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by: Emily Braun II


Emily Braun II
GPA 3.82


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This 212 page Class Notes was uploaded by Emily Braun II on Wednesday September 9, 2015. The Class Notes belongs to BIOEN 508 at University of Washington taught by Staff in Fall. Since its upload, it has received 21 views. For similar materials see /class/192002/bioen-508-university-of-washington in Bioengineering at University of Washington.




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Date Created: 09/09/15
Semiconductor Detectors vs ScintillatorPMT Detectors Semiconductors are emerging technology ScintPMT systems relatively unchanged in 50 years NalTl excellent for singlephoton new scintillation materials being developed for PET lutetiumbased GSO LaBrs LaCls Direct detection in semicond permits superior energy resolution ScintPMT is indirect 2 step process with marginal quantum efficiency which limits energy resolution Puritygrowth of semicond is still a challenge ie expensive ScintPMT is well established stable technology Semicond can be finely pixilated for spatial resolution um theoretically ScintPMT pixel size cannot be made arbitrarily small 05mm 10mm lower limit thus far Note spatial resolution sensitivity tradeoffs Semicond cannot be made arbitrarily thick for photon absorption efficiency sensitivity Scintillation crystals can be arbitrarily thick Note spatial resolution lsensitivity energy resolution tradeoffs Semicond require more sensitive electronics generally and less amiable to signal multiplexing resulting in potential need of many many more processing channels Sample Spectroscopy System Hardware converted to 20 10005 of visible converted to Efccg gsrl nc rlganon Statlsilqal photons electrons uncertainties signal incoming high energy gamma ray larger current or vo age l more electrons l more scintillation photons l higher gamma energy Muftichannel Analyzer amplifier From The Essential Physics of Medical Imaging B ushberg et a1 Interaction Rate and Deadtime Nunparalymbla Gaunt Hale Pamlyzable I I I I I I I I I a 10000 20000 30000 40000 50000 Inleractian Rana Limi paralyzable nonparalyzable From The Essential Physics of Medical Imaging B ushberg et al Interactions of Photons with a Spectrometer Photoelectric Compton Photoelectric Compton Photoelectric with characteristic 5 B Na39m xray escape C W E Compton scattered photon from lead shield F Characteristic xray from lead E Lead Shield Shield F Lead Shield Dowgt A Sample Spectroscopy System Output Ideal Energy Spectrum IDEAL SPECTRUM F39HOTOPEAK E7 MULTIPLE COMPTON SCATTERING 2 z 3 D in m 0 C PHOTOELECTRIC o z m D a LL 0 DD MULTIPLE O 5 Lu COMPTON n m D sgATTERING g a 5 5 2 Z P 3 0 1 m SINGLE 2 E g m compr N m 5 o 3 SCATTERING gt o u E o I I K 4 Lu II PULSE AMPLITUDE NERGY DEPOSITED IN DETECTOR counting mode TIME m Energy Resolution Realistic Energy Spectrum 1 E 137 f Cs g 00 AE 46 F o 80 WHM E7 662 x iOO MAXIMUM HEIGHT 7 a 2 MAXIMUM HEIGHT RELATIVE NUMBER OF COUNTS I I I I I I 00 200 300 400 500 600 700 ENERGY keV Fig lIlll Calculation of FWHM energy resolution of a NalITl detector for mCs 662 keV y rays N v From Physics in Nuclear Medicine Sorenson and Phelps Sample Spectrum Cs137 Detection efficiency 32 keV vs 662 keV o o 1 662keV 32kev g E SSZkeV z E g 32keV 2 Energy keV Energy keV A Photopeak D Backscatter peak 5 Compton COHtlnuum E Barium Xray photopeak C Compton edge F Lead Xrays From The Essential Physics of Medical Imaging B ushberg et a1 Sample Spectrum Tc99m A Photopeak B Photoelectric with iodine Kshell Xray escape C Absorption of lead X rays from shield Number of Interactions Note absence of Compton continuum 20 40 60 BO 100 120 MO 160 Energy keV From The Essential Physics of Medical Imaging B ushberg et a1 Effects of Pulse Pileup I A 140 keV l I I 0 40 80 12 PHOTON ENERGY keV lquot BASELINE SHIFT 8 NUMBER OF COUNTS PULSE PILEUP 40 80 120 60 III Llnf ul nu 20 Fig llelll A WWTC spectrum at low counting rate B Spectral broadening and shift in apparent photopeak energy due to pulse pileup and baseline shift in the spectrometer ampli er of high counting rate From Physics in Nuclear Medicine Sorenson and Phelps Calibrations Energy calibration imaging systemsspectroscopy Adjust energy windows around a known photopeak Often done with longlives isotopes for convenience Cs137 Ev 662 keV close to PET 511 keV T1230yr Co57 Ey 122 keV close to T099m 140 keV T12272d Dose calibration dose calibrator Measure activity of know reference samples eg Cs 137 and Co57 Linearity measured by repeated measurements of a decaying source eg Tc99m Overview of today s lecture Emission vs Transmission Imaging Nature of nuclear radiation Isotopes used in nuclear medicine Detection methods Counting statistics Imaging systems Planar gamma scintigraphy Random Processes in Nuclear Medicine Radiation decay and detection are random processes described by probabilistic statistical distribution functions P t The means and variances of these probability distributions are part of the data analysis process in nuclear medicine Examples governed by quantitative law Radioactive decay in time Ptdt exptrdt probability atom will decay between the times t and tdt r THQI In2 TW2 halflife of isotope Gamma ray absorption in medium PXdx exp4Xdx probability 7 will be absorbed between X and XdX u attenuation coefficient of material at specified energy Examples governed by empirical law Scintillation photons created upon absorption of a gamma ray There is a mean number created per unit of absorbed energy and an associated variance for each scintillation material Have to work with law of averages gt probabilities Scintillation photon creates a photoelectron at a PMT photocathode Photocathodes have an intrinsic quantum efficiency QE the chance that any given photon creates a photoelectron is QE After observing 10005 of photons QE will have generated photoelectrons Characterizind Random Phenomena and Errors Measures of Central Tendency Mode Most Frequent Measurements not necessarily unique Median Central Value dividing data set into 2 equal parts unique term Mean Arithmetic Mean J C 12x i 11 11 Measures of Dispersion Range Difference of largest and smallest values Variance Measures dispersion around mean 1 n 0392 7240c Z n 1 1 Standard Deviation 0407 Statistical Models for Random Trials Binomial Distribution Random independent processes with two possible outcomes Poisson Distribution Simplification of binomial distribution with certain constraints Gaussian or Normal Distribution Further simplification if average number of successes is large eg gt20 Characterizing Random Errors With a Distribution Binomial Distribution Independent trials with two possible outcomes n Binomial Density Function Pbmomlv pr1pn 39 rn r Probability of r successes in n tries when p is probability of success in single trial XPn 6xlpn1p Example What is the probability of rolling a 1 on a six sided die exactly 10 times when the die is rolled for a total of 24 times r 1 n 24 p 15 Pbmmr10 00025 1 in 400 Poisson Distribution Limiting form of binomial distribution as p s o and n a co As in nuclear decay Have many many nuclei probability of decay and observation of decay very very small 139 exp PPoxEEon r I r 2 Only one parameter u Mean Variance gt X C7 Example A radioactive source is found to have a count rate of 5 countssecond What is probability of observing no counts in a period of 2 seconds 7 F 7 7 Pp Wok 710 ex m 104s410 5 00 Gaussian Normal tribution Symmetric about the mean Useful in counting statistics because distributions are approximately normal when ngt Variance and mean not necessarily equal 1 x402 P e x a in exp 202 o Poisson iGaussian PU Variance Error in Countind Photons Poisson process mean variance gt Number measured IV is best estimate of mean number for that phenomenon eg Nemitted gamma rays per unit time Nscintillation photons per absorbed gamma ray gt variance mean N gt standard deviation 0 VVariance W Relative error e in counting experiments 6 3 signaltonoise ratio SNR Nc W X N VN Relative error decreases as number of events increases Emphasizes the importance of detecting as many gamma rays as possible and the sensitivity absorption efficiency of nuclear medicine cameras This applies to individual image pixels in nuclear medicine also applies to gtltray imaging but number of photons is not limited there Also applies to energy resolution in radiation detection systems Simple Prooadation of Error Quantities of interest are often determined from several measurements prone to random error If the quantities are independent then add independent contributions to error in quadrature as follows The simplest examples are addition subtraction and multiplication by a constant If the quantities a and bare measured with known error 6a and 6 then the error in the quantities X y Z when Xab ya b Z ka k constant no error 51 5y 5 5 55mg Overview of today s lecture Emission vs Transmission Imaging Nature of nuclear radiation Isotopes used in nuclear medicine Detection methods Counting statistics Imaging systems Planar gamma scintigraphy The Planar Gamma Camera How it works Positron Emission 18F 180 0 3 o Emc2 397 511keV n p p39 n n n n l3 P 1 I 391 I P Radioactive decay unstable atomic nuclei clue to too many protons relative to the 2 mm number of neutrons decays to stable form by converting a proton to a neutron ejects a 39positron39 to conserve electric charge positron annihilates with an electron releasing two anti colinear highenergy photons Types of Photons Xray photons gamma y ray photons Greek letters for radiation from nuclear decay processes annihilation photons all can have the same energy Molecular Imaging Glu Metabolism 18Ffluorodeoxyglucose FDG glucose FDG HOCH2 gylcolysis glucose 639 FDG 6 anaerobicw phosphate phosphate inef cient i l 0 18 H F pyruvateltlactate i radioactive uo ne TCA what oxidative we ef cient FDG6PO4 is trapped and is a see good marker for glucose metabolic rates How it works Scintillation optical photons 1eV high energy 2 511 keV photon gt pillsreettgr h UV E eSlicizoton detected 4 l gt A K x X l scintillator photomultiplier eg BGO Dense tubes PMTS yet transparent galn of 106 Scintillators used in PET Scanners Materi Cost Effective number Effective Decay Comments al of scintillation Density time us photons 511 determines determine keV determines scanner 5 energy and spatial sensitivity deadtime resolution and randoms NalTl cheap highest lowest long Hygroscop relativel ic y BGO expensiv lowest highest long workhorse e LSO more high high very new expensiv short technology e GSO more very high somewhat very new expensiv lower than short technology e LSO How it works Timing coincidence scanner FCM 3 e39 gt annihilation V record At lt 10 ns 1 pOSitron decay A event Typical PET Scanner Detector Ring PhotomulliplierTu Assembled Together Eight Detector Units Form Detector Form Detector Q3 um mI 38mm 3 Face of Crystal Block 38 mm x 38 mm Crystal Dimensions 47mm x 63 mm Depth of Crystals 30 mm Detector Ring Diameter 8862 mm crystal to crystal Anatomy PET gantry Detector PMT assem blies Typical PET Image Elevated uptake of FDG related to metabolism 39 THAN IAL SAGITTAL DQHDNAL ti 1 Lung cancer example Very obvious What is Attenuation The single most important physical effect in PET imaging The number of detected photons is significantly reduced compared to the number of positron decays in a spatiallydependent manner For PET it is due to Compton scatter out of the detector ring For CT it is a combination of Compton scatter and photoelectric absorption scanner one 511 keV photon scattered out of scanner one 511 keV photon absorbed Simulation of the Effects of not Performing Attenuation Correction of PET Emission Image V atquot Y 39 Locally Enhanced skin increased Reduced interior contrast even neg True PET image PET image without simulation of abdomen attenuation correction Effects of Attenuation Patient Study reduced mediastinal uptake I 39hotlungs i i Non uniform liver Enhanced i Askln uptake 1 6 I 39039 i r PET without PET with attenuation CT image accurate attenuation correction correction accurate Attenuation Correction Transmission scanning with an external photon source is used for attenuation correction of the emission scan The fraction absorbed in a transmission scan along the same line of response LOR can be used to correct the emission scan data The transmission scan can also be used to form a 39transmission39 or 39attenuation39 image same line of response hoton source LOR Ls6 fps p vix rotation a W x i H t ii 1 3 In rkw f XX qu gkx scanner FOV 71 1quotquot Emission scan EM Transmission TX PET Transmission imaging annihilation photon imaging 68Ge68Ga g tag scanner source 7 5 f 511 keV near5ide annihilation detectors photon Using 3point coincidences we can reject TX scatter uxy is measured at needed value of 511 keV nearside detectors however suffer from deadtime due to high countrates so we have to limit the source strength particularly in 3D And if you have PETCT scanner Xray TX orbiting X Xray ray tube and detectors detector at l 1 assembly v 4 ill l W 1 ll i 30 130 keV Xray photon Photon flux is very high so very low noise Greatly improved contrast at lower photon energies Scatter and beamhardening can introduce bias uxyE is measured as an weighted average from 30120 keV so u Xy511keV must be calculated potentially introducing bias Xray and Annihilation Photon Transmission Imaging Xray 30120 keV PET Transmission 511 keV Low noise Noisy Fast Slow Not a physical quantity Linear attenuation coeffcient at 511 keV 100000 75000 Energy spectra I 50000 25000 0 O 100 200 300 400 500 600 keV Quantitative errors in measurement no LOR incorrectly determined LORS scatter Lost attenuated Scattered coincidence Random coincidence event event event 3D versus 2D PET imaging A detected absorbed detected detected 2D