Computed Tomography. What a progress

Transcription

1 IAEA RER/9/135 COURSE ON OPTIMIZATION IN COMPUTED TOMOGRAPHY Sofia, Bulgaria, 2017 Computed Tomography (a little bit of all) Dean Pekarovič UMC Ljubljana, Institute of Radiology Quality and Safety office What a progress

2 CT tecnical evolution Future Technology FPD, more then 2 x-ray tubes, IT implementation of AI 320 Slice CT = 320 slice per revolution 256 Slice CT = 256 slice per revolution Dual Source CT = 2 x ray tubes coupled with 2 detector arrays 64 Slice CT = 64 slice per revolution 8-16, Slice CT = 8, 16, 32, 40 slice per revolution Four Slice CT = 4 slice per revolution Dual Slice CT = 2 slice per revolution Single Slice Spiral/Helical CT = 1 slice per revolution First Clinical CT 1. Generation: Rotate/Translate, Pencil Beam Pencil beam shape -pinhole collimator to ensure only a single beam of x-rays small detectors rarely detected any scattered radiation 1-2detectors- positioned in x ray tube direction 4.5 minute for capturing data and for one slice only X ray tube rotated only 180 degrees 2

3 2. Generation Rotate/Translate, Narrow Fan Beam Narrow Fan beam shape ( 10 ) 30 or more detector elements 20 sec/slice and up to 10 minutes for 40 slice reconstruction 180 degree rotation x-ray tube Very long reconstruction times Detectors were exposed to more scattered radiation Only for head and brain due to its size and limited movement linear movement of the x-ray tube and detectors at each projection angle 3. Generation Rotate/Rotate, Wide Fan Beam Wide Fan beam shape (40 60 ) Next data is collected when x ray tube and detectors changed direction detector elements opposite to x-ray tube 360 degree rotation; rotation/table movement, 1 Gantry rot.= 1 slice Acquisition time to less than 20 seconds Reconstruction time for 1 slice is 1 second Not only for head, but for body scans too : ring artefacts! 3

4 4. Generation Rotate/Stationary Were developed minimize the ring artifacts Fixed ring with detector elements, around patient (~4800) X ray tube is rotating less radiation is needed Detector are capable to detect more radiation By removing the detectors from the rotating gantry and putting them in a stationary ring around the patient, detectors were able to maintain calibration 1 minute for more slices 5. Generation EBCT Stationary/Stationary Wide Fan beam shape CT scanner is composed of no moving parts. Like one Huge x-ray tube. Behind the patient, there is an electron beam, which ejects electrons. This electron beam is electronically deflected down, away from the patient, and makes contact with a large, half-circle tungsten target ring that encircles the patient. The interaction of the electrons with the target ring generates an x-ray beam, which travels through the patient s chest and is detected by a detector ring on the opposite side. Scan time is about 50 msec 4

5 6. Generation Spiral/Helical Slip Ring technology eliminate the time necessary for braking, and reversal of spooled cables X-ray source and detector array rotate continuously as the patient table is moved progressively through the scanner. Faster, helical (spiral or volume) CT, which offered greater spatial and temporal resolution Use of principles of the 3. and 4. generations with the slip ring technology 7. Generation Multiple Detector Array Cone Beam System can acquire an outstanding amount of information in a very short time span, requiring a much higher level of sophistication in the reconstruction process (slice competition between vendors z coverage). Spiral MDCT - submillimetre resolution during one breath hold. 5

6 CT generations overview Source Detector Source detector Generation Source Detector Advantages Disadvantages Collimation collimation movement Scattered Single x-ray Move linearly and Slow 1 G Pencil beam Single None energy is tube rotate unison ( 2.5 min) undetected Fan beam, not Lower efficacy Single x-ray Cannot collimate Move linearly and 2 G enough to cover Multiple Faster then 1G and larger noise tube detectors rotate unison FOV (10 sec) More expensive Single x-ray Fan beam, enough Cannot collimate Faster then 2G 3 G Many Rotate in synchrony then 2 G, low tube to cover FOV detectors,slip RING efficacy (0,5 sec) High scattering - Single x-ray Fan beam covers Stationary ring of Cannot collimate Fixed detectors, Higher 4 G detector are not tube FOV detectors detectors source rotates efficacy collimated (1 sec) Many Tungsten High cost, Stationary ring of Cannot collimate 5 G anodes in a Fan Beam No moving parts Extremely fast difficult to detectors detectors single large tube calibrate 3G/4G plus 6 G A bit more 3G/4G 3G/4G 3G/4G 3G/4G LINEAR Table Fast 3D images Spiral/Helical expensive MOTION 7 G Single x-ray Multiple arrays of Collimated to source Cone -Beam 3G/4G/6G motion Fast 3D images Expensive Multislice tube detectors direction CT components 6

