr measured as a function of activity meaningful way to compare performance of different scanners Factor k : 2 or 1 determined by whether Randoms measured in a delayed time window and subtracted or estimated from the single count rate 2D PET Acquisition Modes 3D PET Acquisition Modes 2D Mode: Interplane septa orthogonal to system axis prevent photons entering at oblique angles efficient rejection of photons scattered in the body reduce single-channel counting rate lower R rate min dead time losses Interplane septa Centre of FOV: sensitivity in 3D significantly greater 3D Mode 3D Mode: septa removed data for all possible LORs significant increase (4-8 fold) in photon sensitivity but considerable increase in S and R rates and system dead time sensitivity: a triangular function peaked at the centre of FOV 3
PET Detectors Properties of ideal PET detector: High stopping power for 511 kev photons High spatial resolution V good energy resolution to reject scatter events V high timing resolution Be inexpensive to produce Physical properties of commonly used scintillators Reproduced from D. Bailey, J.S. Karp, S. Surti, Positron Emission Tomography, Springer-Verlag, London, 2005, pp 31 Possible detectors: Proportional gas chambers Semiconductors Scintillators photo-detectors (PMTs, semiconductor-based photodiodes) Inorganic crystals emit visible light photons 4 main properties: stopping power of 511keV, signal decay time, light output, intrinsic E resolution Single Crystal mounted on individual PMT one-to-one coupling Scintillation Crystal PMT Front-end Electronics 511 kev γ-ray BGO or LSO 8x8 or 12x12 matrix of 6x6x30 mm 3 or 4x4x30mm 3 Block Detector A B γ photons convert to light photons, proportional to γ energy Light converted to electrical signal and amplified Crystal dimensions determine the spatial resolution 30mm depth, 10-30mm high (axial), 3-10mm width (in-plane) Registration and further processing of the signal y x 30 mm C y = x = 4 PMTs D (AB) (CD) ABCD (BD) (AC) ABCD Important Performance Parameters in PET Photon Sensitivity: Ability to detect coincident photons emitted from inside the FOV Determined by: scanner geometry and stopping efficiency of detectors for 511 kev Denser, higher Z, thicker (longer) detector elements to improve the 511 kev stopping power (BGO) Scanner geometry defines solid angle covering object Small diameter & large axial FOV high sensitivity scanner The higher the sensitivity the better the SNR in the reconstructed image High stopping power reduction of parallax error Important Performance Parameters in PET Spatial Resolution: Positron range effect its extent depends on the range of Es of the emitted positrons and the medium Photon Non-Colinearity worse for larger systems Size of the photon detector element pixel size (4-6 mm in typical clinical systems) Directly affects the spatial resolution in reconstructed image R sys = ( R det2 R range2 R nonc 2 ) 1/2 4
Important Performance Parameters in PET Energy and Coincidence Time Resolution: Improved by: Using crystals with brighter & faster light pulses Using low-noise photo-detectors Collecting a higher fraction of the light into the photodetector for larger electronic pulses Good E resolution narrow E window reduce S & R events without compromising sensitivity Good T resolution narrow time window reduce R events Typical values: 25% FWHM at 511 kev, 3ns FWHM Important Performance Parameters in PET Count Rate: For each signal registered a finite processing time If too many photons hit detector saturation of electronics due to pile-up of more than one pulses Pile-up depends on: scintillator decay time; effective integration time of electronics; photon event rate seen by the detector For best count rate performance: Crystals with fast decay time Detector with excellent T resolution Fast processing electronics Limited activity within the sensitive FOV Depth of Interaction (DOI) Effect A photon hits the detector travels a short distance within the detector material deposits its E in the detector the detectors used in PET do not measure this point the exact depth at which photons interact is unknown DOI within the crystal The measured position of E deposition projected to the entrance surface of the detector parallax error Parallax Error due to DOI Effect For photons at oblique angles parallax error produces significant deviations from real position blurring of reconstructed image degradation of spatial resolution DOI depends on: dimensions of the detector element, source location, scanner diameter, crystal length Annihilation photon path Assigned LOR DOI Effect A thin crystal with high stopping power reduces distance travelled by photons in the detector reduces parallax error But thin crystal reduces sensitivity An accurate measurement of DOI within crystal required Need for detectors with DOI measurement capabilities Phoswich detectors: stacking thin layers of different scintillators with different decay time on top of each other and implementing pulse shape discrimination Use photo-detectors at both ends of a thick (long) scintillator Depth sensitive detectors (e.g. CdZnTe) Ongoing research PET Data Reconstruction mathematical algorithms to calculate 3D probe distribution volume from the 2D projection data 2 basic reconstruction schemes: 1. Analytic Methods: acquisition process, measurements, reconstructed image as continuous functions (e.g. FBP) Directly compute an inverse transform formula to convert the recorded hits into an image Require spatial frequency filtering to reduce statistical noise resulting in a loss of spatial resolution Linear, more computational efficient, fast and simple to implement 5