Emission Scan 3D Emission Scan fewer true scattered and more true scattered and random coincidences random coincidences Effect of random coincidence corrections in 2D and 3D FOV for random coincidences 2D Emission Scan 3D Emission Scan Noise Equivalent Counts or NEC NEC 39Effective39 count rate Prompt coincidences are what the scanner sees PTSR but true coincidences are what we want TPSR which adds noise T SNR2 oc NBC 2 Types of coincident events 500 Clinical Prompt detected Range In this measured Trues 400 Randoms example at 10 kBqcc ScEatEer about 6 mCi the m 3007 N C scanner39s count rate for g coincidences will be 200 450 kcps but the effective count rate 100 aka NEC Will be only 2 75 kcps 5 10 15 20 25 Act Con kBqcc NEC comparisons Major arguing point for some vendors Determined partly by detectortype detector and scanner geometry acquisition mode and frontend electronics Important but not sole factor for image quality NEC ates Range for wholebody scans r kcps Partial Volume Effect AQQarent SUV drops with volume Also effected by image smoothing iFillable spheres Final Image PET Resolution Losses O O O 0 Q A 4 true tracer uptake reconstructed values scanner resolution smoothing of noisy data 1 Simulation study with OB typical imaging protocols g Limits quantitation in 306 7 oncology imaging a 04 7 important for following 02 therapy if size changes m 0 2O 25 30 35 Sphere diam mm 0 TI L O 4 L 01 Time of Flight TOF PETCT Uses difference in photon detection times to guess at tracer emission point without timing info emission point could be anywhere along line C 3x10100ms so At 600 ps Ad 10 cm in resolution best guess about location d V Ad o At2 3 e39 annihilation Philips Gemini TF PET scanner LYSO 4X4x22 mm3 28338 crystals 420 PMTs 70cm bore 18 cm axial FOV CT scanner Brilliance 16slice Installation at UPenn Nov 05 Validation and research patient imaging Nov 05 Apr 06 50 patients Beta testing and upgrade to production release software May 06 Jun 06 40 patients to date Heavyweight patient study Colon cancer 13 mCi 2 hr postinj 9 kg 3 minbed BMI 465 nonTOF A MT 1 Improvement in lesion detectability with TOF Administrivia Evaluation 0 For the three guest lecturers 0 Not the evaluation for the course or me there will be a separate evaluation for that later Class Field Trip Saturday Dec 2nd 0 Meet at UWMC Main Entrance at Noon in main hospital lobby if you 39miss the boat39 page me at 5404950 should take 1 to 15 hours depending o No Class Nov 22nd 0 Final report due Nov 29th c Lecturer on Nov 29th will be Adam Alessio please read chps 9 and 10 0 Exam 2 Nov 29th 0 Class Presentation Dec 6th Lecture 7 Magnetic Resonance Imaging MRI Spin density T2weighted image 0 We39ve spent a lot of time thinking about images so how are these three MRI images different From each other From images from other modalities CT PETSPECT US 1 2 3 4 5 6 7 What is MRI The overall picture Put subject in big steadystate 39main39 magnetic field BO hydrogen nuclei 39line up39 on average with magnetic field Adjust magnetic field in a spatiallyvarying manner using 39gradient39 electromagnetic coils Transmit radio waves into subject at the resonance frequency of nuclei using radiofrequency RF coils where resonance frequency depends on magnetic field thus location Turn off radio wave transmitter so nuclei realign with magnetic field and transmit at local frequency which depends on location Store received radio wave data vs time Repeat at steps 26 many times BANG BANG Process raw data to reconstruct images of hydrogen nuclei density Necessary Equipment for MRI Gradient Coil RF Coil Main Magnet 1 Tesla T 10000 Gauss Earth s magnetic field 05 Gauss mail minimum 3 Tesla 3 x 10000 05 60000 X Earth s magnetic field Angular Momentum Classical Mechanics Angular momentum describes rotational motion of a body 1 Orbital Earth orbits the sun 2 Spinning Earth spins on its axis giving us night and day Nuclear Angular Momentum 1 Protons and neutrons orbit within nucleus whole nucleus spinning 2 Protons and neutrons spin gt Together produce nuclear spin Spinning charge produces magnetic dipole same as electromagnet sort of The Reality Quantum Mechanics 0 Quantum Mechanics in sum Energy states exist in discrete amounts quantum Each nucleus in external magnetic field BO has energy described by the equation gtilt eq610iubook E Z YBO 39 r 5 H i Ba Eis Energy inJoules quot a m jj1j 1j 12 or 12 for hydrogen j is spinnumber 1 hZTE h is Planck S constant 6 62x103934Joulesseconds y is constant gyromagnetic ratio of atom in MHzT BU is external eld in T Tesla Nucleus can only exist at discrete energy levels for a given external field Quantum Mechanics Continued Nuclei Unpaired Unpaired Net Spin 7 1H prolong 6 1ng 172 461213 0 Hydrogen influenced by 1H 1 1 1 4109 external magnetic field Most 12C 0 0 0 common atom in body 12C 0 1 112 6729 0 Common carbon and oxygen 1m 1 1 1 1935 isotopes are not influenced by 10 0 0 0 magnetic field In presence of magnetic field two energy levels are established for Hydrogen Spin up or Spin down quantized energy states in external magnetic field known as Zeeman effect 1 E h B 2 7 0 AE Magnetic field BO 2 0Tesla E 1 h B 2 7 0 Magnetic field BO Precession Effect No Gravity ND Ba Atom like a dreidle top Small Gravity Small Ba Large Gravity Large Ba Fae D any quot Process ion Spinning nuclei wobble or precess at a rate Precession wo VBO 27T Applied Magnetic where w0 IS the precessnonal Larmor or Field resonance frequency G 30 Resonance frequency is proportional to B0 Vector Form 0 Magnetic field B0 causes M to rotate or precess about the direction of B at a frequency proportional to the size of B for hydrogen 42 million times per second 42 MHz per Tesla of B Mz stays same Working with a lot of nuclei M 0 Small B0 produces small net magnetization M 9 Thermal motions try to randomize alignment of nuclei magnets M 0 Larger B0 produces larger net magnetization M lined up with BO 9 Reality check 00003 of nuclei aligned per Tesla of BO Making M not parallel to B 0 Now have a majority of nuclei in line with large magnetic field each precessing at its resonance frequency 0 Basic Idea We want to perturb the nuclei and see how long it takes until they realign o A way that does not work BO 30431 Turn on a second big magnetic field B1 perpendicular to main BO for a few seconds M would drift over to vector sum B1 of B0 and B1 Then turn B1 off M is now not parallel to magnetic field BO 0 This fails because cannot turn huge Tesla magnetic fields on and off quickly But it contains the kernel of the necessary idea A magnetic field B1 perpendicular to B0 Making M not parallel to B Mechanical Analogy Swingset Person sitting on swing at rest is aligned with externally imposed force field gravity 0 To get the person up high you could simply supply enough force to overcome gravity and lift him and the swing up Analogous to forcing M over by turning on a huge static B1 c The other way is to push back and forth with a tiny force synchronously with the natural oscillations of the swing Analogous to using the tiny RF B1 to slowly flip M over Apply force in Resonance Making M not parallel to B Excitation with Radiofrequency RF B1 is excitation RF field Apply B1 so fluctuates at the resonance frequency and points perpendicular to B0 Cl w k RF energy is absorbed An observer in the I quot surrounding laboratory will see Mia spiral down to the XY plane or even to the Z axis x An observer riding on the Mo vector sees the external world rotating about him Mia then seems to tip 0c towards the Y39 axis Making M not parallel to B Excitation with Radiofrequency RF Varying frame of reference B1 B0 Magnetization M0 RF energy is absorbed An observer in the An obsener riding on the 8 vector sees the surrounding laboratory will see Mo spiral down to the external world rotating about him Mo then seems XY plane or even to the Z axis to tip on towards th m39v axisps stmxdedmb ammomr Common RF Pulses M begins along 2 axis 900 pulse M lines up with y axis M OIMOIO OL 90 Quantum Mech Both energy levels are occupied by same number of spins 1800 pulse M lines up with z axis M 00 M0 OL 180 Quantum Mechmajority of spins occupy highest energy level Turn RF Off Precessing spins cause a change in flux CD in a transverse receive coil Flux change induces a voltage across the coil This is the quotNMRquot signal At the resonance frequency The frequency of this precession is proportional to the applied magnetic Bo field Signal proportional to Proton Density Laboraory Frame J Turn RF Off Mz begins to recover Exponential recovery of M2 Time constant is called T1 Longitudinal or SpinLattice Relaxation Spins Equilibrium Equal numbers parallel Reversion to Equilibrium and antiparallel equilibrium was saws am sad 015 de 59 5 900 I 100 quot56 O MZ 630 0 t0 t T1 Time gt Turn RF Off SpinLattice Relaxation T1 Immediately after RF signal M0 cosa After long time returns to this M0 HOW I t M1t M0 cosoze1 M0 1 e71 T1 Spinlattice Relaxation Water T1 0 500 1000 1500 2000 2500 Timems RF Off SpinSpin Relaxation T2 Spins Mxy begin to dephase due Exponential decay of signal Time constant is called T2 or T2 Transverse or SpinSpin Relaxation Mxy Mxy Mxy maxumum decreasrng zero C Mxy 100 I k 37 A 0 tT2 Time 90 pulse CASE 1 Slows Dawn Speeds Up amp CASE 2 u btr acts SJ btr acts Beth spins slow down as theyr move together RF Off SpinSpin Relaxation T2 relaxes as 2 MN M 0eT 2 Eq 623 in text Transverse Magnitude o T2 SpinSpin Relaxation 120 H O O 80 Water T2 1500ms Fat T2 50ms 0 500 1000 Timems 1500 2000 2500 Relaxation Summary L I time time http WWW7II11 SI 1 Stanford eduNbrianintromr Why does MXy decrease Why does MZ increase Current Status Have a big magnet B0 Forces all same atoms to precess at same rate Larmor Frequency Can perturb atoms with RF signal and measure precession and relaxation Imaging with NMR 0 Slice selection Magnetic Gradient coils provide a linear variation in B2 with position 0 1973 Paul Lauterbur Nobel Prize Laureate in 2003 Result is a resonant frequency variation with position B Z l l l in a b POSIthl l o o W 1090 Gzzgt 27r Protons at 22a Will process at different frequency than those at 22b Position Magnitude Exciting a Slice m Frequency RF Amplitude Time httpwawemrsrlStanfordeduNbrianintromr Image Acquisition o Gradient causes resonant frequency to vary with position 0 Receive sum of signals from each spin Signai Frequency 0 o H O quot v o39 a Slgna2 39i If I H s l quotu x l I l Position T0tal 0 O 0 0 0 Image Reconstruction 0 Received signal is a sum of tones o The tones of the signal are the image 0 This also applies to 2D and 3D images Founer Transform ltZgt Image Received Signal MRI Readout imaging Next we apply a second gradient field in a perpendicular direction say the xdirection to force some change in the signal across the plane otherwise we can t tell where the signals are coming from in the plane So now the precession frequency of the hydrogen nuclei depend on the xlocation wow 2 yGxxB We can then describe the motion of the transverse component which generates the RF signal output using phasor notation as M trx y t M trx y 0e 7Gx The RF signal output is the integration of all the signals in the slice st T Tpxye clx dxdy oo oo proton density what we want to know Fourier transforms in MRI The ktheorem o If we define kx nytZn39 we can write the RF output signal as st T Tpxyequot 2 quot dxdy oo oo 0 Now recalling that Skxky 1quotsxy j jsxye quotx quoty dxdy oo oo we see that our signal is part of the Fourier transform of pxy ykp0 To sample the other parts of kspace the Fourier transform of our object of interest we apply a predetermined gradient in the perpendicular y direction prior to the readout gradient GX this gives us 2 I l 717kyyklx stfy J JPWYV dxdy lt1 MRI equation oo oo where ky 30ny 2717 2D Imaging Sequence ky Bare bones MRI pulse sequence 90 degree ip selective excitation RF c0IS RF t zslice selection 62 Gradient coils Gy ty phase encoding Gy is changed with each repetition l39 R to sample k space along different ky GX m values causes banging ktheorem or N N MRI equation 30 I I px yexp 2 kyykxxdXdy kxkyGXtGyty 2D Image Reconstruction 0 So to image the patient slice remember this is tomographic imaging we sample out all of kspace and compute the inverse Fourier transform Freq uencyspace k space Image space KSpace Illa Mnna Lisa in kSpace ksnace if r quotquot L quot ky kx N Low Frequency Mona mm mm Mm Source Traveler s Guide to K sgce CA Mistretfa Resolution 0 A lot of MRI development is to figure out how we can more efficiently andor more accurately sample kspace 0 Image resolution increases as higher spatial frequencies are acquired Takes time to sample more of kspace Review l Tissue protons align with magnetic eld lt J equilibrium state RF pulses E l Relaxation PrOtons absorb Spatial encoding processes RFenergy using magnetic EXCIteE state field gradients I I Relaxation i processes l I I I Protons emit RF energy lt return to equilibrium state NMR signal detection 4 RAW DATA MATRIX Fourier transform IMAGE Dephasing Phenomena The bulk magnetization vector M0 macroscopic group of spins has two components longitudinal M2 and transverse Mtr or Mxy which actually generates the signal Both of these components change separately with time due to physical effects In the rotating frame these are given by the Bloch equations dMZ MZ M0 dt T1 dMn dt T 2 With a 90 degree flip the resulting time behavior is given by Mzt M01 e m Mtrt Moe tTZ SpinLattice Relaxation T1 2 M20 MOO e T1 Long udmalMagn ude H N O H O O 00 O m 0 4 o N O O 500 T1 Spinlattice Relaxation Fat T1 100ms Water T1 2000ms 1000 1500 2000 Timems 2500 SpinSpin Relaxation T2 2 M t M 0eT 2 Eq 623 in text Transverse Magnitude 00 120 H O O 00 O 0 1 O 4 o N O 0 T2 SpinSpin Relaxation Water T2 1500ms Fat T2 50ms 500 1000 1500 2000 Timems 2500 Additional T2 effects T2 Two factors contribute to the decay of transverse