7 CT needs In-plane resolution: mm Nominal slice thickness: S = mm Effective slice thickness: Seff = mm Tube (max. values): 120 kw, 150 kv, 1300 ma Effective tube current: maseff = 10 mas 1000 mas Rotation time: Trot = s Simultaneously acquired slices: M = Table increment per rotation: d = mm Pitch value: p = (up to 3.2 for DSCT) Scan speed: up to 73 cm/s Temporal resolution: ms Ref: Kachelrieß, Basics of Tomography 1 X ray Tube 7

8 X Ray Tube Characteristics X Ray Tube Tube voltages from 70 to 150 kv High instantaneous power levels (typ. 50 to 120 kw) ma High cooling rates (typ. > 1 MHU/minute) Must with stand centrifugal force Fast Tube Cooling Rates. Centrifugal force at 550 mm with 0.5 s: F = 9 g with 0.4 s: F = 14 g with 0.3 s: F = 25 g with 0.2 s: F = 55 g 8

9 State of The art CT performance GE Revolution Philips IQ Spectral Siemesn Force Toshiba One Vision Ref: CT scanners for, King's College London Detector needs Available as multi-row arrays Very fast sampling (typ. 300 μs) Favourable temporal characteristics (decay time < 10 μs) High absorption efficiency High geometrical efficiency - narrow gaps between active elements High count rate (up to 109 cps*) Small physical size Ref: Kachelrieß, Basics of Tomography 1 9

10 Detectors Properties: High detection efficiency(~90%). High geometrical efficiency (~80%). Small physical size of detector elements. Most commonly used detectors in new scanners. Hybrid, Matrix, Phased Array Matrix Phased Array Hybrid Hybrid 1. CT Generation 2. CT Generation 10

11 Detector Configuration 2 Hybrid Matrix 32 SLICE CT 64 SLICE CT Detector array configurations for state of the art MDCT 11

12 Detector Configuration Represents number of data channels along patient length (or z-axis) multiplied by the effective detector row thickness of individual data channel. 4 x x x 1.5 GE Philips Siemens Toshiba Detector configuration Collimation N X T (mm) Detector Configuration Detector Configuration Detector - collimation Siemens Sensation 64 Possible Configurations at spiral mode: 64 x 0.6* (19.2 mm nominal beam width ) Modes: 0.6, 0.75, 1, 1.5, 2, 3, 4, 5, 6, 7, 8, 10 mm Thinner slices, less volume per rotation, longer time for same volume. 24 x 1.2 (28.8 mm nominal beam width ) Modes : 1.2, 1.5, 2, 3, 4, 5, 6, 7, 8, 10 mm Bigger volume per scam, thinner slice is 1,2 mm. *Z-flying focal spot: In real 32 x 0.6 mm 12

13 Detector configuration- Impact on dose! 15 % 27 % Fixed mas AEC Off What is important? High detection capability : Detector must percept big amount of radiation in change it to light. Faster it is better the detector is. Afterglow : When radiation is converted to light, detector starts to glow. When detector is black" again, it can accumulate new amount of radiation. High Image Resolution: Number of pixels, number of channels,, number of rows in size of element below millimetre define ImQ. 13

14 Collimators Before Patient Single array collimator determine slice thickness MDCT different solutions by vendors Behind patient( reduce scatter radiation to improve contrast). Cone-Beam Artefacts mm from iso center 4 slice CT/ no processing 16 slice CT/ no processing 16 slice CT/ with processing 14

15 Anti-scatter grids 1D Anti Scatter Grid 2D Anti Scatter Grid As multi-slice scanners with large-area detectors have advanced, so has the need to reduce x-ray scatter through the use of improved anti-scatter grids. Increases low contrast for larger patients Anti-scatter grids are aligned to the detector pixels. Anti-scatter grids reject scattered radiation. Bowtie filter 15

