Imaging And Detectors For Medical Physics Lecture 3: X-ray Imaging
Joint CI-JAI advanced accelerator lecture seriesImaging and detectors for medicalphysicsLecture 3: X-ray imagingDr Barbara Camanzibarbara.camanzi@stfc.ac.uk
Course layoutAM 09.30 – 11.00PM 15.30 – 17.006th JuneLecture 1: Introduction tomedical imagingLecture 2: Detectors formedical imaging7th JuneLecture 3: X-ray imagingDayWeek 18th JuneTutorialWeek 213th JuneLecture 4: Radionuclides14th JuneLecture 5: Gammacameras16th JuneLecture 7: PETLecture 6: SPECTWeek 322nd JuneTutorialPage 2/49
Books1. N Barrie Smith & A WebbIntroduction to Medical ImagingCambridge University Press2. Edited by M A FlowerWebb’s Physics of Medical ImagingCRC Press3. A Del GuerraIonizing Radiation Detectors for Medical ImagingWorld Scientific4. W R LeoTechniques for Nuclear and Particle Physics ExperimentsSpringer-VerlagPage 3/49
X-ray in the bodyRef. 1 – Chapter 2, Ref. 2 – Chapter 2 X-rays going through patient’s body get attenuated:𝐼 𝑥 𝐼0 𝑒 𝜇 𝐸 𝑥𝐼0 X-ray fluence in entrance𝐼 𝑥 X-ray fluence at position 𝑥 fluence in exit𝜇 𝐸 X-ray linear attenuation coefficient X-ray linear attenuation coefficient μ 𝑐𝑚 1depends on X-ray energy In tissue mass attenuation coefficient often used𝜇/𝜌 𝑐𝑚2 𝑔 1 , with 𝜌 𝑔/𝑐𝑚3 tissue densityPage 4/49
X-ray transmission imaging Basis differential absorption of X-rays by tissues for ex. bone absorbs X-ray more than soft tissueTissue𝝁 𝒄𝒎 𝟏𝑰 𝒙 /𝑰𝟎 𝒙 𝟏 𝒄𝒎Difference tomuscle (%)Air0.0001.0 20Blood0.1780.837 0.2Muscle0.1800.8350Bone0.4800.619-26 Contrast agents chemicals introduced in patient’sbody to enhance contrast between tissuesPage 5/49
X-ray transmissionimage formation Image formation:1. X-rays from source directed toward patient some X-raysabsorbed some X-rays transmitted2. X-rays transmitted detected in exit from patient3. Measured in exit from patient fluence distribution linearattenuation coefficient distribution Some X-rays scattered inside patient image noise /backgroundPage 6/49
X-ray imaging techniquesPlanar radiographyComputed Tomography Image 2D projection of alltissues between X-ray source anddetector X-ray source and detector fixed Image 3D image of body region X-ray source and detector rotateat high speed around patient patient moved in third direction Disadvantage respect to planarradiography much higher dosePage 7/49
Other X-ray imaging techniques X-ray fluoroscopyImages are acquired continuously to study passage of X-raycontrast agent through GI tract Digital mammographyImages are acquired with lower X-ray energies than standardX-ray scans to obtain images with much finer resolution Digital subtraction angiographyImages are acquired at extremely high resolution to imagevasculature Digital X-ray tomosynthesisHybrid planar radiography – CT: fixed screen rotating sourcePage 8/49
X-ray tubeX-ray source for transmission imaging X-ray tube Cathode filament focusing cup Anode target that rotates at highspeed to reduce localised heat Filament and target usuallytungsten Efficiency for e conversion inX-rays 1%, rest dissipated inheath Strong vacuum inside tube unimpeded path between cathodeand anode Oil surrounding the envelope dissipates heat from anodePage 9/49