magnetization 0 Molecular interactions said to lead to a pure T2 molecular effect 0 Variations in Bo said to lead to an inhomogeneous T2 effect The combination of these two factors is what actually results in the decay of transverse magnetization The combined time constant is called T2 star and is given the symbol T2 The relationship between the T2 from molecular processes and that from inhomogeneities in the magnetic field is 1T2 1T21T2inhomo Undo effect of T2 with SpinEcho Imaging Spin Echo 0 Apply 90 degree pulse to start signal 0 Apply 180 degree pulse to invert direction of dephasing spins gtinduces an echo to reform In opposite direction 7 4 l l pse TE2 180 TE 9 l l pulse l I l C9 9 e e 6 l l l Signal gradually increases gt Signal gradually decays with Spin echo decay rate T2 amplitude depends on T2 Spin Echo SpinEcho T2 Decay T2 decay SpinEcho Pulse Sequence Two Parameters to vary 1 TE 2 TR Repeat sequence time RF pulse Slice select w E E 4 90 pulse 180 pulse Phase encode Frequency encode r L Signal out WWW V TE 4 V TR SpinEcho Pulse Sequence Basic signal output with a 90 degree spinecho pulse sequence TR TE Pxay1 CXP EDCXP Ej Adjust image qualities by changing TR and TE all values in msec T1 Weighting T2 Weighting Spin Density short TR and long TR and long WEithing short TE TE long TR and short TE TR 400600 15003000 15003000 TE 1030 60150 1030 SpinEcho Pulse Sequence What Parameters should we set for T1 Weighting TR Tl differences in longitudinal magnetization minimized because not enough time for everything to return to equilibrium Short TE So T2 decay effects minimized T2 Weighting Long TR reduce Tl effects has time to return to equilibrium TE T2 allow for T2 decay to be emphasized between tissue Spin Proton Density Weighting Long TR reduce Tl effects has time to return to equilibrium Short TE keep signal high and reduce effects of T2 Examples of Different Weightings in SpinEcho 2DFI39 MRI Spin H proton density T1 weighted image T2weighted image long TR short TE TRT1 short TE Long TR TENT2 x l ex ex p y p T1 p T2 T1 weighted spinecho ZDFF MRI is the most common form of MRI Again this is an example of 1 the differences between the data world and the visual world and 2 that even the data representation does not correspond to a single physical property Examples of Different Weightings in MRI T1 weighted T2 weighted Exampies i iiferent iiiweightings in MRE spin H proton density weighted T2 weighted Aspects of MRI we won t have time to discuss Flow imaging functional brain activation imaging fMRI Angiography The use of dynamic contrastenhanced MRI Bioengineering 508 Physical Aspects of Medical Imaging 1 I Iquot quot12 ht39tln39Irnursps 39 39 1 Organizer Paul Kinahan PhD Adam Alessio PhD Ruth Schmitz PhD Lawrence MacDonald PhD Imaging Research Laboratory httpdeptswashingtonedunucmedIRL Department of Radiology University of Washington Medical Center Alessio BIO508 Bioengineering 508 Physical Aspects of Medical Imaging Introduction to Medical Imaging 1 Medical Imaging Modalities 2 Modern Image Generation 3 Intro to Image Quality Adam Alessio PhD Department of Radiology University of Washington Medical Center aalessiouwashingtonedu Alessio BIO508 Nature of Medical Imaging For this class Medical Imaging Noninvasive imaging of internal organs tissues bones etc Focus on 1 Macroscopic not microscopic in vivo in the body not in Vitro in glass in the lab Primarily human studies Primarily clinical diagnostic applications bfDN Alessio BIO508 Nature of Medical Imaging QUICK CAVEAT Powerpoint Slides are just a vehicle for major topics These do not have all the information discussed in class Taking notes to supplement slides is probably a good idea Alessio BIO508 Types of Medical Imaging Modalities Grouped by underlying physics XRaleT Ultrasound Major 4 that dominate Magnetic Resonance Imaging MRI Elfirgrigcigigrg focus Nuclear Medicine Optical 4 Primarily microscopic Magnetic Field Electric Field Therma Mainly research based Optoacoustic Elastography Alessio BIO508 Types of Medical Imaging Modalities Electromagnetic Spectrum MI MAM Magnetic Resonance Optical XRay Ultra Sound Tomography Mammography Wavelength 103 102 0392 1039 10 m m7 10 10 9 in meters v xr lt SWe 039 longer wavelength Common name of wave 1 LTRAVIOLE I soF x RAV Sources Frequency one photon electron volts 5 10quot 10 7 1O 39 1t at 10399 104 102 103 i7 For comparison this is wavelengthfrequency range of US but US is NOT 39 Alessio BIO508 Types of Medical Imaging Modalities Classifications of Medical Images 1 Anatomical vs Functional AnatomyStructureFeatures vs Physiology 2 Emission vs Transmission Where does energy imaged originate 3 Projection vs Tomographic Projectiongt 2D imaging single plane no depth information Tomographic tomoquot slice graphyimage gt volumetric Alessio BIO508 Modern Image Generation From continuous real world to a meaningful image on computer 1 Sampling Continuous Information Information and sampling technique varies widely for each modality Topic for later lectures Computer can only hold discrete chunks of data Pixel a single picture element Voxel a single volume element Quantizing Samples Each discrete chunk must be represented by certain number of bits Visualization Techniques of quantized sampled image volumes N a Alessio BIO508 1 Sampling Continuous Information Given a signal such as a sine wave with frequency 1 Hz Alessiu r ElOSUB Intro to Sampling Theory We can sample the points at a uniform rate of 3 Hz and reconstruct the signal Alessiu r ElOSUB Intro to Sampling Theory We can also sample the signal at a slower rate of 2 Hz and still accurately reconstruct the signal Alessiu r ElOSUB Intro to Sampling Theory However if we sample below 2 Hz we don t have enough information to reconstruct the signal and in fact we may construct a different signal an alias Alessiu r ElOSUB Intro to Sampling Theory Intro to Sampling Theory Aliasing occurs when your sampling rate is not high enough to capture the amount of detail in your image Can give you the wrong signalimage an alias Where can it happen in graphics To perform sampling correctly In Image space need During image synthesis to understand structure of dataimage sampl39 g In continuous signalinto discrete signal u eg ray racingv line drawingvmnmon planing m Fourier Any periodc function can be rewritten as a weighted During image processing sum of sines and cosines of different frequencies Fourier resampling discrete signal at a different rate Series eg Image warping zooming inzooming out etc Nyquist criterion Must sample at two times the highest frequency in the signal for the samples to uniquely de ne the given signal Samplinngz Fm Sampling below the Nyquist frequency can cause aliasing CD sampling example 1 f samvles a39llased uuipui signal i 1 L Lsampimgpici analog input signal Alessio r BlO508 Alessio r BlO508 A sum of sines Fourier Transform Signal fx Our building block A smalx Add enough ofthem to get any signal fx you want Which one encodes the coarse vs fine structure of the signal 1D Example A signal composed of two sine waves with frequency 2 Hz and 50 Hz The Fourier Transform of the signal shows these two frequencies Fourier Transform of fx In 2D Usually represent low M l What would an image look f requencIes near orIgIn high like With a IOt Of high frequencies away from origin frequency content ll 0 so WI m 39 What COUId you do to reduce High Freq High Freq frequency speckled noise from an image High Freq High Freq ftarget f1 f2 f3 f Alessio r BlO508 Alessio r BlO508 2D Fourier Transforms t Mmquemmm mavem mew mm WEE HPEEWDWquot mmmdeattveuen camanem tamannudeattveuen 2D Fourier Transforms t Mmquemmm mavem mew mm Gamma WEE HPEEWDWquot mmmdeattveuen camanem nnudeattve n Gamma Ongmat After lowpass After highpass Atesstu r EtOSUE Atesstu r EtOSUE Frequency Content Frequency Content Atesstu r EtOSUE Atesstu r EtOSUE Modern Image Generation From continuous real world to a meaningful image on computer Sampling Continuous Information Information and sampling technique varies widely for each modality Topic for later lectures C mputer can only hold discrete chunks of data Pixel a single picture element Voxel a single volume element N Quantizing Samples Each discrete chunk must be represented by certain number of bits Fquot Visualization Techniques ofquantized sampled image olumes Alessin r ElOSUB 2 Quantization Only have finite storage available for each picture element Digital images have digitized intensity values Continuous values are quantized into discrete values Example Truecolor on computer displays use 24 bits for each pixel 8bits blue 8 bits red 8bits green256x256x256 possible colors Many medical imaging modalities use intensity values of 1 2 bits per pixel 2quot124096 possible gray levels Alessin r ElOSUB Color depth a pits per pixel 5 pits per pixel 4 mg per pixel 2 pits per pixel Alessin r ElOSUB Overview of today s lecture Nature of nuclear radiation Isotopes used in nucI med Detection methods Counting statistics Imaging systems Planar gamma scintigraphy The Planar Gamma Camera Gamma Camera Instrumentation Typical Gamma Camera acquisition Parameters NalTl crystal 500m X 300m Prggefsmg PMTs75 cm 3quot lSp ay 30 50 PMTs tota EleCtromcs Collimators holes hex 2 6mm boards lei crystal collimator Crystal and light guide Light Guide N 38 thick 95 mm Crystal NaITI Density 867 gcm3 Attenuation Coefficient p 140 keV 264 cmquot gt 1e2640mllo950m 92 PE fraction 80 Light output 40keV gt 40140 5600 scint photons Decaytime 230 nsec Wavelength 410 nm Light guide distributes scintillation light over PMT array Light response function versus position light sharing gt spatial resolution E xi E i Intrinsic spatial 55 1 Resolution 2E1 lt 4 mm FWHM i lt PMT size MIA 1 1 i I I I d I I I I I L XV pMI S PMT signals Ei LG gt Crystal 7 absorption creates 1000s of scintillation photons Spatial Positioning I Corrected Posmon pOSItIon and and ADCs allergy energy Signals I gt1 Signals 3 Pulses Digital from W Y I I Y d d I gt pOSItIon Dlglta InPIRATua 2 Circuit energy S 39 X and X computer spatial linearity Digital LUIIcc Iiun summing Z circuits Z circuit enerQY Analog J Digital A FIGURE 215 Electronic circuits of a modern digital scintillation camera From The E ential Phy ic of Medical Imaging Bu hberg et al Gamma Camera Energy Spectra Summed signal from all PMTs 139 Source Scattered events have changed Source behind I in air direction hence they will be 10 cm water I mis positioned by the image l Counts generation algorithm gt this tends to diffuse sources and reduce image contrast Energy Energy Windows 0 Balance between accepting all good events importance of sensitivity and rejecting scattered events 0 Most gamma cameras can acquire data using multiple energy windows Allows for simultaneous imaging of different radioisotopes for example Tc99m 140 keV and 1131 364 keV Collimators Septal Penetration Collimator l septa p 67 Minimum septa thickness t 2 H for lt5 septal penetration l Collimator Efficiency Collimators typically absorb well over 9995 of all photons incident on them Tradeoff between spatial resolution small collimator holes and detection ef ciency large collimator holes Hexagonal holes good symmetry good packing fraction foil fabrication gt 26 are double walls Collimator Resolution Collimator line spread function projected radiation profile mV Am Line source FIGURE 2112 Line spread function LSF of a parallelhole collimator as a function of source tocollimator distance The fuIlwidthathalfmaximum FWHM of the LSF increases linearly with distance from the source to the collimator however the total area under the LSF photon fluence through the collimator decreases very little with source to collimator distance in both figures the line source is seen quotendonquot From The E ential Phy ic of Medical Imaging Bu hberg et al Gamma Camera spatial resolution 16 we wequot h6 6 a 14 VW X A 9 E v 4 3 12 39 2 x E 2 2 RS RIR 9 c 5 B 2 08 E 2 gtquot 06 v 0 39 I x 39f 04 39 Typical r39 organ depths o 2 gt 00 0 2 4 a 10 12 14 16 Sourcelocollimalor dlslance cm Image in crystal Image in crystal llllllllllli Parallel hOIe Pinhole Object Object Converging Image in cwstal Diverging Image in crisial Object Object 20 250 Diverging 16 200 Parallelhole 12 150 System resolution mm Relative geometric e iciency B 100 Pinhole 4 g Converging 50 o I I I 1 o 0 5 10 15 20 A B Collimator Resolution and Sensitivity Converging Pin hole I Parallelhole Diverging 5 Sourcetocolllmator distance cm 10 15 20 Figure I4 2I Performance characteristics A system resolution B pointsource geometric ef ciency in air versus sourcetmcollimator distance for four different types of gamma camera Lollimutorst lReprinted by permission of the Society of Nuclear Medicine from Meyer RA A Iow energy multihole converging collimator compared with a pinhole collimatort J Nucl Med 1559 64 1974 From Physics in Nuclear Medicine Cherry Sorenson and Phelps Collimator Resolution and Sensitivity TABLE 213 THE EFFECT OF INCREASING COLLIMATORTOOBJECT DISTANCE ON COLLIMATOR PERFORMANCE PARAMETERS Collimator Spatial resolution3 Efficiency Field size Magnification Parallel hole Decreases Approximately Constant Constant m 10 constant Converging Decreases Increases Decreases Increases m gt1 at collimator surface Diverging Decreases Decreases Increases Decreases m lt1 at collimator surface Pinhole Decreases Decreases Increases Decreases m largest near pinhole 3Spatial resolution corrected for magnification From The E ential Phy ic of Medical Imaging Bu hberg et al The Scintillation Camera Corrections and QA Gamma Camera Processing Electronics energy correction Energy channel vs event location 6 1n4 I 1 I 50 100 1 U 2 0 energy keV Gamma Camera Processing Electronics with and without energy correction Gamma Camera Processing Electronics linearity correction Barrel distortion Pincushion distortion From Physics in Nuclear Medicine Cherry