16 z-over-scanning Helical scanners must irradiate a larger volume of patient than is ultimately displayed in the image data set : Necessary evil to avoid having incomplete data for reconstruction of first and last slices of interest. Typically an extra rotation is required a beginning and end of imaged volume. Pitch Is the Table Feed per gantry rotation divided by the beam width/collimation. Pitch is the ratio of two distances and therefore has no units. Users should monitor other parameters when changing Pitch. The scanner may or may not automatically compensate for changes in Pitch (for example, by changing the tube current). Pitch > 1 Some view angles are not covered by the beam width at certain table positions Pitch = 1 No overlap of Beam Width at each view angle and no view angles not covered at certain table positions Pitch < 1 Beam Width has some overlap at each view angle from rotation to rotation 16

17 Pitch (helical) Pitch (IEC,2003) T = Detector width in z axis N = nu. Of detector elements per rotation (N>1 at MDCT), N x T full width of detector during acquisition I = table speed in z direction in one full rotation mm/rot N*T = collimated x-ray beam width I = 15mm/rot N = 4 data channels T = 2,5 mm detector width P = (15 mm/rot) / (4 x 2,5 mm) P = 15/10 =1,5 Collimator and detector pitch T = table movement in 360⁰ ration of gantry C = width od collimation D = width of detector Collimator pitch = T/C Detector pitch = T/D Collimator pitch = Detector pitch / N Ref.: The Essential Physics of Medical Imagining;Bushberg et.all N = no. of detectors E.G.: FOV 20 mm Table speed 20 mm/rot Collimation 4 x 5 mm Collimator pitch = 1 Detector pitch = 4 High or low pitch? 17

18 Resolutions Image Quality Spatial Resolution Contrast resolution Temporal resolutions Spatial Resolution (SR) Ability to observe or separate small objects Minimal observable size depends on object contrast (CNR) Spatial resolution is also called high contrast resolution Spatial resolution is described by In axial plane: point spread function or MTF In z direction: slice sensitivity profile Gives the actual width of the imaged slice Usually quantified as FWHM 18

19 Factor affecting SR Detector pitch : is centre-to centre spacing along the array. Detector aperture : is the width of active element of one detector. Smaller detectors increase cut off(nyquist) frequency of the image (improve SR). No. of views : more views less artefacts (aliasing). mas: big influence in same FOV Focal spot size (FS): larger FS cause more geometric unsharpness and reduce SR Object magnification : Increased magnification amplifies the blurring of the focal spot. Slice thickness : the slice thickness is equivalent to the detector aperture in CR-CA direction. Large slice thickness Helical pitch : greater pitch reduce SR (increased slice sensitivity profile) Reconstruction kernel : Bone filters have the best SR and soft filters lower SR Pixel Matrix : direct influence (256², 512² view 1024²) FOV : 10 cm FOV in 512² ;0,2 mm pixel size 35 cm FOV in 512² ;0,7 mm pixel size The limiting factor is noise Contrast Resolution (CR) Difference in signal between a structure and its environment The contrast resolution of CT is not intrinsically high, because the difference in x-ray attenuation between different tissues is generally small (vs. MRI) High nu. of projections vs. object overlapping e.g. chest x ray 25 mas 200 mas The limiting factor is Noise -Dose 19

20 Factor affecting CR mas : directly influence the number of x-ray photons used to produce image, thereby affecting the SNR and CR (double mas-41% better CR-but) Dose : increase linearly with mas per scan Pixel size (FOV) : if FOV increase, pixel dimension increase, and no. of x- rays passing through same pixel increase(scan parameters and patient size is same) Slice thickness : thicker slice use more photons and have better SNR Reconstruction filter : Bone filters has lower CR and soft filters improves CR Patient size : Larger patients attenuate more x-rays, resulting in detection of fewer x-rays, lower SNR and therefore CR Gantry rotation speed : compared 1 s with 0.5 s results in reduced mas to produce each CT image, reducing CR Temporal resolution Ability to resolve fast moving objects For image reconstruction at least 180 is needed (180 MLI) If in plane are no moving organs 360 State of the art CT 3-4 rotations/s Influences motion artefacts Influences motion induced blurring Particularly important for moving organs (heart) children, uncooperative patients Requires fast tube rotation (or 2 tubes ) 20