Materials for thefilament and target Tungsten: most commonly usedCharacteristicsAdvantagesEmission at 2000 oCHigh and stable e thermionic emissionMelting point 3370 oCCan withstand very high temperaturesgenerated in the anodeHigh 𝑍 74High X-ray production efficiency1Good thermal conductivity Low vapour pressureCan operate in very high vacuum1Bremsstrahlungyield increases with 𝑍 Molybdenum: used in digital mammography thatrequires very low energy X-rays less heatgeneratedPage 10/49
X-ray tube parametersTube parametersValuesAccelerating voltage Δ𝑉𝐶 𝐴 , kVp 25 140 kV1Tube current 𝐼 from the cathode 50 400 mA for 2D radiographyto the anodeUp to 1000 mA for CTExposure time125Limited by anode heatingkV for mammography, 140 kV for bone and chest These parameters are chosen by the operatoraccording to the specific application 2D radiography and CT scanners different set-up same X-ray tube cannot be used for bothPage 11/49
Power rating Power rating – DefinitionMaximum power dissipated in an exposure time of 0.1 s ExerciseQ What is the maximum exposure time of a tube with a powerrating of 10 kW, when operated at 125 kV with 1 A of current?What modality is this?A 𝑃𝑜𝑤𝑒𝑟 𝑑𝑖𝑠𝑠𝑖𝑝𝑎𝑡𝑒𝑑 𝑘𝑉𝑝 𝐼 125 𝑘𝑉 1 𝐴 125 𝑘𝑊𝐸𝑥𝑝𝑜𝑠𝑢𝑟𝑒 𝑡𝑖𝑚𝑒 𝑃𝑜𝑤𝑒𝑟 𝑑𝑖𝑠𝑠𝑖𝑝𝑎𝑡𝑒𝑑 𝑃𝑜𝑤𝑒𝑟 𝑟𝑎𝑡𝑖𝑛𝑔 𝑃𝑜𝑤𝑒𝑟 𝑟𝑒 𝑡𝑖𝑚𝑒 80 𝑚𝑠𝑃𝑜𝑤𝑒𝑟 𝑑𝑖𝑠𝑠𝑖𝑝𝑎𝑡𝑒𝑑 125Modality is CTPage 12/49
A couple of definitionsField-of-view (FOV)Penumbrahttps://en.wikipedia.org/wiki/Field of viewRef. 2 – Chapter 2.5.5 FOV of optical instruments orsensors solid angle throughwhich detector is sensitive toradiation solid angleimaged by the detector Penumbra 𝑃 unsharpness/ blurring in the image dueto finite size of X-ray source:𝑑1𝑃 𝑑2𝑃Page 13/49
Instrumentation for2D X-ray radiography X-ray tube: generates the X-ray beam Collimator: reduces patient’s dose and amountof Compton scattered X-rays Anti-scatter grid: reduces amount of Comptonscattered X-rays background/noise increases image contrast Digital detector: converts transmitted X-rays intolight and then into electric signal Read-out electronics: digitises and reads thesignal from the detectorPage 14/49
Collimators and gridsCollimators Sheets of lead placedbetween X-ray source andthe patient Restrict dimension of thebeam to the FOV in 1D or 2D reduce amount of X-raysreaching the patient onlyX-rays inside FOV reach thetissue dose reduced scattered reducedAnti-scatter grids Parallel or slightly divergentstrips of lead foil withaluminium spacers Amount of scattered X-raysabsorbed depends onlength, thickness andseparation of lead strips Some non-scattered X-raysare absorbed increase indose to get same imageintensity of one without gridPage 15/49
Detectors and electronicsRef. 1 – Chapter 2.7 Computed radiographyInstrumentation detector plate separate reader Digital radiographyInstrumentation detector and reader are one unit1. Indirect X-ray converted into light by scintillator light converted into electric signal by photon detector2. Direct X-ray converted into electric signal bymaterials such a:Se.Less efficient than indirect conversion devicePage 16/49
Signal-to-noise ratio (SNR)Ref. 1 – Chapter 2.8.1 Signal 𝑁 of X-rays arriving on detector Statistical fluctuations in number of X-rays detectedper unit area noise Statistical fluctuation follow Poisson distribution 𝜎𝑛𝑜𝑖𝑠𝑒 𝜇 with 𝜇 mean value𝑁𝑁𝑆𝑁𝑅 𝑁𝜎𝑛𝑜𝑖𝑠𝑒𝜇 Exercise: What is the dose increase if doubling 𝑆𝑁𝑅?A: 2 𝑆𝑁𝑅 2 𝑁 4 𝑁 4 𝑁 4 𝐷𝑜𝑠𝑒Page 17/49