Sorerrson and Phelps Gamma Camera Processing Electronics linearity correction ltb ill 0304 alooQID Additional Gamma Camera Corrections sensitivity uniformity Acquired from long uniform flood after energy and linearity corrections have been applied Multiplicative correction Adjusts for slight variation in the detection efficiency of the crystal Compensates for small defects or damage to the collimator Should not be used to correct for large irregularities Daily Gamma Camera QA Tests Flood uniformity Photopeak window From The E ential Phy ic of Medical Imaging Bu hberg et al Multienergy spatial registration eg Ga 67 93 185 and 300 keV gamma rays properly adjusted improperly adjusted From The E ential Phy ic of Medical Imaging Bu hberg et al 11 Pulse Pileup A A 140 keV 4O 80 120 PHOTON ENERGY keVl BASELINE SHIFT 39 A B NUMBER OF COUNTS l i t L O 40 80 12O 60 WINDOW k 20 a Figi 1110 A 9939quotTc spectrum at low counting rate B Spectral broadening and shift in apparent photopeak energy due to pulse pileup and baseline shift in the spectrometer ampli er at high counting rate Source Pileup Sgurce Figure l46 lmnges if two quotwwl39c point sources ul39 relatively high activities i3 MHq each Events Appearing in thv hand between the two pointsource locations are mispnsitimietl events luv tn pulse pileup Pileup in image From Physits in Nuclear Medicine Sorenson and Phelps and Cherry Sorenson and Phelps Image Acquisition Frame mode data stored as an image static single image acquisition can have multiple energy windows dynamic series of images acquired sequentially gated repetitive dynamic imaging used for cardiac imaging Listmode data stored event by event time stamps are included within data stream allows for flexible postacquisition binning can result in very large data files 12 Region of Interest ROI and TimeActivity Curves TAC Quantitative Results mun m dnuy mush Kidnnu 1599 1433 xzaa teen see 666 486 259 FIGURE 21Z4t Regions Ul 39 From The Essential Physics of Medical Imaging B ushberg et al Example Clinical Images l 1 i 39 7 To evaluate the hyperparathyroidism double phase technetium99m sestamibi parathyroid scintigraphy was performed Parathyroid Scintigraphy was performed 20 minutes and 2 hours after injection of technetium99msestamibi The 20 minute scan showed uptake in a normal appearing thyroid gland as well as uptake in two ovoid areas in the upper mediastinum The 2 hour image showed wash out of activity from the thyroid and persistence of activity in the upper mediastinum 13 Example Clinical Images 7 x L n Collimator artifacts 7 I from high energy gammas quot 364 keV l L R 39 1 O 39 y K g 39 Q 0 ml uptake in primary differentiated thyroid carcinoma arrow and in rib and pelvic metastases arrowheads MT east 99mTcMDP bone scintigraphy demonstrating multifocal increased uptake due to skeletal metastases from a renal carcinoma note right nephrectomy Example Clinical Images 99mTcMIBI scintimammography supine and prone left lateral views showing a primary tumor in the left breast arrow and axillary lymph node metastases arrowhead Example Clinical Images whole body renal excretion gt55 mm x mm 5 xx wow quot I C 939 1 z 3 as gt y quotm a 1 B 397 I C I f I 5 G M 13 5 QQWTC 201T Typical PET Image Elevated uptake of FDG related to metabolism DDHDNAL 39 THAN IAL SAGITTAL a L m g 2 Lung cancer example But where exactly is it located PETCT Oncology Imaging Anatometabolic fusion images are useful in the management of patients with cancer Wahl JNM 1993 PETCT scanners are used to provide accurately aligned functional and anatomical information Beyer JNM 2000 recurrent thyroid cancer localized to the right retropharyngeal space A secondary synergy of PETCT scanners is to use the CT image for attenuation correction ofthe PET emission data Kinahan Med Phys 1998 lownoise attenuation correction factors no transmission scan shorter total scan time o no bias from emission contamination of post injection transmission scans Growth of PET procedures in the US 1000000 100 Procedures 800000 80 I PETCT of Sales m g 600000 60 395 8 2 400000 40 a 200000 20 0 0 1998 1999 2000 2001 2002 2003 2004 1998 Reimbursement for FDGPET 1st PETCT prototype built The number of procedures has been doubling every 19 months Over 90 are FDG cancer imaging for diagnosis and staging Recent figures indicate 40 annual growth in number of procedures 1998 Pittsburgh PETCT prototype PET Fused image viewer 2006 Six Commercial PETCT Scanners All rely on CTbased attenuation correction Siemens Phillips General Electric Biograph Pico and Hires Gemini GXL and TF Discovery ST and DSTE LSO GSO LYSO Imaging FDG uptake PET with anatomical localization CT Thyroid cancer example m Anatomy Function FunctionAnatomy Improved Integration of PET and CT T t 39 i 4 l v 39 5 quot f 139 r gt i l V I 3 Scanners now support listmode flexible protocols and improved display facilities Basic PETCT Architecture coscan length lt Helical Stationary CT PET Detectors patient bed patient port lt gt axial separation Attenuation Correction Transmission scanning with an external photon source is used for attenuation correction of the emission scan The fraction absorbed in a transmission scan along the same line of response LOR can be used to correct the emission scan data The transmission scan can also be used to form a 39transmission39 or 39attenuation39 image same line of response hoton source LOR Ls6 fps p vix rotation a W x i H t ii 1 3 In rkw f XX qu gkx scanner FOV 71 1quotquot Emission scan EM Transmission TX And if you have PETCT scanner Xray TX orbiting X Xray ray tube and detectors detector at l 1 assembly v 4 ill l W 1 ll i 30 130 keV Xray photon Photon flux is very high so very low noise Greatly improved contrast at lower photon energies Scatter and beamhardening can introduce bias uxyE is measured as an weighted average from 30120 keV so u Xy511keV must be calculated potentially introducing bias Xray and Annihilation Photon Transmission Imaging for Attenuation Correction Xray 30120 keV PET Transmission 511 keV Low noise Noisy Fast Slow Potential for bias when Quantitativer accurate scaled to 511 keV for 511 keV 100000 75000 Transform I 50000 25000 0 O 100 200 300 400 500 600 keV CTbased Attenuation Correction BiIinear scaling methods apply different scale factors for bone and nonbone materials Should be calibrated for every kVp andor contrast agent C 0 U E 015 C waterbone 9 m 4 6 3 010 alrwater mixture 2 mixture g 005 U 5 E 03900 air 39 soft tissue 39 c39lense boneI 1000 500 0 500 1000 1500 CT Hounsfield Units Typical PETCT scan protocol 1 Scout scan 5 20 sec 4 ll CT PET 3 Helical CT 2 O 60 sec 2 Selection of scan region 1 2 min 4 Wholebody PET 6 40 min i 4 Data flow CT images are also used for calibration attenuation correction of the PET data Xray Anatomical CT CT acquisition 7 Reconstruction Image Display gt Of PET Smooth to PET Translate CT to PET I Images V PET Emission Attenuation Correct Functional PET PET Acquisition PET Emission Data Reconstruction Image Note that images are not really fused but are displayed as fused or sidebyside with linked cursors Potential problems for CTbased attenuation correction Artifacts in the CT image propagate into the PET image since the CT is used for attenuation correction of the PET data Difference in CT and PET respiratory patterns Can lead to artifacts near the dome of the liver Use of contrast agent or implants Can cause incorrect values in PET image Truncation of CT image due to keeping arms in down in the field of view to match the PET scan Can cause artifacts in corresponding regions in PET image Bias in the CT image due to beamhardening and scatter from the arms in the field of view Effect of Contrast Agent on CT to PET Scaling The presence of Iodine confounds the scaling process as Iodine cannot be differentiated from bone by CT number alone In general does not seem to lead to artifacts Can use contrast scaling but then bone values are incorrect 3 0120 E Bonewater mix 0391 Curve that 5 BiIinear should be 4 0100 39 diquote used for g contrast agent 0090 T 100 o 100 200 300 CT Hounsfield Units Patient shifting Large change in attenuation going from spine to lung Impact of Wholebody Respiratory Gated PETCT in worst case Static wholebody Single respiratory phase 1 cc lesion on CT 1 of 7 so noisier The value ofthe lesion goes from 2 in the static image to 6 in one phase ofthe respiratorygated image sequence Respiratory Gated CT images 10 phases 8 mAs 5mm slices Wholebody Respiratory Gated PET 9 phases Note changes in 9s lesion intensity I g PETCT Applications and Challenges Primarily for Cancer Imaging works very well Diagnostic imaging and staging for cancer Expanding Areas with significant challenges Radiation treatment planning using PET and CT Cardiac imaging Assessment of therapeutic response PETCT and RTP using BTVs FDGbased boost volumes AnatomicalFunctional Mapping of the Heart rims Rest Stress Rb NH3 H2O Quanti cation MBF MFR Combining coronary imaging CT with perfusion PET Quantitative Assessment of Response to Therapy Basenne sp 4 mos Ietrozole Example Change in SUV measures of FDG and fluoride quot incorporation for bony metastases from breast cancer FDG before left and after hormonal therapy right 74 124 19 45 Bone images look similar but have very different values CT helps with precise F18 i realignment of ROls in serial studies 3 i 717 S 118 206 f 99 u z F1 339 SPECTCT Hybrid Systems Like PETCT SPECTCT acquires both scans with the patient in the same position Very new type of system not clear how this will be useful clinically but a lot of interest CT is also used for attenuation correcton of SPECT data Having the gamma camera and CT scanner on the same gantry allows straightfonvard fusion of the two data sets The CT provides accurate anatomical localisation of the functional information within the gamma camera scan It is claimed that the accuracy of radionuclide therapy planning can be increased by using the CT attenuation corrected SPECT data Applications in development include combined coronary CT angiography and myocardial perfusion imaging SPECTCT Hybrid Systems Very different approaches by the 39big 339 Kxi lquot 399 e x x v 4 n v 39 V V 7amp1 77 5 rth V Siemens Philips GE Highend CT not a real CT Entw Ievel CT Image Quality Alessiu r EIOSUB Image Quality Alessiu r EIOSUB Image Quality In art and advertising image quality can be entirely subjective For scienti c and medical purposes objective de nitions of image quality are needed Medical Image quality must be assessed on the basis of average performance of some task of interest by some observer or decision maker Image quality must be de ned in terms of as What information do you Wantfrom an image7 The Observer How Will you extract information from the image7 Object and image statis ics H H Barrett and K J Myers Famdananxaflmage 5mm Hubuken NJ Wiley znm Alessiu r EIOSUB Image Quality I Task detection of ventilation abnormalities Observer trained human rea er Image statistics patient dependent etc Alessiu r EIOSUB Alessio BIO508 Image Quality A Point Stimulus Stationary response Alessio BIO508 B Isotropic PSF Some Image Based Metrics for Image Quality Spatial Resolution VariantNonStationary C NonIsotropic PSF Nonstationary response Often assume stationary reponse but in reality most systems suffer from this Some Image Based Metrics for Image Quality Spatial Resolution Depends on pier density dots per inch ALSO Depends on resolution of imaging process Ideal Real system system Point Spread Functi T e s stems response to an in nitely small point 0 nexpress resolution in terms of full width at half m 39mum 4cm Alessio BIO508 Pier density 10 pixelscm FWHM 03cm Some Image Based Metrics for Image Quality Resolution Alessio BIO508 Some Image Based Metrics for Image Quality Resolution same dpi different resolution Alessiu r aroana Some Image Based Metrics for Image Quality Spatial Resolution Contrast Measure of differences in brightness in adjacent regions 0 ma e Noise Imaging process is a form of a random process information is random photons ux etc collection device introduces random noise electronic noise e Therefore noise is always resent Often quantify with signaltonoise ratioSNR strength of signal strength 0 noise Or contrasttonoise ratio CNR 7 Possible inverse problem discussion Artifacts Alessiu r aroana Artifacts Artificial image features From a variety of sources physics of imaging problem with scanner motion ofpatient image generationprocessing steps e c Alessiu r aroana Artifacts Artificial image features From a variety of sources physics of imaging problem with scanner motion ofpatient image generationprocessing steps e c Alessiu r aroana Assignment for Next Class Read chapters 1 and 4 Find 2 medical images of abnormal anatom or physiology pathology formed from the next lecture s modality xray radiographs Place these images i t Write 12 brief sentences describing each image Write 12 brief sentences describing differences between the Images Alessiu r ElOSUB Ultrasound Basic Idea Send waves into body which are reflected at the interfacs between tissue Return time of the waves tells us of the depth of the reflecting surface History First practical application 1912 unsuccessful search for Titanic WWII brought massive military research SONAR SOund Navigation And Ranging Midcentury used for nondstructive testing of materials First used as diagnostic tool in 1942 for localizing brain tumors 195039s 2D gray scale images 1965 or so realtime imaging Sonography relativey portable inexpensive and safe so is often the first choice of a medical imaging method where feasible Sound waves 0 Sound wave propagate by longitudinal motion compressionexpansion but not transverse motion sidetoside 0 Can be modeled as weights connected by springs Bab Spring Equlllbllum i l I l l 1 Acnush 0000 U rarelacrmn compression Ultrasonic Waves and properties 1 Propagation Amnlrmde