21 Temporal resolution CCTA Ref: CT scanners for, King's College London Reconstruction settings Reconstruction order ma prescription under TCM may be linked to specific recon Reconstruction kernel determines spatial resolution, image noise and texture filtered back-projection, iterative reconstruction selection may affect ma prescription under TCM Recon mode e.g. Recon Full, Recon Plus selection may affect ma prescription under TCM Ref: E. Castellano, IAEA CT Workshop 21

22 Reconstruction algorithm Soft Standard Smooth - soft tissuee Detail Lung Bone Edge Bone Plus Sharp - bone GE recon algorithms Reconstruction kernel Represents a feature on the scanner which influences the smoothness and sharpness of images in transverse plane B20 B31 B50 B70 Lower noise Poorer edge delineation Better contrast Higher noise level Better edge delineation Poor contrast Standard Details Lung GE Philips Siemens Toshiba Algorithm Reconstruction filter Kernel Filter Convolution 22

23 Kernels Head 1 vendor Siemens H 40s H (Head), B (Body), C (Child Head), S (Special Applications) s standard mode f- fast mode h- high resolution mode H21, H31, H41 are like H20, H30, H40 but with finer grain noise and a milder edge enhancement. H22, H32, H42 don t include iterative beam hardening correction (PFO). Reconstruction speed is faster, but the reconstructed images may contain significant beam hardening artifacts. H37 is comparable to GE Soft H45 is intermediate sharpness between H40 and H50 H47 is comparable to GE Standard H48 is like H47 but a bit sharper Reconstructed slice thickness 10 mm 5mm 2mm 1 mm 0,5 mm 23

24 Reconstruction Possibilities-AIP 1. 0,8 mm Coronal view/plane 2. 4,0 mm Coronal view/plane ImQ more soft, better Contrast Resolution, almost like on axial planes Average Intensity Projection MIP MIP (cca 10 % basic image info) AIP (2,5 mm) MIP (2,5 mm) MIP 5 mm MIP 10 mm MIP 15 mm Overlaping with structures with high HU (White, Bone, other vessels..). 24

25 MIP slice Thickness 7mm /5 mm 2. 4mm/ 2mm Image reconstruction 1972 CT discovery Geofrey Hounsfield (UK) Technical part Allan Cormack (SA) Mathematical reconstruction 25

26 Hounsfield Cormak Improved Radon transformation 26

27 Mathematics (grrrr) I 0 Intensity of x-ray beam entering object h material thickness μ μ attenuation coefficient (tissue characteristic) I Intensity of x-ray beam exiting object I I 0 e h More then one tissue I 0 I I 0 e h μ is average value for all tissues in path 27

28 Matrix can be used for calculation I 0 I I μ1 μ2 μ3 μ4 μ5 μ6 I I 0 0 e e ( x x x3 4x4 5x5 6x6 ( ) x ) Reconstruction = using mathematical algorithms to reconstruct an image from projections 28

29 CT from X ray tube to Image RECON TYPE : FBP IR Noise Reduction X ray Tube: kv mas Bow Tie Filter Detector : Data Collection DAS: Detector Efficiency Multiple calculations

30 Reconstruction algorithms FBP Filtered Back projection IR Iterative reconstruction Filtered Back Projection Smeares measured projection data back to image More projections better image 30