PET Data Reconstruction 2. Iterative Methods: acquisition process, measurements, reconstructed image as discrete quantities start with an initial estimate of the 3D distribution and go through iterative modifications of that estimate until a solution is reached (e.g. MLEM) may incorporate statistical methods Appropriate for photon count limited data & PET systems with non-standard geometry Allow an improved trade-off between spatial resolution and More computationally intensive Corrections for PET Data There are several undesired physical effects inherent in the detection process of annihilation photons in PET measured PET data must be corrected for these physical factors either before or during image reconstruction Photon Attenuation within the tissue Most important correction For any given LOR, the attenuation depends on the total path travelled by the 2 annihilation photons (the total thickness of the body along that line) and is the same whether a point source (origin of the 2γ emission) is outside or inside the object Attenuation may be corrected for every LOR by measuring the total attenuation factor of an external radiation source that transmits activity through any LOR In clinical PET/CT use of X-ray CT the 511 kev attenuation coefficients determined from an appropriate scaling of attenuation coefficients measured at X-ray Es Detector Response Non-uniformity Variations of the coincidence detection efficiency between LORs affect the image uniformity in PET Variations result from imperfections related to physical, geometric, mechanical and electronic properties of individual detector elements This artefact is normalised by measuring the non-uniform response for every LOR with an external radiation source and applying this normalisation to every measured data set for correcting them Detector Dead Time or Saturation Dead time: the finite time required to process and record an event while no other events can be recorded Also includes pulse pile-up loss of counts Causes loss in spatial and contrast resolutions Saturation when the incoming photon flux is higher than the system processing bandwidth allows Dead time losses significant in high count rates with continuous or block detectors An analytic model of the dead time/saturation to calculate the correction factors that applied to every LOR Random Coincidence Events Cause loss in quantitative accuracy and contrast resolution R effects worse for high count rates Estimates of the R coincidence rate for every LOR are obtained from measurements or calculations and are subtracted or estimated from the single count rate Single count rate as a function of time is required Common technique is to measure the count rate in a delayed time window where true coincidences are impossible Then real-time subtraction for each LOR implemented in hardware 6
Scatter Coincidence Events Cause degradation of quantitative accuracy and contrast resolution Worse for larger objects, higher γ rates, poor E resolution Using a narrow E window rejection of large angle S events but small angle S events still present Small angle S events are calculated for each LOR and subtracted Most promising scatter compensation use simplified Monte Carlo simulations to estimate the scatter distribution Time consuming but already practical implementation in commercial PET scanners Isotope Decay Compensates for changes in tracer activity over time Knowledge of the half-time of the isotope as well as the time record of when each data acquired is required Partial Volume Effect Causes spatial resolution blurring Due to the finite spatial resolution of the system and the inherent sampling of discrete pixel image representation reduces intensity for structures that are on the order of the system resolution or smaller The activity concentrations of such structures either overestimated or under-estimated depending on the regional distribution of the radioactivity Correction factor by measuring the intensity reduction effects vs structure size in a phantom with known activity concentration in spheres of various known diameters Detector Developments New scintillators with fast decay time, high light output, high stopping power PS-PMT and MC-PMT Si PIN Diodes APDs SiPMs CdZnTe HpGe Si microstrip Avalanche Photo Diodes (APD) In a single package or arrays Compact size Position sensitive Large excess noise Internal signal amplification, so improved SNR High quantum efficiency Need for low-noise, fast frontend electronics Gain sensitive to small temperature variations and changes in applied bias voltage MR compatible Reproduced from T.K. Lewellen, Phys Med Biol 53 (2008) R287-R317 Silicon Photomultipliers (SiPM) Geiger-mode APD High intrinsic gain Low operating voltage Excellent timing resolution Single photoelectron sensitivity Excellent SNR Robustness V low excess noise MR compatible Low cost, wide range of pixel sizes Dead space between cells reduces overall QE Lower detection efficiency than APD at short wavelengths High dark current (cooling) 7
CdZnTe Direct conversion --- true semiconductor High density and effective Z high absorption efficiency High intrinsic spatial resolution obtained by the electrode pixellation on the front and in the depth rather than cutting crystals High energy resolution (better scatter & random rejection, multiple-isotope imaging, energyresolved CT) CdZnTe Operate in room temperature High resistivity due to wide band gap lowleakage current low-noise characteristics Small system FOV versatile, flexible, light weight and configurable detectors Compatible with MRI: operate normally with MRI acquisitions up to 7 and 9.4 T Multimodal detector ({MRI/}SPECT/PET{/CT}) Time of Flight (TOF) PET In TOF PET for each annihilation event the difference in arrival times between the 2 coincident photons is also measured Images have higher SNR than images without TOF info Scintillators with v fast timing decay, high light output and high stopping power (LSO, LaBr 3 ) V fast electronics Potential to improve the image quality in heavy patients (more attenuation and scatter) CT Non-TOF MLEM TOF MLEM Patient 1: colon cancer, 119Kg, BMI=46.5 Patient 2: abdominal cancer, 115Kg, BMI=38 Reproduced from J.S. Karp et al, J Nucl Med 2008; 49:462-470 Multimodality Imaging PET/CT PET/MR TOF PET New Developments Specific applications dedicated systems breast, brain, prostate V high resolution pre-clinical imaging systems small animal imaging 8