Wavelength Mechanical waves are longitudinal oomprssion w v Ultrasound refers to frequencies greater than ZOkHz the limit of human hmring Medi l 39maging typically 100 Times higher frequency than audible by human typically 2 to 20 MHz Transmission and Reflection Trunamiucd Scrillcrud plunc wuw spherical mu ylimirr made l39rurmlucer 39 39 h39 lrunstluur li CL Transducer l ltLrLcl 1 Pulnl object Propagation of ultrasound waves in tissue Scattering Specular re ector is a smooth boundary between media cohvehuohai View ofre ecuons Acousnc scattering arises from objecB matare size ofwavelengm or smaller igiidiui SPECUUR ECHOES scATrERED ECHOES Specular e echoes orrgrheuhg from relatively large regularly shaped ObJSCB With smooth surfaces These echoes are relatively intense and angle dependent r e valves Re ectlon from large surface Scattered e echoes originating from relatively small Weakly re ective irregularly shaped objects are less angle dependent and less intense r e blood cells 7 Re ection from small surfaces Basic Idea momma 73 mm rmm skin swim Eulw mm organ hm Inca rnhn mm nah hack lace r t Volingc 4h Time llmlll ll l l 0 Along each line we transmit a pulse and plot the reflections that come back vs time Mnev 5km induce The Speed of Sound 0 The compressibility K and density p of a material combined with the laws of conservation of mass and momentum directly imply the existence of acoustic waves 0 Ultrasound waves travel at a speed ofsound c given by Variations in Speed 0 Speed of sound for Mil different materials A up 11 m Aluminum l39llll rum 1 ix mm mm c r asquot isnu pK new mu um mum on ma Iuno Hm lib 157V 1 mm was wan VlU HSU Illm lib 10b WC 4in m I HM Physics of Acoustic Waves Three dimensional in nature and depend on time Whatever the physical quantities that are used to des 39be t e sound waves they must depend upon three spatial variables x y 2 an time t Particle displaoement UX y z t associated with the comp ssion and expansion of the aooustic wave Particle velocity VX y z t Acoustic pressure pX y z t which is zero if there is no wave For longitudinal waves it is straightfonNard to relate the acoustic prssure to the underlying particle velocity p VZ where Z pC is called the characteristic imgdance This is a like VIR Note that V 75 C Variations in Speed and Impedance 0 Speed of sound M for different Munmi k materia s M 39139 l A mm 1 Aluminum 1m mm mu quotm 0 9m pic mm 1 y c w yin o Impedance relating pressure to particle 10W haw mo veIOCIty mm 910 H 7 Mil lion 17 7 v2 mm 1m 4m p 1039quot Z pf 7 mm K mm 1 w Wave Equation 0 The acoustic pressure p must satisfy the threedimensional wave equation 82 82 82 pxyzt 1 32pxyzt 2 2 2 2 8x 8y dz c at2 o For a plane wave traveling in the zdirection thus reduces to ow i 82m 8z2 62 Biz 0 An example solution is pzt COS kZ tc which has cyclic frequency in Hertz of f kc 275 C which also leads to the important relation f I Propagation of ultrasound waves in tissue Ultrasound imaging systems commonly operate at 35 MHz which corresponds to a wavelength of 044 mm when c 1540 ms Material 1 Refraction 0 When a wave passes from one medium to another the Material 2 frequency IS constant and since c changes then so must the wavelength Slnce M lt 7n C we have c2 ltc1 A f Propagation of ultrasound waves in tissue Bending of waves from one medium to another is 39refraction39 Follows Snell s Law sin 6i sin 6 sin 6 Cl Cl 62 since M lt 7L1 we have c2 ltc1 and 92 lt 91 Total Internal Reflection 0 Since M gt kl in this case we have c2 gt c1 and 92 gt 91 0 There can be a 39critical39 incident angle 91 2 6C where 92 90 deg ie there is no transmitted wave In that case there is 39total internal reflection of the wave Incident VJIVL Mcdi um Medium 3 Interface Re llcclcd l39C I 39l39rnnsmilted 39il39L Attenuation of ultrasound waves in tissue Attenuation is the term used to account for loss of wave amplitude or signalquot due to all mechanisms including absorption scattering and m The model of attenuation is phenomenological meaning it agrees well in practice but is not easily supported by theory de conversion We model amplitude decay as Az Anew where M is called the amplitude attenuation factor and has units cm l Sinoe 20 log1D AzA is the amplitude drop in decibels dB it is useful to define the attenuation ooefficient oc as 1 2010gme LlA E 874A The absorption ooefficient of a material is generally dependent on frequency f and a good model for this dependency is 1 afb The rough approximation that b 1 is often used Attenuation of ultrasound waves in tissue Assuming b1 vl iumrl itiii mr ll Ir i Aw Anew lmiu MW 1 w AWN n1 i L hm mu le l5 lllmnl ll 5 limit lll il mm ill i mil uvl l ummnm uni 1W 4 i w n 94 l I i imm r Vluu mnmr mm mm viiut mm mm t rm mmrui TimeGain Compensation Depth of signal is related to reflection time so as time progresses the signal will be increasingly attenuated Timedependent attenuation causes severe signal loss if not compensated All systems are equipped with circuitry that performs timegain compensation I39GC a timevarying amplification In practice most systems have additional frequency dependent slide potentiometers which allow the gain to be determined interactively by the operator This permits the user to quot39 I H3 f h manually adapt the system to special mm Mum lII circumstances requiring either more or less gain so that subtle features can be seen in the images lunggum winpunmuun iiiy hllg Zr Generation of Ultrasound A 39transducer39 converts energy from one form to another The Piezoelectric effect was described 1880 Pierre and Jacques Curie Lead zirconate titanate or PZT is the piezoelectric material used in nearly all medical ultrasound transducers It is a ceramic ferroelectric crystal exhibiting a strong piezoelectric effect and can be manufactured in nearly any shape The most common transducer shapes are the circle for single crystal transducer assemblies and the rectangle for multiple transducer assemblies such as those found in linear and phased arrays Plosiic Housmg Acousiic Insuloior Backing Block 4 4 9 v 4 quotLivzz39I Electrode Crystal Elemeni IIGroundH Elecirode Insulated Cover An electric eld realigns the IL poles in a piezoelectric crysta Beam Pattern Formation 0 Simple Field Pattern Model D VA Fresnel region G t t Some rlc approx39ma loquot Fraunhofer or farfield region Approximate field pattern for a focused transducer Collect the Echo V 11 delayed pulses array of PhasedArray concept for transmission and reception Focused l ll H IIlll i i generated wave transmission 39e 0e 39 39 39 p39 Z lectr39 senSItIve region reception c crystals Transducer Arrays A Linear Array B Phased Array C Transducer element 4 Width Thlckness A subset of transducer elements activated All transducer elements activated ili ear cugra39 Phased arrays IntracaVItary ultrasound probes 12 Array Transducers inear Array Convex curvilinear array Rectangular Field of View trapezoidal eld of view Linear arrays composed of 256 to 512 discrete transducer elements 15 to 20 adjacent elements simultaneously activated sequentially across surface to sweep FOV Beam steering Time Transmit Receive Timing Delay quot Phased array transducers 7 composed of 64 128 or 256 elements phase delay f 7 7 K compos39te beam area varied to sweep across FOV 39 r xi if Ki if X l l i l l l l l i l K x Sidelobes Focused arrays typically have larger 39sidelobes39 of signal power for transmission and sensitivity for reception 13 Amplitude AMode Tmmmillcd Eclm 1mm quot kln Unde thn mm f myun mm Lice l FL hnlIum mu hm R10 Time d ll ll lall Cum Skin m lam Along each line we transmit a pulse and plot the reflections that come back vs time Unfortunately it is very difficult to associate a precise physical meaning with the received signal amplitude vs time Ultrasonic Imaging Modes Stationary Transducer M mode Bmode Amode Incident pulse Returning echo w Resultant Mmode display Valve Leaflets 14 Ultrasonic Imaging Modes Echo Display Modes 0 Amode amplitude display of processed information from the receiver versus time Speed of sound equates to depth only used in ophthalmology applications now Bmode brightness Conversion of A mode information into brightnessmodulated dots Mmode motion uses Bmode information to display the echoes from a moving organ AMode Example Prnwrz ill I Vailnu tun an I M Ito inllnrml Transmission pulse in red re ected waves in blue Forming an Image 0 The amplitude values are converted to brightness along a line and displayed on a screen 0 The line direction is swept across an angular range either mechanically or electromagnetic beamforming Pmi rit beam sweep Forming Clinical Images Probe locations Two common clinical ultrasound examinations L an echocardiogram showing the four chambers of the heart R fetal ultrasound showing a normal fetus at the seoond trimester of gestation Complete System Acquisition and Recon Time For external imaging each line corresponds to 20 Velocity ofsound in soft tissue is 1540 ms Travel distanoe from and to transducer 40 cm Acquisition of line takes 260 us Typical image has 120 lines for total time of 31 ms Ima es reconstructed in real time So can have temporal resolution of 30 Hz 30 images a s d Mo ern scanners collect multiple scan lines simultaneoust usually frame rates of 7080 Hz Clinical Uses Cardiac Imaging Le Mm Le Emu12 wage m 2 mm heavt Empwe ofMMode be w zo Ewude 1w Clinical Uses Neonatal Bmode image of a fetus The dark region is the uterus which is filled with fluid 1 2 3 Doppler Imaging Blood moving towards transducer produces higher frequency echoes FIGURE 1645 Doppler ultrasound Sound waves reflected from a moving object are compressed higher frequency when move ing toward the transducer and expanded lower frequency when moving away from the transducer compared to the incident sound Wave frequency The difference between the incident and returning fre quencies is called the Doppler shift fre Blood moving away from transducer produces lower frequency echoes quency Continuous Wave CW Doppler Continuous sinusoidal wave transmitted with one crystal and reflected wave received with second crystal Pulsed Wave PW Doppler Pulsed waves transmitted at constant pulse repetition frequency and only one sample as function of time is collected Color Flow CF imaging 19 Doppler Imaging Doppler Imaging Color Flow CF imaging39 pp er equivalent of Bmode scansevera pulses instead of one are transmittedreceived along each line Calculates phase shift between two subsequent pulses Velocity information in color is superimposed on anatomical gray scale image Red 7 ow towards tiansducer Blue 7 ow away from tiansducer 20 3D Image Formation Reordering of the known slice locations provides surfaceshaded wire mesh MIP or other renditions of the anatomy Linear Wedge Freeform Rotationai Comparing 2D to 3D US 21 Dangers of Ultrasound very minimal in comparison to other methods I 0 development of heat tissues or water absorb the ultrasound energy which increases their temperature locally 0 formation of bubbles cavitation when dissolved gases come out of solution due to local heat caused by ultrasound 0 high intensity systems actually used for therapy Some Ultrasou nd Uses shortlist Obstetrics and Gynecology e measuring tne size ortne retus to determinetne due date 7 cnecking tne felus s growth rate by making many measurements oyertime impianted in tne mother s Fallopian tubes instead or in tne uterus baby seeing and ayoiding tne baby during amniocentesis sampling ottne amniotic uid Witn a needle tor genetictesting Years ago doctors use to pertorm tnis procedure o panying use ofullrasound tne risks ottnis procedure 7 seeing tumors ottne oyary and breast Cardiolo y eseeing tne inside ottne neart to identity abnonnai stmctures or tunctions emeasuring blood ow tnrougn tne neart and maiorbiood yesseis Urology emeasuring blood ow tnrougn tne kidney edetecting prostate cancer early 22 Breast Cancer Example 0 Not same dimension scale o In US we terns like hypoechoic or hypore ective for low intensity regions and hyperechoic or hypen39e ective for high intensity regions Dynamic Fetal Ultrasound Imaging Normal Brain scan example Fluid from intraventn39cular hemorrhage BIOEN 508 R Schmitz Projection X Ray Imaging V Toda 39 Bioengineering 508 Y Physical Aspects of Medical Imaging XRay Radiography httpcourses quot39 J nn edIIbioen508 Physics of Xrays For questions remarks discussions errors in the book 39 InteraCtions 0f Radiation With Matter Class Discussion Board link from class website Projection Radiography Monitored by instructors frequently Organizer Paul Kinahan PhD Adam Alessio PhD Ruth Schmitz phD Ruth E Schrniti PhD Lawrence MacDonald PhD I IDelJal m39lenit 0 Ra 390 09y UniverSIty of Washington Medical Center Imaging Research Laboratory rschmitzuwashingtonedu http deptswashingtonedunucmedIRL Department of Radiology University of Washington Medical Center BioEn 508 R E Schmitz October 4th 2006 0 BioEn 508 R E Schmitz October 4th 2006 1 Ci in V Some Elementary Particles V X rays Photons o Electromagnetic radiation of rather high energy Electrons large sparse outer cloud of atoms ordered into shells T fEM W I th E ype o ave eng nergy 9 Chemistry Table of the elements radiation cm ev Charge 391 gamma rays 1044108 104 108 Small mass x rays 109107 10 105 UV radiat on leO 7 4x10 5 3 25 visible light 4x10 5 7x10 5 2 3 infrared heat 10 4 01 0001 2 m m crowaves 01 1 10 4 10 3 Light particles or electromagnetic waves radowaves 1105 1079104 No Charge AC 60 Hz current 5x108 4X10 14 Massless Move at the speed of light BioEn 508 R E Schmitz October 4th 2006 2 BioEn 508 R E Schmitz October 4th 2006 3 October 4th 2006 1 BIOEN 508 R Schmitz Projection X Ray Imaging quot Discovery of Xrays V Xray energy relationships V Energy E and frequency v proportional E h v discovered in 1895 by Wilhelm Roentgen in cathode tubes 0 Energy E and wavelength 9 related inversely E h c it rst N0be39 