31 FBP 2 projections OBJECT FBP M M ,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,4 1,4 1,4 1,4 1,4 1,4 1,4 1,4 1,4 1,4 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,0 1,1 1,1 1,1 1,1 1,1 1,1 1,1 1,1 1,1 1,1 1,7 1,7 1,7 1,7 1,7 1,7 1,7 1,7 1,7 1,7 2,2 2,2 2,2 2,2 2,2 2,2 2,2 2,2 2,2 2,2 1,7 1,7 1,7 1,7 1,7 1,7 1,7 1,7 1,7 1,7 1,1 1,1 1,1 1,1 1,1 1,1 1,1 1,1 1,1 1,1 1 st proj. 2 nd proj. 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,7 2,2 1,7 1,1 1,0 1,4 0,9 1,0 1,0 1,1 1,4 1,6 1,4 1,1 1,0 1,2 1,0 1,0 1,0 1,1 1,4 1,6 1,4 1,1 1,0 1,2 1,0 1,0 1,0 1,3 1,6 1,8 1,6 1,3 1,2 1,4 1,2 1,2 1,2 1,1 1,4 1,6 1,4 1,1 1,0 1,2 1,0 1,0 1,0 1,1 1,4 1,6 1,4 1,1 1,0 1,2 1,0 1,0 1,0 1,1 1,4 1,7 1,4 1,1 1,1 1,3 1,1 1,1 1,1 1,4 1,7 2,0 1,7 1,4 1,4 1,6 1,4 1,4 1,4 1,7 2,0 2,2 2,0 1,7 1,6 1,8 1,6 1,6 1,6 1,4 1,7 2,0 1,7 1,4 1,4 1,6 1,4 1,4 1,4 1,1 1,4 1,7 1,4 1,1 1,1 1,3 1,1 1,1 1,1 IMAGE FBP limitations noise,artefacts.. FBP M IMAGE M OBJECT 1,1 1,4 1,6 1,4 1,1 1,0 1,2 1,0 1,0 1,0 1,1 1,4 1,6 1,4 1,1 1,0 1,2 1,0 1,0 1,0 1,3 1,6 1,8 1,6 1,3 1,2 1,4 1,2 1,2 1,2 1,1 1,4 1,6 1,4 1,1 1,0 1,2 1,0 1,0 1,0 1,1 1,4 1,6 1,4 1,1 1,0 1,2 1,0 1,0 1,0 1,1 1,4 1,7 1,4 1,1 1,1 1,3 1,1 1,1 1,1 1,4 1,7 2,0 1,7 1,4 1,4 1,6 1,4 1,4 1,4 1,7 2,0 2,2 2,0 1,7 1,6 1,8 1,6 1,6 1,6 1,4 1,7 2,0 1,7 1,4 1,4 1,6 1,4 1,4 1,4 1,1 1,4 1,7 1,4 1,1 1,1 1,3 1,1 1,1 1,1 31

32 FBP projections FBP what is in calculations Point source Pencil Beam Point Voxel Point detector 32

33 What contribute to noise on image Technology upgrade Improvements in x ray tubes Bowtie filters Rotation time/ collimation Antiscatter grid Detector efficiency (spectral imagining) High pitch scanning modes Reconstruction algorithms Not possible to change Patient (positioning, metal implants, moving.) FBP way to.. Lowering the dose to get still diagnostic quality Research, evaluation years Adopting different noise level (age, indication orientated ). Goal Define lower limit in LCR, no decrease SR mas ,7 mgy Pekarovic, Ref:Ramirez-Giraldo, IAEA Sofia Simens USA Ref:Radiation dose from MDCT, Tack et all 33

34 FBP limitations - detailed Simplified Geometry Infinetly small: Focal spot Thin beam Point detector Physics Perfect sampling Simplified physics Scatter Absorption Beam filtration Noise In reconstrucions it use even measurements with very low response same good as perfect ones. High noise is compensating with increase in dose : kv mas Rot time Pitch etc.. Artifacts In this type of reconstruction we can`t avoid : Metal artifacts Streak artifact Beam Hardening Windmil HU are higher in tissue where artifacts are Next.. Maybe we got maximum output from FBP ImQ Dose Is Balance Achieved? Is iterative reconstruction a solution? 34

35 Iterative reconstructions (IR) WHY IR?? Dose - CT highly contribute to population dose Number of CT scans rapidly increase More indications Post processing power of CT, visibility. Fast and accurate Experience from NM Possibility to reduce noise Market needs Iterative reconstruction 1st generation using statistical model 2nd generation using mathematical model of imaging system Vendor 1st generation 2nd generation Philips idose(irt) IMR General Electric ASIR ASIR-V, Veo* Siemens IRIS SAFIRE, ADMIRE Toshiba QDS/Boost 3D AIDR 3D Not included SIEMENS Itrim (cardiac, TR) 35