Factors affecting SNR1. X-ray tube current 𝐼 and exposure time 𝑡𝑒 :𝑆𝑁𝑅 𝐼 𝑡𝑒2. X-ray tube kVp: the higher kVp the higher the X-ray energy greater penetration in tissue signal increases 𝑆𝑁𝑅increases in a non-linear way3. Detector efficiency: the higher the efficiency the more X-raysare detected signal increases 𝑆𝑁𝑅 increases4. Patient size and body part to be imaged: the greater thetissue thickness the higher the X-ray attenuation signaldecreases 𝑆𝑁𝑅 decreases5. Anti-scatter grid: attenuates Compton scattered X-rays reduces signal 𝑆𝑁𝑅 decreasesPage 18/49
Spatial resolutionRef. 1 – Chapter 2.8.2 Factors affecting spatial resolution:1. Set-up geometry: penumbra 𝑃 unsharpness / blurring inthe image due to finite size of X-ray source generates ideal set-up:a.b.c.Smallest possible X-ray spot sizePatient on top of detectorLarge distance between source and patient2. Detector’s properties: detector’s intrinsic spatial resolutionPage 19/49
Contrast-to-noise ratio (CNR)Ref. 1 – Chapter 2.8.3 Factors affecting CNR:1. X-ray energy: the higher the energy the more X-rays undergoCompton scattering CNR decreases2. FOV: up to 30 cm the larger the FOV the higher the numberof Compton scattered X-rays reaching the detector CNRdecreases; above 30 cm there is no change3. Thickness of body part being imaged: the thicker the sectionthe more X-rays undergo Compton scattering the more Xrays get absorbed CNR decreases4. Anti-scatter grid: reduces the Compton scattered X-raysreaching the detector CNR increasesPage 20/49
Computed Tomography (CT)scannersDiodeScintillatorCollimatorX-ray sourceCT scanner:- X-rays rotating source- Diametrically opposite detector unitMarket: 30,000 scanners worldwide, 60 millionsCT scans performed annually in USACourtesy Mike Partridge (Oxford)Page 21/49
Computed Tomography (CT)Ref. 1 – Chapter 2.12 Basic principle:– Conventional CT:1. Series of 1D projections at different angles is acquiredcontinuously by synchronously rotating the X-ray source anddetectors through one complete revolution around the patient2. The 1D projections are combined by the process of filteredbackprojection to form the 2D CT image, also called slice– Spiral / helical and multi-slice helical CT1. 2D slices are acquired as in conventional CT2. Multiple adjacent slices are acquired by moving the patient’scouch along the direction perpendicular to the slices’ planeto give 3D imagesPage 22/49
Instrumentation forconventional CTInstrumentationNotesX-ray tubeSame as in planar radiographykVp 80, 100, 120, 140 kVCollimatorsSame as in planar radiographyAnti-scatter gridsSame as in planar radiography butusually integrated in the detector arrayDetectorsOnly one detector unit 1D array ofseveral hundred 15 1 mm2 detectors1along circumferenceHeavy gantryHas fixed to it X-ray tube and detectorunit and rotates at high speed1Detector scintillator (converts X-rays into light) photodiode (converts light into electric signal) Note: detector’s orientation wider side (15 mm)along couch axis slice thickness determined bywidth of collimated beam that is 15 mmPage 23/49