Prize in PhYSiCS in 1901 frequency and wavelength related through speed of light 39 attenuated by VarGUS materials differenty v c it Revolutionized Medicine 102 to l mi 102 103 we iii5 we 107 108 1091010301113121013 first Medical Imaging Wavellengilhi mll I I I I I I I I I I I I I and beginning of scientific medicine Photon energy ieV l I l l i0 5 10 7 1076 10395 10394 10393 10 2 i0 1 i 10 102 103 104 105 106 to wedding ring floating around bone Rad ograph of Frau Roentgen s hand w th From Roentgen s announcement letter BioEn 508 R E Schmitz October 4th 2006 4 BioEn 508 R E Schmitz October 4th 2006 5 33 4 Xray Production 39 1 eVacuated tUbe heated cathode current releases electrons Electrons are accelerated to anode by voltage U Energy E qU EQ E Closer Look at Xray Spectrum V Characteristic radiation from knockouts in inner electron shell and subsequent filling of the lowenergy empty position In field of anode atoms electrons release their energy as Energy emitted is characteristic of the energy levels of the anode bremsstrahlung material I lead shield absorbing mostx rays EXray peak EAshells h39Vxray h39C A Wlth C A 39V Intensity vacuum Characteristic radiation 5 I 25kV 4 A l l 20 kV Continuous quot radiation w lvll a U30100kV 2 Yields a continuous Xray spectrum Wavelength A R 0 05 But what are the spikes on the highest energy spectrum BioEn 508 R E Schmitz October 4th 2006 6 BioEn 508 R E Schmitz October 4th 2006 7 Wavelength L October 4th 2006 2 BIOEN 508 R Schmitz K1 g Xray Source Parameters 0 Amount of xray photons Cathode current time mAs 0 Energy of the emitted photons controlled by voltage between anode and cathode kV but a spectrum of energies produced Often peak photon energy quoted kVp Intensin Characteristic radiation K Continuuus radiation 05 l 15 2 2 5 Wavelength i R BioEn 508 R E Schmitz October 4th 2006 8 to v Secondary Ionization o In Photoelectric Effect and Compton Scattering atoms are ionized o In pair production we have two charged particles produced 0 These energetic charged particles moving in matter ionize more atoms producing many free electrons o This secondary ionization is the basis for most detector systems 0 Ionization also leads to breaking molecular bonds basis of most radiation biological effects 0 Dose monitoring necessary when working with radiation BioEn 508 R E Schmitz October 4th 2006 10 October 4th 2006 A EQ E Interaction of Photons with Matter Projection X Ray Imaging The dominant photon interaction mechanisms for y and Xrays o Photoelectr c absorption Interaction with initially bound atomic electron Incident photon disappears Photon energy absorbed by electron momentum by atom probabil ty increases at low incident photon energy and high electron density in medium mass density X Z 0 Compton scatter Interaction wth free electron photon energy gtgt Binding E Scattered photon changes direction and loses energy 0 Rayleigh coherent scatter Interaction with with entire atom or molecule elastic scatter Photon changes direction though usually by small angle 0 Pair production y ray must have sufficient energy to create e e39 pair Ev gt 1022 MeV gt 2 me BioEn 508 R E Schmitz October 4th 2006 9 a Q Effect of interactions in matter Attenuation Narrow Beam Approximation Single interaction statistical event but overall effect governed by exponential law Lambert s Law L Scattered gamma rays lt gt Field of View ncident radiation of detector Vquotlt IO 5 i gt 4 Transmitted 39 I beam 7 gt Detector Attenuating object Colllmator Attenuation coefficient u BioEn 508 R E Schmitz October 4th 2006 11 BIOEN 508 R Schmitz Projection X Ray Imaging 3945 Linear Attenuation Coefficients 4Q Attenuation the fuller Story Function of photon energy and material u uE material I ID epL only valid for homogenous material not body Units lCm and one incoming photon energy For inhomogenous material u ux need udx integrated over that material in the direction of the beam M Lgt r Mxdx gtIOMIO eXpK f MOOdye m Exa m p l e Attenuation Xra beam We can break this problem down skull into 3 homogenous regions skull brainskull brain 3 l dx 2 luskull A39xskull lubrain Ax Attenuatlen enefflelent Hem Attenuation eneffieient Hem brain out 10 39eXp2 lusle 39Axskuzz lubrain 39Axbmm V Photon Energy keV depth BioEn 508 R E Schmitz October 4th 2006 13 4 quot Attenuation 4Q What ha 39 4 ppens In the real world 5 what makes X ray imaging possible 5 X rays image radiodensity amount of absorption in material Narrow beam approximation does not apply when Muse am We scatter can get measured or absorbed in dosimetry I calculations Figure 112 Radiodensity as a function of composition with thickness kept constant ncident radiation Attenuation is Energy dependent 100 kV u uupuu m filivpuu lcm layers l l l For cacu atng shedng dosmetry orfuxto an magng system scattered gamma rays must be taken nto account 74 84 98 21 68 96 attenuation Much more comp Gated For thin layers lower energies distinguish diff material better BioEn 508 R E Schmitz October 4th 2006 14 BioEn 508 R E Schmitz October 4th 2006 15 October 4th 2006 4 BIOEN 508 R Schmitz Projection X Ray Imaging as X39ray deteCtors Q Characteristics of film 1 ScreenFilm detector Graininess larger grains for faster film Film has silver halide crystals grains in emulsion When exposed to enough 0 Contrast slope of sensitometric curve D vs ogE phOtonS39 the grams tum metal with D optical dens ty darkness of film after exposure silver which darkens the film i i Very low quantum efficiency QE 2 o E exposure E Iin source duration FLUORESCENT COATING only useful in linear reg on 2 X45quot 504 i Fluorescent screens phosphor on Di senSitometric cu i FLUORESCENT COAT39NG both sides of the film absorb Xray y photons and covert them to visible 0 M inverse of amount of light w to to m 25 but add blur funct on of grain Size and scatter Film is indirect detector because need l l 39 0 Resolution l l the phosphor first i 1 Useful m or optical density funct on of grain size and scatter 3 Eggozure Fluorescence glowmg In VlSlbie light Immediately when h t by photons i39 i 9 Phosphorescence afterglow after the radiat on has stopped 09exposure IogE BioEn 508 R E Schmitz October 4th 2006 16 BioEn 508 R E Schmitz October 4th 2006 17 9 Xray detectors V Q Xray detectors for Computed Radiography 2 Image intensifier coupled to camera 3 Storage phosphors Xrays converted into light at phosphor screen then into electrons at Special phosphor screen no immediate absorption of Xray energy in phOtocathOder the phosphor Stored until stimulated with laser light Electrons accelerated and focused in tube with electromagnetic fields Converted back into light at output screen coupled to camera 0 Incident xrays boost electrons into conduct on band and trap them inpm there by impurities Latent image of stored energy stable for long Phosphor time periods lanil Screen Photocaihode Evacuated Tube Adva nta es dynamic reatime imaging 0 Extract stored informat on by pierWIse scanning wth a laserbeam Xrays Output wh ch lets electrons fall back into valence band and release VlSlbie light Screen gt 39 Disadvantages gt 0 Light captured by optic array and passed to photomultiplier gt g quot9 quot 39 worse Spat39a39 reso39Ut39on converted to electr cal signal El l l 39 39 ecron crises camera I I I I I gif glgiw more noise conversions Signal is digitized and recorded 0 eometric in cushion g p o Phosphor ready for next use after strong light HighVol1ago Powchupply distortion espeCIally at edges source irradiation BioEn 508 R E Schmitz October 4th 2006 18 BioEn 508 R E Schmitz October 4th 2006 19 October 4th 2006 5 BIOEN 508 R Schmitz Projection X Ray Imaging V V V 4 r 4Q Qv Storage Phosphors Advantages V Xray imaging chain Complete radiographic imaging chain Much wider useful exposure range than film screen no grains must 0 Xray tube be blackened Aluminum filter absorbs low energy photons beam hardening Collimator limits irradiated area 0 Patient attenuates and scatters x rays Linear exposure range ie no contrast reduction in low and high density areas of the image gt Very tolerant to over and underexposure o Collimating scatter grid mtenergy H 1 b CO ma mg absorbs large angle HQE39W scattergnd Image can be post processed image enhancement quantification scatter photons I i 0 Digital image easy to store transport distribute Detector Immediate availability through dig tal image database m i l l x eray source collimator detector BioEn 508 R E Schmitz October 4th 2006 20 BioEn 508 R E Schmitz 39 uctober 4th 2006 21 q 4 9 Finally39 Images V Radiographic Image Quality V 0 Resolution how many line pairs per mm can be distinguished lpm 39 Quality Of anode tip 900d angle for 900039 beam fOCUS Xrays produce transmission images Shadowgrams thicker patients more scatter less resolut on use collimator grid Structures visible due to different attenuation in diff materials light scattering properties of phosphor film resolution mainly determined by grain size See 4 primary densities gas fat watersoft tissue bone sampling size for image intensifiers computed radiography laser spot size for read out in computed radiography Ideally high resolution and high speed but interdependent Muscle Blood Liver Butter W Photon counting is poisson process SNR square root number of counts faster detector or higher speed less photons needed lower SNR n m u m dose req u reme nt Fig e 112 Radiodensity as a functionjof conosition with thickness kept constjani g a S Contrast base Water fa39t determined by film contrast or can be manipulated in digital methods higher contrast lower useful exposure range Aside X ray projection images should be called Radiographs Artifacts generally artifact free except for pin cushion effect and organ overlay BioEn 508 R E Schmitz October 4th 2006 22 BioEn 508 R E Schmitz October 4th 2006 23 October 4th 2006 6 BIOEN 508 R Schmitz Projection XRay Imaging w w 45 The need for gt1 Pr0jectlon Q What is this 2 dimensional projection of 3D object gt always need at least 2 projections to identify correctly A Radiographs of geometric objects In which ways can different objects make the same radiograph What is the structure on the outside vs middle end BioEn 508 R E Schmitz October 4th 2006 24 BioEn 508 R E Schmitz October 4th 2006 25 q V Clinical Use Advantages and Disadvantages Then what about this Advantages Cheap 0 Fast 0 Available Disadvantages Overlap of structures a Poor contrast resolution A Frontal B Frontal C Lateral Trick cream containing a metallic salt was applied to finger Real radiograph of finger looks like image on left BioEn 508 R E Schmitz October 4th 2006 26 BioEn 508 R E Schmitz October 4th 2006 27 October 4th 2006 Lecture 5 Tomographic nuclear systems SPECT Field trip this saturday at 11 AM at UWMC meet in main hospital lobby at 11 AM if you miss the 39boat39 page me at 5404950 should take 1 to 15 hours depending No class next week Exam will will be emailed to class email list Wednesday afternoon Submit to class website closes at 10 PM Nov 1st Next homework Read chp 7 Seutens Find 2 medical images of abnormal anatomy or physiology pathology formed using ultrasound AKA 39sonography39 Place these images in a document Write 12 brief sentences describing each image Write 12 brief sentences describing differences between the images Write 12 sentence what the image values represent physically Project Groups 1 Rodriguez Dones 2 Romig Martinez Sung Harmelin 3 Pham Wong Legesse 4 Braddock Mauro 5 Jeddi Miller Zhang 6 Morrow Dahl Clayton 7 Rundgren Lam Deadlines Nov 1 Outline due 30 of mark for project Nov 29 Final report due 50 of mark for project Dec 6 Class presentation 20 of mark for project Xray Projection Imaging overlay of all information non quantitative source Xil y detector 7X pxzgt f fxyz dy Xray Tomographic Imaging 39Tomo 39graphy Greek 39slice 39picture orbiting source detector data for all angles true crosssectional image Nuclear Medicine Tomographic Imaging single photon emission computed tomography SPECT r C Underlying Principles of SPECT H Rotating NalT Detector Module with PbW Collimator in Front Direction of Rotation Signals to electronics Scintillation site for y ray parallel to collimator holes 39 39f kx ra w Radioisotope decay by 39Yemission Nuclear Medicine Imaging RVL shinquot W at Projection Images WIMI IMI 39 v A z t 1 a 9 3 s r p a i I 1 7 lf 7 n UKT I IWIIQI gm Ir Tomographlc Images m v 3933quot 1 7 Wit 3 h A an nu I Iquot IiiL rum iquot r add Ll 2h 739 w Basic SPECT A Xial level of sinogram 2D planar projection Sinogram Angle of above projection Early Clinical SPECT I 5quotle mum I GE 400T Rotating Anger Camera ca 1981 Modern Clinical Systems EM39II 39 V ph39r c d39 60 Siemensecam G lemum G Hps ar lo VariableAngle Conventional Anger Camera PMTs coupled to large continuous NaT crystal Spatial resolution 3 4 mm FVVHM Energy resolution 8 10 FWHM Mature technology DoB 1957 Largearea gt40cm gtlt 40cm typical Simple and costeffective SPRINT II camera miodulei Photon Absorption 0 Ideal Photon Detection o Absorbed Photon 5mm NOTE at 30cm cvlinder center SPECT p0153cm Mo 010 Attenuation Effects 5 cm dia 10 cm 20 cm 0 O 30 cm Attenuation 015cm 140 keV photons in water Reconstructions of disks should be uniform Attenuation causes a distortion that increases with object size Possible Attenuation Artifacts Activity in myocardium should be uniform but isn t Could be interpreted as abnormal Breast Attenuation Diaphragmatic Attenuation Should be uniform intensit Tc99m Sestimibi Myocardial Perfusion Images 1 TCTECT Acquisition Geometry Transmi sion line source Detector 1 Transmission and emission 65cm focal distance fanbeam collimator Detectors 2amp3 Emission only Parallel hole collimator Transmission Source Am241 60 ke V photons This is one configuration Many others are in use Photon Scatter Ideal Photon Detection Scattered Photon NOTES Additive Loss in resolution and contrast Typica scatter fractions T1 0 