36 Iterative reconstructions Forward projection instead of back Several steps: Assume an image (empty or FBP), Compute projections from the image, Compare to the original projection data and Update the image based on the difference between calculated and actual projections. SIR 1 st Generation SIR uses FBP data set as the building block for image reconstruction assuming the beam to be a perfect point source. Use this or raw data to compare to other data (e.g. forward projection data, noise smoothed image) --- compares, makes corrections, does that again, until?? It aims to improve image quality by focusing on noise reduction. 36

37 SIR Comprehensive noise model includes: Photon, Beam hardening, electronic noise Severe photon starvation (such as through metal), which creates star, streak other artefacts. Can use adjacent projections and voxels to reweight. Attempts to preserve edges. How Really Works To complicated to understand without appropriate background knowledge Ask Medical Physicist to explain it softly. Important to know : How it looks on Images 37

38 Difference in FBP IR Point source Pencil Beam Point Voxel Point detector Real source Fan Beam 3D Voxel shape Detector Geometry Tube and detector response Huge calculations SIR FBP all vendor is same principle IR different approaches like : Assumptions, calculations (physics and optics). Algorithms, kernels (filters). Blending options. Choices (levels, percentages). Guidance (kv, mas, etc.). Software, hardware evolving rapidly. 38

39 Advantages of IR Dose Noise Artefacts Speed ImQ IR statistical Denoising 39

40 IR Homogenous phantom FBP 120 kv, 300 mas, CTDI vol 12,5 IR -50 % (SIR)- half Dose More Noise less softening vs. IR 100 % IR -100 % (SIR) half Dose Less Noise- softening vs. FBP Ref: Hara et all, Iterative Reconstruction Technique for Reducing Body Radiation Dose at CT: Feasibility Study, AJR:193, September 2009 Validation? 50 mgy FBP Average: 35; Min/Max : % IR - 12,5 mgy Average: 31 Min/Max :7 51 In both cases: ROI in Center : HU ± 3 40

41 Low Contrast Resolution 50 mgy FBP 50 % IR 12,5 mgy 4 x less dose same contrast- same ImQ IR and SR A : regular dose B : low dose C : low dose +IR 60% D : low dose + IR 100% Ref :O.Rapalino, S.Kamalian et all, Cranial CT with Adaptive Statistical Iterative Reconstruction: Improved Image Quality with Concomitant Radiation Dose Reduction, AJNR

42 Way to go Protocols can remain unchanged Keep the same dose better ImQ Lower dose same ImQ 120 kv 10mAs 0,5 rot. time 0,625 slice thickness 0,61 msv Image Courtesy, dr.barbau,cnn,france 0,61 msv DLP 36,02 On Images- Low Contrast Rresolution Low Dose CT, FBP: no IR 120kv, 3,75 mm CTDI 8 Low Dose CT, FBP: + IR 120kv, 3,75 mm CTDI 8 Routine CT, FBP: no IR 140kv, 3, mm CTDI 22 Same Dose Same ImQ Ref : Hara et all,iterative Reconstruction Technique for Reducing Body Radiation Dose at CT: Feasibility Study, AJR :September 2009, Volume 193, Number 3 42

43 On Images Spatial Resolution Low Dose CT, FBP: no IR 120kv, 3,75 mm CTDI 11 Low Dose CT, FBP: with IR 120kv, 3,75 mm CTDI 11 routine CT, FBP: no IR 140kv, 3, mm CTDI 20 Edge of cyst soft or sharp? 1 st generation FBP FBP low dose FBP AIDR processing ASIR 67 % dose reduction IRIS 60% dose reduction 43

44 IR and CR A : regular dose B : low dose C : low dose +IR 60% D : low dose + IR 100% Plasticy Waxy Blotchy Pixelated Ref :O.Rapalino, S.Kamalian et all, Cranial CT with Adaptive Statistical Iterative Reconstruction: Improved Image Quality with Concomitant Radiation Dose Reduction, AJNR 2012 Steps or % betwwen FBP and IR 44