Instrumentation for helical CT X-ray source and couch moved at the same time X-ray path helical Conventional CT set-up modified as follows:1. Power supply and signal transmission cables aresubstituted by multiple slip-ringsReason: impossible to have fixed cables for powersupply and signal transmission to read-out system2. X-ray tube: specially designed to withstand very hightemperatures in anodeReason: X-rays produced (almost) continuously no cooling period anode reaches very hightemperatures higher than in conventional CTPage 24/49
Instrumentation formulti-slice helical CT Same operation as helical CT but bigger detector unit larger volumes can be imaged in a given time Same set-up as of the helical CT but with differentgeometry of the detector unit 2D array of smallerdetectors1. Along couch axis detector size is much smaller (can be 0.5 mm) but there are multiple rows that cover up to 16 cm slice thickness determined by detector width smallerthan in helical CT2. Along circumference detector size (1 mm) and number ofdetectors per row are the same as for helical CT 1D arrayPage 25/49
Dual-source CT Dual-source CT 2 X-ray tube multi-slice detector chainsReason: increases temporal resolution 2 x temporal resolution ofsingle-source CTGantry’s rotation gravitational forces on scanner rotation speedlimited ( 100 160 ms for 180𝑜 ) temporal resolution limited Features:1. Set-up: 1 standard chain (can be used alone) 1 chain withnarrow-arc detector smaller FOV ( 2/3) (only used with other)2. Data acquisition modes:a.b.Single energy both tubes operated at same kVpDual energy tubes operated at different kVp 140 keV and 80 keV better contrast between different tissuesPage 26/49
2D image reconstruction in CTRef. 2 – Chapter 3.2 Data acquired and used to reconstruct image transmission measurements:1. Exit (attenuated) X-ray beam intensity detected2. Ratio attenuated (exit) / unattenuated (entry) X-raybeam intensity projections3. Reconstruction extract linear attenuationcoefficients from projections4. Image display of linear attenuation coefficients’distributionPage 27/49
Transmitted intensity Transmitted intensity 𝐼𝜙 𝑥 ′ :𝐼𝜙 𝑥 ′ 𝐼𝜙0 𝑥 ′ 𝑒𝑥𝑝 𝜇 𝑥, 𝑦 𝑑𝑦 ′𝐴𝐵𝐼𝜙0 unattenuated, entry intensityTaken from Ref. 2 pg. 104 𝑥𝑦 frame centred on body 𝑥 ′ 𝑦 ′ rotating frame centredon scanner X-ray source on 𝑦 ′ axis𝜇 𝑥, 𝑦 2D distribution of linearattenuation coefficients Assumptions:1. Very narrow pencil X-ray beam2. Monochromatic radiation3. No scatter radiation reaching thedetectorPage 28/49
Projections A single projection 𝜆𝜙 𝑥 ′ is defined as:′𝐼𝑥𝜙′𝜆𝜙 𝑥 ln 0 ′𝐼𝜙 𝑥 𝜇 𝑥, 𝑦 𝛿 𝑥 cos 𝜙 y sin 𝜙 𝑥 ′ 𝑑𝑥 𝑑𝑦 𝛿 Dirac delta function picks out 𝐴𝐵 path Reconstruction to invert equation above recover 𝜇 𝑥, 𝑦 from set of measured projections 𝜆𝜙 𝑥 ′Page 29/49
Image reconstructionRef. 2 – Chapter 3.6 Mathematics of transmission CT and theory ofimage reconstruction from projections researchfield on its own Reconstruction techniques:1. Convolution and backprojection methods also calledfiltered backprojection methods2. Iterative methods3. Cone-Beam reconstruction extract spatial (2D) distribution of linearattenuation coefficientsPage 30/49