Tc03 04 PET 015040 Camera Energy Spectrum 140 keV 5000 Energy WIndow lt 4000 39 3000 39 15 20 Energy window Counts 2000 39 1000 39 200 2300 Energy Channel Typical Anger camera has from 8 10 FWHM energy resolution at 140 keV Collimation Systems The collimation system is the heart of the SPECT instrument it s the frontend and has the biggest impact on SNR Its function is to form an image by determining the direction along which gammarays propagate Ideally a lens similar to that used for visible photon wavelengths would be used for high efficiency not feasible at gammaray wavelengths Absorbing collimation typically used Parallel Hole Collimation Detector Fl Collimator Photon reaches a h L detector hoton absorbed by collimator channel Source ltUom Um mm Iom 003581 in a 1 n iguana sa ampk aw 35an auras sung Parallel Hole Collimation 2 16L2 L Efficiency Intrinsic Resolution FWHM Response Single collimator channel X distance V Response is trapezoidal becoming triangular at larger distances Essentially solidangle ef ciency of single channel Resolution vs Depth h1mm h2mm L 30mm L 20mm L 10mm Point source depths 10 15 20 25 30 35 cm Sad Facts About Absorbing Collimation In the best case efficiency is only a few photons detected for 10000 emitted Performance of absorbing collimation decreases rapidly with increasing energy Photoelectric absorption coefficient changes as Z4 E3 Typical collimator for 140 keV Tc99m 9mm FHWM 10cm 23e4 Typical for 364 keV l131 12mm FWHM 10cm 1e4 Radionuclides with even a small fraction of higher energy radiation will result in severely compromised imaging performance Alternatives for Increasing Efficiency P llllllllllllllll Surround patient with more cameras Use fanbeam or conebeam collimation to trade FOV size for efficiency at given resolution IlllIllllllllIll Xl lllllllllll llllllllllllllll ll llllltllll RARD 6 i llllllllllll lllllillllll Modern Image Generation From continuous real world to a meaningful image on computer 1 Sampling Continuous Information Information and sampling technique varies widely for each modality Topic for later lectures Computer can only hold discrete chunks of data Pixel a single picture element Voxel a single volume element 2 Quantizing Samples Each discrete chunk must be represented by certain number its 3 Visualization Techniques of quantized sampled image volumes Image Visualization Alesslo r BIO508 Intro to Color additive subtractive Re ected colors add up Re ected colors subtract Primary colors re ect only one light Primary colors absorb only one light Re ect greenbluered light white light As add together absorb more light As add together re ect more light Alesslor BIO508 Intro to Color Saturation Hue dominant wavelength of light color Saturation intensity of speci c hue high saturation is a vivid color Brightness luminance ofvisual target amount oflight Alesslo r BIO508 Intro to Color Chromatic light needs three descriptors Humans basically have 3 types of cone cells in their eyes receptive to 3 primary colors Achromatic light needs only one Intensity Achromatic light has a saturation equal to 0 Medical images are generally just intensity based single descriptor Alesslor BIO508 Intro to Color Colormaps map an intensity value to a set display value or color ex Pixe1 have intensity value 296 Map1pixe1 R100 G100 Bo gt pixel looks yellow Linear colormap for single color described bywlrlduw and level WI ncl ow lt display lrlterlslty value input lrlterlslty value level Alessle e ElOSUB Colormaps CT example Reader selects preset viewi ng window to level and window the colormap for particular intensity values crlelee Hu 7 lzzanu hone window clwum 2500 zanHu olsn w 29 HLJ ozan HU luuu WU Alessle e ElOSUB Colormaps Colormaps offer exibility over how image data is displayed lmageJ Example Alessle e ElOSUB Colormaps Alessle e ElOSUB 3D Visualization general information Anisotropic vs Isotropic Isotropic allows for true 3D visualization Aiessiu r ElOSUB 3D Visualization general information Isotropic vs Anisotropic Multiplanar reformat Axial or Transaxial XY Plane 0m Damn Coronal Section Sagittai Sechon Hunzuriial Section Y Aiessiu r ElOSUB general information 3D infonnation can be Viewed many different Ways Following ifnages are from Same 3D CT data set Aiessiu r ElOSUB 3D Visualization general information Common Options 1 Multiplanar Reformat arguably the most important in diagnostic imaging 2 MIP maximum intensity projection 3 Surface Shading 4 Volume Rendering extensions to surface shading Aiessiu r ElOSUB Volume Visualization i Aiessiu r EiOSUB Assignment for Next Class Read chapters 1 and 4 Find 2 medical images of abnormal anatom or physiology pathology formed from the next lecture s modality xray radiographs Place these images i t Write 12 brief sentences describing each image Write 12 brief sentences describing differences between the Images Alessiu r ElOSUB Lecture 3 My Cnmpuled Tnmngrzphy Easeg Em chamev g1 SUE EHS when 1h uvdev Humewuvkmmextweek Read chamevE g1 sge1ehs up u ahg1he1gg1hg seem a A F1hg2 meg1ea1 wages gvahhgwa1ahammymphysmmgy pa1hg1ggwgweg gs1hg Nge1eayMeg1e1he p1ahaygamma cameva Wage hm PET m SPECT P1aee 1hese wages 1h a ggegmem WME 172 hhe1 seh1ehees gesehhmg eaeh wage Wme 172 hhe1 seh1ehees gesehhmg gmeyehees hemeeh 1he wages Wme 172 seh1ehee wha1 1he wage va1ges vepvesem phys1ea11y Fuvm gmgps mm by nextweek 1w111ass1gh gmgps 11 yuu wam wesh eeme pvub12m1nvu1v1ng meg1ea1wag1hg ahg pvesem yum eghe1gs1ghs 1h a WWWquot vepun ahg15 whme gya1 pvesematmn Em Dee 51h The mvESHgaHun shgg1g have 1he 1g11gw1hg 01122 BumpunEMS 111shgg1g1ayge1a speeme uvgan msease andDYDWEVEDnden Z 11shggg s ee y g 1he wag1hg mggahhes g1segsseg 1h dass ahg 311shgg1g gehhe 1he ghTeehve heh1hg1he use 11 meg1ea1 wag1hg a Class Project Examine g1 each g11he1h1ee cgmpgheh1s ave g g as Madam meme amhmmm my De emanmivnoss Lghgemey ayeesaemey age 1 VWVvessmn emhewmewmsese hgeemeaehe wegsmh Shake gmsggha Wave vumed may ngnh was MR1 Se mum VuuvaPun shgg1g addvessmesE gues11ghs 1 a11s1he mgmem he1ng1hves11ga1eg ahgwm 1s11wpgnahm 2 WW 1s 1he chgseh megahw g1 mgga1111es1he bes1 chgmem addvessme Wuh12m7 Th1sshgg1gheahavggmemhaseggh1he1eehh1ca1hehems e g VESD UHDH SNR Speed e12 mvggvehgme ascumpaved 1g g1he1 gmmhs 3 ng1s1he mgmem EUHEHW addvessed gs1ng1h1s mudamw Wha1 wage mgcess1ng1s1egg11egv cue apmgpha1e1e1e1ehees11gm1he111e1a1g1e a vvha1cgg1ghegghe1ghenevaggvess1h1spvgh1em1h1he1mg1ev Wha1 abum 1he pe11phe1a1 g1 suPPUn eggmmemvT a1 1s 1Wuu W212 askedm wmme1heme1hggg1ug11wha1mehgeswgmgvggpmsgemsw Class Project Examine Pvmects 1 Eva1nmmuvMR 1IPYUEvessmanhe gga11s1g measuve ehahges1h1gmms1ze Pmmessmn we 1we Th1s can be gseg 1g assess the VESPUHSE enhe when m heawem MR1 pmwdes gggg sn 11ssge dEYmmun heeessarv 1g 1geh11rv 1gmgyhgghgahes 2 Lwev g1sease1 cT ahg g1uasgghg1Reg1sha11gh a The gga11s1g a11gh1he cT ahg g1hasgghg wages yeg1sha11gh su 1hev can be g1sp1a1eg h a Eummed wage Th1s can be gse1g11h m1n1maHv1nvas1v2 suvgerv Wneveme g1hasgghg1s gseg 1h 1231 1we 1g gg1ge1he sgvgegh ahg1he cT pmwdes h1gh dEWHHun wages gnhe anammv nemlmes New ou111hegge 3ng1mavk1gvp1g1ec1 Ngvza r1ha1 vepgngge 5ng1ma11lt1uvp1g1ec1 Dec 6 c1ass pvesemangh 2m gvmawmmmeu Projection Imaging Versus Tomography The heeg gymme1hah ghe pyegeehgh L a Depehgmg gh1he wew ang e the same g1 eyeh1 gmeen A cuquavy 1s1ha1 WEI dMevem gmens eah 1gg1lt1he same1h une uvaVEW 39 v12ws2xamp1es7 sg hgw mahy wews gg We heeg 1g gh1gge1y 1geh111y ah gmeew And wha1 can We gg huh a111he 1h1gwa11gh7 Basic Principles Conventional Radiograph Forms a 39collapsed39 projection view without any depth information like a gamma camera image Example of quotProjection Imagingquot Here two views PA and RL lateral are shown 9 quot chest gtltray image 3 mei scintillator and ilm pack Projection Imaging overlay of all information non quantitative detector pxz f fxyz Tomographic Imaging Tomo graphy Greek slice picture orbiting source detecto data for all angles r true crosssectional image Formation of a CT image Measurement Data in m HIkam 1 nmnm walk Mm A ray a single transmission measurement through the patient made by a single detector at a given moment in time A projection or view a series of rays that pass through the patient at the same orientation Two Types of Tom ograpny Transmission and Emission b detector A i T quoti I Transmission nuciear Magnetic resonance imaging MRi or MR and uitrasound US are somewnere in between in tnat tney use emission stimuiated by an externai source SPECI39 or PEI39 Emission Major Medical im aging Modalities Modality Resolution TX or EM Mode x r r a Pianar Nuciear Medicine 10 e 20 mm EM Xcray CT 1mm Tx Tomo ra nic Uitrasound 3 mm TXEM sound Tomograpnic 1 mm TXEM RF MRI Tomograpnic PETSPECT 5 c 10 mm EM Tomographic xray imaging anay imaging nas been used tor medicai imaging tor oyer a century and pianenfiim eray scanning is tne Wurknurse o radioiogy departments VMtn radiograonsi noweyeri it is d to see iuvvncuntrast obiects imcuit ComputedTomograonyCTwa deyeiooed in tne eariy iams at EMi by Godtrey Hounstieid to generate crosssectionai mageso aoa en a me tomog onici t ti tAttn tti EMin diarge rotit s tne Eieaties were recording undertneir Pariurpnune bei Terms sucn as computerized transyerse axiai tomograony x mmputedrassistedtumugrapny or computerized axiai tomograony CAT naye been used Tne term computed tomograony CT was standardized by tne Radiuiugicai Society ot Mortn America Comparing Projection to Tom ograpiiic Images Hnunsfieid s insignt wastnat by imaging aiitne way around a patientwe snouid nave enougn infurmatinn to farm a urnssnsectinnai image Radingraphs ypmaiiy nave nigner iesoiution butmucn iower contrast and no deptn infnrmaIinri By stacking a series or 20 Xcray cT imageswe can get ayoiumetric image ordata seti wnicn is tnen dispiayed by making at pnncipai sectionstniougn tne image yoiume r nest iadogvapn Emmi section ct a so Timage yoiurne Historical Development of CT oondlbum 197 Ma hu hum I1972 v nu m In l 4 Wmuu mamIi my 39 Why mun nu hum mm in mm mm 317 quot s 39 5 km ink Mr 1 rK a L i mu Historical Development of CT Socalled 39Fi h39 Generation CT StationaryStationary 1990 s AKA Electron beam scanner Primarily for cardiologists 50 msec scan times Uses tungsten target and highenergy electron beam Now largely obsolete electron bending data acquisition system bea coils I y r detector rings electron fucussing gun coils H jjl xrays i ii vacuum pumps patient table target rings CT Developments 1972 Invention of CT 1978 Head scan takes 30 min 1986 Slipring technology 1 second scan 1989 Helical CT 1998 Multidetector CT 12 second scan 2000 57 million CT examinations done in 7645 facilities 2001 Commercial PETCT 2002 4 8 and 16 slice CT 2003 32 slice CT 2003 Head scan takes 3 seconds 2004 64 slice CT 03 second scan 2006 Dual tube CT scanners How it works CT Scan Concept Third Generation CT Xray Tube Rotating fan beam 03 to 2 seconds to acquire an image 3 rpm to 200 rpm Workhorse for CT scanners Rotating gantry Xray Beam Patient Table Detector array arc CT Tube and Detectors Xray tube Xray beam 39conditioning39 Detectors CT Scanner in Operation 64slice CT weight 1 ton speed 033 sec 180 rpm Xray CT Scanning Stationary anode tube low power version suitable for radiographs focal spots Rotating anode tube dissipates heat for higher beam currents needed for CT r r lliml mm m1er l ivlt m mummy mm m m ummlmurimmgnrlurwmlulmntrm mlw Modern XRay Tube Electron Collector reduce offfocal radiation Lower patient dose Rotation speed s 05 High PeakPower Target amp Bearings High peakmA for fast rotation typical mAs 200 mA needed Large Patient Large Patient Detectors Used in CT Scanners O dergeneratmn used XEHDH as dEtE DrS E fast mmzatmn detemurs ere reeenuy suhd same date mm syste ms 6 man Xray c1 Physics The delecmvs ave basmaHy svm av e methuds used m nuc eav magmg sys ems semuuauen eweweu by ham eeueeuen The scmHHamv e 3 cs cunvens the mghrenevgy phumn m a ham pu se whmh s demoted by phmu umues Eg m scmmlaxian mum collimation cmhmamysaye useum pvmec he pan mbyv mmgmexrvaybeamm me anatumy uhmeve Pyepauemeemmauen sm uenced by me neax sputswze because Mpenumbva Theaygeymemeaxspe s zewe gveatenhe penumbva and wave Detectuv m pustpauem eemmauen my beam and vemuves sceneved Va atmn cmhmauen a su he ps uevme shee thmkness u 5 mm m 11 mm dependmg un scannev 1 z mu tpmmmmm x WWW Mum R Dmecinr onlumuar mmmmummm PrePatiem collimation Camus pauem vamauun expusuve xeray mbe co h mator asserer xeray sht PostPatient Collim ation Adjusts Wage shmnmkness coH mater detectors Major Component of a CT Scanner System Typical con guration GanW Paws GanW Buth swung 5 Tube detatcr Amman mam and yawn 093339 RWDE CUPACE svunmnn EWSD E d su sv pnntavs Pahent bed m ammmmm I K a a J Patient Couch Cuuches avemeauv capame msuppumng wman puunusmn an tamE pusmunmu accuvaEvvemudummhw m n 25 mm Hunzuma muvemem vange s Wmcauv 17mm znn cm axnnum hunzuma muvemem speeds m mu m an mms Gantry Slip Rings AHqu m cummuuus vmanun Helical CT Scanning Patient is transported continuously through gantry while data are acquired continuously during several 360deg rotations The ability to rapidly cover a large volume in a singlebreath hold eliminates respiratory misregistration and reduces the volume of intravenous contrast required The helical data set is interpolated into a series of planar image data sets 0 Production of additional overlapping images with no additional dose to the patient 0 Images can be reconstructed at any level and in any increment but must have a thickness equal to the collimation used Helical Scanning and Image Interpolation Pitch I Pitch K distance that table travels per 360 rotati0n slice thickness Collimator Pitch o A pitch of 10 implies axial scanning best image quality in helical CT scanning o A pitch of less than 10 involves overscanning some slight improvement in image quality but higher radiation dose to the patient 