45 2nd Generation Model Based IR Whilst SIR still relies on FBP data sets, model-based iterative reconstruction builds a forward projection using dedicated system optics, taking into account every x-ray projection (in its true three-dimensional domain), and produces an image based on the raw data. Multiple iterations are performed to correct the residual error between the forward projection and acquired image. Much more complicated algorithms These algorithms also incorporate statistical noise information in the reconstruction process. The combination of system optic modeling and statistical modeling helps in noise reduction and results in truer image characteristics compared to FBP and SIR. In addition, MBIR also accounts for noise from photon flux as well as system noise (e.g. electronic noise) from the CT system itself. AEC - effect on IR Courtesy of GE 45

46 IR and CR 100 mas 25 mas 12 mas Ref : P.Kinahan, The Future of Low Dose CT, Univ. Of Washington 1st and 2nd Sub msv FBP ASIR 50% VEO Low dose GE Analytical rec. idose IMR Images: courtesy of Kevin Brown, Philips 46

47 1st and 2nd Sub msv FBP IRIS SAFIRE IV Images: courtesy of Thomas Flohr and Rainer Raupach, Siemens Analytical rec. QDS/Boost AIDIR3D Images: courtesy of proff.sujith Seneviratne,Toshiba IR and patient size 40-cm phantom FBP vs IR Diameter % noise reduction FBP IR Mild 24 cm 23% 30 cm 36% 40 cm 51% IR Std. IR Strong Ref;:Kim M. Et all, Adaptive Iterative Dose Reduction Algorithm in CT: Effect on Image Quality Compared with Filtered Back Projection in Body Phantoms of Different Sizes, Korean J Radiol

48 Recipe Dose 100 % 50 % 25 % 120 kv and 100 Kv Different reconstruction algoritms Slice thickness 1mm and 3mm 3 mm FBP IR 2 IR 3 1 mm FBP IR 2 IR 4 IR may reduce CT patient dose depending on the clinical task, patient size, anatomical location, clinical practice But An aggressive dose reduction can reduce lesion detection So.. A consultation with a radiologist and a physicist should be made to determine the appropriate dose to obtain diagnostic image quality for the particular clinical task. 48

49 ImQ- SR Courtesy of Ge Healthcare IR + metal artefact reduction 49

50 2 nd Artefacts and bones WW/WL 350/50 WW/WL 2000/400 Ref : Intel Xeon Processors + GE, Whitepaper, New Levels of CT Image Performance and New Levels in Radiation Dose Management 3 D (VR +MIP) + IR Ref: Philips PM Ref :Improved Delineation of Arteries in the Posterior Fossa of the Brain by Model-Based Iterative Reconstruction in Volume-Rendered 3D CT Angiograph, Machida et all,

51 IR Impact on CTDIvol and DLP Ref : IR dose reduction factors Ref: Qiu and Seeram; Does Iterative Reconstruction Improve Image Quality and Reduce Dose in Computed Tomography? 51

52 DECT First article! Two pictures are taken of the same slice, one at 100 kv and the other at 140 kv so that areas of high atomic numbers can be enhanced. Tests carried out to date have shown that iodine (Z = 53) can be readily distinguished from calcium (Z =20). Hounsfield 4 decades ago DECT Basic Principle The principles of DECT are based largely on the photoelectric effect and can be achieved by exploiting the energy-dependent attenuation of materials when exposed to 2 different photon energy levels ( 2 basis materials). These physical principles can be exploited for in vivo human imaging, because DECT is based on dissimilar tissue characteristics with respect to their energy dependent x- ray attenuation. 52

53 DECT - basics predominant at higher x-ray energy levels Materials in human body hydrogen (Z=1) calcuim (Z=20) carbon (Z=6) iodine (Z=53) nitrogen (Z=7) oxygen (Z=8) higher at the lower x-ray energy levels Compton scattering is proportional to the electron density. Lower noise on image. Photoelectron absorption is directly proportional to the atomic number and electron density. Emission of characteristic radiation, Higher noise on image. Can we see more with DECT Materials can be chosen arbitrarily, as long as their K edges are sufficiently different (i.e., attenuation profiles), such as water and iodine. Material with an attenuation spectrum different than that of the chosen basis materials will be reflected as a combination of the 2 basis materials Ref: Danad et all, New Application of Cardiac CT 53