Filtered backprojectionreconstruction algorithmsRef. 2 – Chapter 3.6 Two steps to extract 𝜇 𝑥, 𝑦 :1. Filtered / Convolution step: measured projection 𝜆𝜙 𝑥 ′ isfiltered to give a filtered projection 𝜆†𝜙 𝑥 ′ measuredprojection 𝜆𝜙 𝑥 ′ convolved with filtering operator 𝑝 𝑥 ′ :𝜆†𝜙 𝑥 ′ 𝜆𝜙 𝑥 ′ 𝑝 𝑥 ′2. Backprojection step: filtered projection 𝜆†𝜙 𝑥 ′ is backprojected distributed over the 𝑥, 𝑦 space to give 𝜇 𝑥, 𝑦 :𝜋𝜆†𝜙 𝑥 ′ 𝑑𝜙 𝑥 ′ 𝑥 cos 𝜙 𝑦 sin 𝜙𝜇 𝑥, 𝑦 0Page 31/49
Iterative reconstructionalgorithmsRef. 2 – Chapter 3.7 Developed in early days, abandoned, now back in use Basic principle:1. Computed backprojections 𝜆′ 𝜙, 𝑥 at position 𝜙, 𝑥 :𝜆′ 𝜙, 𝑥 𝑁𝛼𝑖 𝜙, 𝑥 𝜇𝑖𝑖 1𝑁 number of 2D pixels in the image𝛼𝑖 average path length of projection through 𝑖 pixel𝜇𝑖 linear attenuation coefficient density in 𝑖 pixel2. 𝛼𝑖 calculated once at start3. 𝜇𝑖 calculated iteratively until 𝜆′ closely resemble measuredbackprojections image created from 𝜇𝑖Page 32/49
Cone-Beam reconstructionalgorithmsRef. 2 – Chapter 3.8 Two main categories:1. Exact Cone-Beam reconstruction algorithmsConvert measured 1D projection data into planeintegrals use backprojection complex and requirehigh dose considered impractical for medicalapplications2. Approximate Cone-Beam reconstruction algorithmsDo not calculate full set of plane integrals simpler andrequire less dose widely usedPage 33/49
Data interpolation in helical CTRef. 2 – Chapter 3.5 Data acquired along helix and not within 2D plane one (single-slice scanner) or few (multi-slice scanner)projections available in given plane interpolation toget full set of projections for image reconstruction1. Interpolation techniques for single-slice scanners: 360o LI (Linear Interpolation): See for ex. W A Kalender et al.,Radiology 176, pg. 181-183 (1990) 180o LI (Linear Interpolation): See for ex. C R Crawford & K F King,Med. Phys. 17, pg. 967-982 (1990) Other techniques: J. Hsieh, Med. Phys. 23, pg. 221-229 (1996)2. Interpolation techniques for multi-slice scanners: See for ex. H. Hu, Med. Phys. 26, pg. 5-18 (1999)Page 34/49
CT number 𝐶𝑇 𝑛𝑢𝑚𝑏𝑒𝑟 of tissue fractional difference oftissue linear attenuation coefficient 𝜇𝑡𝑖𝑠𝑠𝑢𝑒 relativeto water 𝜇𝑤𝑎𝑡𝑒𝑟 measured in units of 0.001 Hounsfield units (HU):𝜇𝑡𝑖𝑠𝑠𝑢𝑒 𝜇𝑤𝑎𝑡𝑒𝑟𝐶𝑇 𝑛𝑢𝑚𝑏𝑒𝑟 1000𝜇𝑤𝑎𝑡𝑒𝑟 Data acquired are rescaled in terms of 𝐶𝑇 𝑛𝑢𝑚𝑏𝑒𝑟Page 35/49
2D image display Image formation steps:1. Backprojections are measured2. 𝜇𝑖 are calculated from backprojections for each 𝑖 pixel3. 𝐶𝑇 𝑛𝑢𝑚𝑏𝑒𝑟𝑠 are calculated and displayed “Display” 512 512 matrix of 2D 12 bits pixels 𝐶𝑇 𝑛𝑢𝑚𝑏𝑒𝑟 range 1000 3095 HU. Somemanufacturers offer increased range to 20,000 HU(useful for areas with metal implants) Display monitor typically 256 grey levels windowing techniques map selected range of𝐶𝑇 𝑛𝑢𝑚𝑏𝑒𝑟𝑠 (window width) onto grey scalePage 36/49
CT numbers of some tissuesTissueDensity and 𝝁𝒕𝒊𝒔𝒔𝒖𝒆𝐶𝑇 𝑛𝑢𝑚𝑏𝑒𝑟 (HU)1BoneHigh 𝜇𝑏𝑜𝑛𝑒 𝜇𝑤𝑎𝑡𝑒𝑟1000 3000BloodLow 𝜇𝑏𝑙𝑜𝑜𝑑 𝜇𝑤𝑎𝑡𝑒𝑟40MuscleLow 𝜇𝑚𝑢𝑠𝑐𝑙𝑒 𝜇𝑤𝑎𝑡𝑒𝑟10 40Brain (grey matter)Low 𝜇𝑏𝑟𝑎𝑖𝑛,𝑔.𝑚. 𝜇𝑤𝑎𝑡𝑒𝑟35 45Brain (white matter)Low 𝜇𝑏𝑟𝑎𝑖𝑛,𝑤.𝑚. 𝜇𝑤𝑎𝑡𝑒𝑟20 300WaterLipidVery low 𝜇𝑙𝑖𝑝𝑖𝑑 𝜇𝑤𝑎𝑡𝑒𝑟 50 100AirVery low 𝜇𝑎𝑖𝑟 𝜇𝑤𝑎𝑡𝑒𝑟 10001At70 keV Soft tissues low density 𝐶𝑇 𝑛𝑢𝑚𝑏𝑒𝑟𝑠 very closeto each other and to zero. Can still be resolved andreconstructed in CTPage 37/49
Signal-to-noise ratio (SNR) Sources of image noise:1. Poisson fluctuations2. Reconstruction algorithm3. Electronic noise small contribution Poisson fluctuations propagates through reconstructionalgorithm object of uniform density 𝜇 appears mottled:𝜇𝑆𝑁𝑅 Δ𝜇Δ𝜇 RMS fluctuation in 𝜇 reconstructed around mean Contrary to other imaging modalities, CT image noisenot affected by pixel sizePage 38/49