0 A pitch greater than 10 is not sampling enough to avoid artifacts Faster scan time often more than compensates for undersampling artifacts Also reduction in patient radiation dose Increasing Pitch to Reduce Scan Duration Faster acquisition mode same region of body scanned in fewer rotations thus less motion effects Pitch 1 Pitch 2 Single versus Multislice Detectors Can image multiple planes at once 1 detector row 4 detector rows Helical MultiDetector CT MDCT Fastest possible acquisition mode same region of body scanned in t fewer rotati Single row ons even less motion effec s scanners have to either scan longer or have bigger gaps in coverage or accept less patient coverage The real advantage is reduction in scan time yquotquot ill l 1 detector row pitch 1 and 2 001 4 detector rows pitch 1 34 sizchannai by l mw 14592lndlvldual elem CT Detectors Single Slice vs Multislice CT Each channel ls t mm wma Each alemeni ls 2n mm tall Singleslice CT detectors are based on a singlerow detector array Each channel isi mm was Each alemenll51 25 mm tall Multislice CT uses multiplerow detector array with each row consisting of individual a r r elements 5 77 7 ants an a Ssrdegres arc Four channel system Although there are 16 detectors in direction in example only four groups can at a time This is calle CT scanner Detector collimators ie postpatient to match be Multirow Detector Arrays axial this be read out multiple detector array d a 394slice39 2 detectors binned lS adjusted collimated xray am Width beam width using Llr 50 mm detectors L 4 detectors nned collimated xray beam Width usan four 125 mm detectors A E C D detector configurations Xray Computed Tomography For an ideal narrow beam ofmonoenergetic photons the fractional reduction ofthe beam intensitylis given by 7d I udt This can be integrated to give Ir In exp rjudr n The solution is LambertBeer law 1 In epoxt Xray Computed Tomography For spatially varying attenuation coef cients Ax which is what we really want right we can convert it to a simple integral with reference to the initial unattenuated beam intensity In em mm 2 we I a Xray Computed Tomography The function pre is formally called the pro39ection ofuxy along the direction 9 This can is more easily visualized as a row or column sum of a matrix sum in Is direction pr9 We need to have pre for all re to determine the original object uxy Estimating ux from pre knowing the relation pr9 J pxyds is a classic inverse problem a Xray Computed Tomography quot 39 calculate r 39 39 angle i J rotate the image with any program and sum in columns or rows r sum in this direction pr9 We can group all the data for pre in a 2D array to make a sort of image which is called a sinogram Onedimensional projections I Projection pgtltr gtltr pxR idyR my Yr y gt l XR l l cos sin l l x l Xr LxRJ Lsin cos JyJ gtlt ISinogram sgtltr single projection 1 lt sine wave traced out by a point at x0y0 gtltr Object fgtlty XavYO Xray Computed Tomography Given the sinogram pre how do we get backto uxy The details are beyond the scope of our course but are described in section 532 The core concept is backpro39ection which is placing values back into the image matrix place back in this direction 1909 Then repeat for all anglese 11 Formation ora CT image Filtered back projection IL mm mm Wm W 3 1 v vmrmm amp a l H e in 1 2D Central Section Theorem pmhamwhe swmme ww m uheevs1ahe the necessary cumumhuhmerahezo Wage vecuhsmnw cummessed mm 3 shdes Am 012 mutemunrshce memem P011 W Projec uon pm w j mm XR1DFT y quot1gt magmg forward mode Objea ay 2 Eackprojection The vesuh e1 backpvujectmg a shgxe pmjactmn 5 wuwa em m p acmg the the Fuunev hahsmhhee va ues mm the vecunshuctmn mam whmh vepvesems magmmde and phase enhe the spahax heeuehmes enhe Wage a s hm Jew pmw I IdxdyexpGZIMm Hm y 2 Eackprojection and Filtering Thezo Feuhemahsmhh sbmh up hem the 1 D FTs enhe pvujectmns h the hmmng ease asAVMandrgtEHhe samphng dEnSNy h heeuehey space 5 pvupuv unatu1 V Thus 3011410 F111 11h11 su F111 11v113111 11v We normahze the unevemy samp ed Fou ertransform of the omen huh the ramp ther M Thxs samphhg s the same for PET SPEC CT not MR h We do not use ms horma zauom We get heavhy b urred Wages Reconstruction Demo 1 Filtered Backprojection Image Noise vs Resolution Edge enhancement algorithm Standard algorithm Image detail Smoothing algorithm Image noise Bone filter fine detail increased noise Soft tissue filters smoothing decreases image noise and spatial resolution The choice of the best filterto use with the reconstruction algorithm depends on the clinical task Image Noise vs Resolution Smooth Reconstruction Demo 2 Noise versus Resolution 13 What Is Being Measured During acquisition each detector element is related to the average linear attenuation coefficient ofthe tissue contained in each voxel along the line of response LOR from tube to detector xray tube D detector HighEnergy Photons and Interactions with Matter The names for highenergy photons are determined by the source not the energy 0 Xrays come from bremsstrahlung electrons are accelerated from cathode to an anode by a voltage Vp and bend around orbital electrons in the anode target The bending is an acceleration that releases energy mostly heat y rays come from nuclear decay processes that release energy 0 Annihilation photons come from the mutual annihilation of electrons and positrons Their mass is converted to energy according to E mc2 often called yrays by mistake Interactions in the energy range of 301000 keV are 0 Photoelectric absorption Photon energy absorbed by electron Dominates at lower energies or for materials with high 2 values 0 Compton scatter photon scatters off a free electron and changes direction and loses energy Dominates at higher energies or for materials with low 2 values biological materials Charged particles interact in less than a mm photons take many cm Xray CT Image Values 0 With CT attempt to determine uxy but due to the bremsstrahlung spectrum we have a complicated weighting of uxy at different energies which will change with scanner and patient thickness due to differential absorption 100000 75000 I 50000 25000 Energy dependent linear attenuation ooef cients 11X cle bone and mus Input xray bremsstrahlung spectrum intensity vs photon energy r a commercial xray CT tube set to 120 kVp Componens of linear attenuation coefficients Ermatotal 39Ennephataelectric y for CT Numbers or Hounsfield Units We can39t solve the real inverse problem since we have a mix of densities of materials each with different Compton and photoelectric attenuation factors at different energies and a weighted energy spectrum The best we can do is to use an ad hoc image scaling The CT number for each pixel xy of the image is set to x y uwater CTx y 1000 limiter px y is the attenuation coefficient for the voxel pme is the attenuation coefficient of water and CT image CT xy is the CT number or Hounsfield unit that comprises the final clinical 14 CT Numbers or Hounsfield Units Typical values compact CT value bone HU 1000 80 50 70 her 600 60 blood sponglous 00 bone 50 pancreas 200 water 39 kidquot 0 at 40 39 200 V 30 400 Inn 5 20 600 we 1 1000 an i o M2000 Digital Image Display WindowLevel I o The window width W determines the contrast ofthe image with 39 narrower wmdows wmasr was warm mu warns L any resulting in greater W a1 contrast White L The level L is the CT number at the center of 939 the window Display Intensity hlark 4mm 1 z taonu CT Number P1 L 12W P2 L 12 w Digital Image Display WindowLevel The dynamic range of Xray images is broader than can be displayed or perceived thus we use the windowlevel settings to control what we see CT value HU 22su W 3003 bone window CW I DUO 2500 7250 quotU 2000 15 nu Iooo f quot eduslmai w n w CW 750 400 250 Hu 25 my 9 window moo ciwraoqjip mm r I I I I V I gt V 59 Contrast Agents o Iodine and bariumbased contrast agents very high Z can be used to enhance small blood vessels and to show breakdowns in the vasculature Enhances contrast mechanisms in CT Typically iodine is injected for blood flow and barium swallowed for GI air is now used in lower colon in Iodine 0 WP gImil luau CT scan without 1 CT scan with contrast 15 in Applications wide iange ei anneimamies ei diseases in anv pan eiine eedv caicium Scuvm Radiaiien iieaimeni pianning Tiauma inieciiennnnamaiien FuHWrup ei cemeniienai cnesi way iindines aenenaciuie spinai ceiumn damage xray c1 Scanning oiinicai Uses The aduaniage eici cempaied ie pieieciien iadiegiapny is impmved ceniiasi and deiaii iei cempiex siiuciuies se ii is used in cases wneie inis deiaiied inieimaiien is impenani eneugn ie euiweign ine cesi ii can aise be used in an pans eiine eedy uniike US ei MW and is e c ideiec ing e ne iiaeiuies ie ie ine densiiy iange iei imaging Wiin ceniiasi agenis can be used iei iinding iegiens ei anneimai eieed iiew eg canceiin ine iwei CT siices inieugn ine eiain snew a sueduiai nemeiinage as a nypeidense iegien eieed meie dense inan eiain iissue aieng ine innei SKuH waii xray c1 Scanning oiinicai Uses CT eiine cnesi a Mediasiinai and e iung windewieuei seiiings and c ceienaiiesiiced image The images snew a cengeniiai maiieimaiien eiine iung iecaied in ine ieii ieweiieee Nuiicemeiwu cempeneniseiine iesien a dense muiiiieeuiai epaciiy anew suiieunded by an aiea eidecieased iung anenuaiien anew needs xray c1 Scanning oiinicai Uses CT siice inieugn ine ceien snews a ceienescepy a eaies a d Aiieieanumisusedasaceniiasia i 39Technique39 Technique refers to the factors that control image quality and patient radiation os kVp kV potential energy distribution of Xray photons recall lower energy photons are absorbed more readily mA number of Xray photons per second controlled with tube current s gantry rotation time in seconds mAs total number of photons photons per second X seconds pitch slice collimation filtration filters placed between tube and patient to adjust energy andor attenuation not discussed here Radiation Dose in CT CT acquisition requires a high SNR to achieve high contrast resolution and therefore the dose to the slice volume is higher because the techniques used are higher a PA Chest xray 120 kVp 2 mAs r Chest CT 120 kVp 200 mAs Radiation dose is linearly related to mAs At same kVp and mAs number of detected photons increases linearly with slice thickness SNR improves Larger slice thickness at same technique yields better contrast resolution higher SNR but spatial resolution in the slice thickness dimension is reduced Smaller slice thickness improves spatial resolution in slice thickness dimension and reduces partial volume averaging Noise will increase with thinner slices unless mAs is also increased to compensate for loss of xray photons from collimation In CT there is a wellestablished relationship between radiation dose pixel dimensions A SNR and slice thickness T 2 SNR D olt A3T Signal to Noise Ratio SNR Signal mean photons used to produce the imageunit area Noise injects a random or stochastic component into an image many sources SNR signalnoise increases with increase in the number of photons detected Quantum noise is the statistical fluctuation in the number of photons detected Quantum noise and structure noise both affect the conspicuity of a target Radiation Dose kVp kVp not only controls the close but also controls other factors such as image contrast noise and xray beam penetration through patient Parameter 80 W 120 kV 140 kV Image Contrast est Intermediate Poor Noise Most Average Least Penetration Least Average Most Patient Dose Lowest Intermediate Highest per mAs 17 Effective Dose Comparison with Chest PA Exam E I t f Approx period of Procedures Eff Dose mSv qu39va en 039 0 background chestxrays radiation Chest PA 002 1 3 days Pelvis 07 35 4 months Abdomen 10 50 6 months CT Chest 8 400 36 years CT AngWe or 1020 500 45 years elvrs Typical Background Radiation 3 mSv per year Some CT Artifacts Physics based beamhardening partial volume effects photon starvation scatter undersampling Scanner based centerof rotation tube spitting helical interpolation conebeam reconstruction Patient based metallic or dense implants motion truncation Beam Hardening Physicsbased CT Artifacts 0 unique to Xray due to preferential energy absorption of lower energy photons unlike nuclear imaging Tube spectrum gt U C 0 a C CD gt a I D I After 200m water After 30cm water BeamHardening Causes 39dishing39 artifact in uniform cylinders Can be corrected by scanner calibration tables 18 Photon Starvation Occurs when insuf cient photons reach the detectors typically across the shoulders usually corrected by 39adaptive ltration39 Cente rofRotation Similar artifacts occur with de cient detector channels and tube 39spitting39 Metallic Objects Occur because the density of the metal is beyond the normal range that can be handled Additional artifacts from beam hardening partial volume and aliasing are likely to compound the problem Patient Motion 19 Breathing artifacts in CT at the dome of the diaphragm and other areas Respiratory motion during a helical CT scan can lead to artifacts Tru n catio n o For accurate topographic image reconstruction the entire patient must be viewed from every direction 0 Unfortunately the patient port diameter is typically 70 cm in diameter while the imaging field of view FOV is typically 50 cm in diameter Extended Field of View EFOV CT Reconstruction cnverv sr svsnnsrnca uniusrsity fwashingtnnmedttr a a an 15175 Im26 nrnv 7 acm PET At Extended Field of View EFOV CT reconstruction for an offset 50 cm diam plastic phantom kv ran mA asn large 39 5 mnM5SE nqau1 quotIt n n Wl nss39ui 172a52nsa Not used for diagnostic CT yet Port 70 cm I CT FOV 50 cm Mean value m not constant for regions 1 2 and 3 Some Challenges in CT Patient radiation exposure Information overload from modern multislice scanners The pace of technological progress 20


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