54 DECT + Tube current and tube filtration can be optimized for each tube potential independently. Relatively low degree of spectral overlap, which improves CNR in material-specific images. Beam-hardening corrections are applied prior to image reconstruction, allowing material-specific images to be created in the image domain. - A 90 phase shift between low- and high-energy data. Simultaneous use of both x-ray sources allows scattered radiation whose original primary photon came from one tube to be detected by the detector of the other tube, requiring specialized scatter correction. Ref: Danad et all, New Application of Cardiac CT DECT + Data acquisition of the low- and highenergy data set allows dual-energy material-decomposition algorithms to be implemented by using either projection data or reconstructed images. Reduces beam-hardening artefacts in calculated virtual monoenergetic images. Same HU. - There is no current modulation. Relatively high overlap of the energy spectra (80 kv and 140 kv). Gantry rot/time is 0,5 s between (80 kv and 140 kv) prolongs time of aquisotion Ref: Danad et all, New Application of Cardiac CT 54

55 DECT + Simultaneous data acquisition of the low- and high-energy data set. All image data are acquired in a manner that supports material-specific imaging. - Relatively high overlap of the energy spectra. Noise level may differ between low- and high-energy images. Ref: Danad et all, New Application of Cardiac CT DECT + Can be performed on any CT scanner - Single source-detector pair with tube voltage switching between sequential gantry rotations. Any patient motion occurring between the two scans may cause severe degradation of the resultant images and material composition information. a change in contrast opacification Ref: Danad et all, New Application of Cardiac CT 55

56 DECT DECT mix series 20 % - 80 % 80 kv 150 kv 60 % - 40 % 80 kv 150 kv 90 % - 10 % 80 kv 150 kv 56

57 DECT Mono energetic Monoenergetic 44 kev Monoenergetic 50 kev Monoenergetic 60 kev Monoenergetic 80 kev Mono-energetic reconstructions allow to reconstruct the images at any given virtual kev setting of a CT scan to optimal contrast needed for diagnostic purposes. Low kev reduced volume of CM -higher High kev artefacts/obese patient Monoenergetic 100 kev Monoenergetic 120 kev Monoenergetic 150 kev Monoenergetic 189 kev SECT DECT example 3 cm scan length Arterial phase 120 kv 80 kv 150 kv 80 kv mixed 57

58 SECT DECT example DECT postprocessing Iodine Map Metal artefact Table/bone removal Perfusion - PE Renal Stones 58

59 Flat Panel Detector- - CT Detector Detector Matrix: elementov (40cm x30cm). Effectiv esize (magnification) :25 x 25 x 18cm. Element size : µm. FOV in z axis/1 rotation: 18 cm. Spatial Resolution (x, y, z): µm. ANY ANGLE frames/second( CT 1000x/s). Speed big FOV in detector. Z axis Organs: Heart, Brain, liver in 1 rotation. Diagnostic Conventional modalities CT Angiography Fluoroscopy image in real time Flat Panel Detector- - CT Quick detection. Big detector instead rows (30x 40cm, volume). Bigger FOV in 1 rotation (in z-axis). More radiation contribute to image formation vs. Amorphous Siliceous. Less noise contribute to wider dynamic range (grey levels - attenuation). Faster readout(cardiac, hold breath, moving artefacts...). Smaller pixel size then in crystal(powder) structure. Smaller Pixel- higher Spatial resolution. Similar images like in Interventional and Cardio modalities. 59

60 Flat Panel Detector- CT Bad material became old (improvements?). Worse scintillation. Lower Contrast resolution (on thinner slices). Higher dose for equal SNR ratio in comparison with MDCT. Longer time of acquisition. SR and LCR Up to 30 lp/cm in entire FOV (npr. 18cm). Routine CT aprox LP/CM UHR up to 25 lp/cm limted by FOV size. MDCT ± 3 HU FPCT ± 5 HU or more 60

61 Basic difference All standard dose description are not usefull (CTDI, DLP), NEW ONE should be estimated. Some image characteristic are not equal (training for radiologist) between MDCT in FPCT. Image formation is different. Especially for Interventional (angio), Cardio, Skeletal and ORL. Automatics : MDCT vs FPCT Old CT: constant kv, mas and time. Modern CT : constant kv (2 sources, 2 kv), ma, constant time. FPCT: kvp, ma, constant time. Conclusion... Correct use of technology requires didactic education, hands on training and continuous re training of users. 61

62 Thank you! 62