Spatial resolutionRef. 2 – Chapter 3.9.1 Spatial resolution two terms:1. In the scan plane2. Perpendicular to the scan plane Factors affecting the spatial resolution:1. Spatial resolution in the scan plane: acquisitionparameters (sampling frequency and bandwidth) andreconstruction algorithm2. Spatial resolution perpendicular to the scan plane:collimationPage 39/49
Low-contrast resolutionRef. 2 – Chapter 3.9.1 The smaller are the details with low-contrast thatcan be resolved the higher is the imaging efficacy Low-contrast resolution diameter of the smallestlow-contrast detail visible on the image Factors affecting low-contrast resolution:1. 𝑆𝑁𝑅2. Spatial resolution3. Reconstruction algorithmPage 40/49
Artefacts1. Partial-volume artefactsDue to X-ray beam divergence or anatomical structures not perpendicularto slice regions with density not corresponding to any real tissue2. Beam-hardening artefactsDue to faster absorption of low-energy X-ray beam components beamhardens false reduction in density false details ex. dark bands3. Aliasing artefactsDue to wrong sampling4. Motion artefactsDue to patient movement during scan inconsistencies in the projections “artificial” sudden changes in attenuation5. Equipment-related artefactsDue to changes in performance artefacts depend on faulty components ex. rings due to drifts in detector performancePage 41/49
Effects of reconstructionalgorithms on image qualityRef. 2 – Chapters 3.9.9, 3.9.10 and 3.9.11 Effect of spiral interpolation algorithmsSome degree of blurring of the image is introduced Effect of iterative algorithmsNoise is lower dose could be reducedNoise texture is different challenge for the radiologistas not used to it Effect of Cone-Beam reconstruction algorithms‘Wave’ or ‘windmill’ artefacts can be introducedPage 42/49
Quality control of CT scannersRef. 2 – Chapter 3.11 X-ray tube testsScan localisationCT dosimetryImage qualityHelical scanningPage 43/49
X-ray imaging dose X-ray imaging ionising radiation associated dose Dose damage:1. Deterministic effects2. Stochastic effects Damage side effects concern Dose needs to be quantified: Absorbed dose in tissue 𝐷𝑇 Equivalent dose in tissue 𝐻𝑇 Effective dose in tissue 𝐸𝑇Page 44/49
Dose quantification in CT X-ray beam divergent beam profile acrossslice not uniform CT dose index 𝐶𝑇𝐷𝐼 𝐶𝑇𝐷𝐼 measured not on patients but on dosimetryphantoms Dose delivered to patients is complex function of:1. Scanner parameters geometry, X-ray beam qualityand filtering2. Size of patient3. Acquisition parameters Empirical relation between dose on phantom andeffective dose on patientPage 45/49
CT dose index CT dose index 𝐶𝑇𝐷𝐼:𝐶𝐷𝑇𝐼1001 𝑁𝑇 50 𝑚𝑚𝐷 𝑧 𝑑𝑧 50 𝑚𝑚𝑁 number of slices𝑇 slice width𝐷 dose profile along axis of rotation 𝑧 Dosimetry phantoms used two cylindrical Perspexphantoms:1. Diameter 16 cm2. Diameter 32 cmPage 46/49
Other CT dose indexes 𝐶𝑇𝐷𝐼 depends on where on plane weighted 𝐶𝑇𝐷𝐼𝑤 :12𝐶𝑇𝐷𝐼𝑤 𝐶𝑇𝐷𝐼𝑐𝑒𝑛𝑡𝑟𝑒,100 0 𝐶𝑇𝐷𝐼100 at centre of 𝑟𝑦,100 𝐶𝑇𝐷𝐼100 1 cm under phantomsurface Average dose in volume irradiated 𝐶𝑇𝐷𝐼𝑣𝑜𝑙 :𝐶𝑇𝐷𝐼𝑤𝐶𝑇𝐷𝐼𝑣𝑜𝑙 𝑝𝑝 pitch of helical scan 𝑐𝑜𝑢𝑐ℎ 𝑖𝑛𝑐𝑟𝑒𝑚𝑒𝑛𝑡 𝑖𝑛 𝑜𝑛𝑒 ��𝑐𝑒 𝑡ℎ𝑖𝑐𝑘𝑛𝑒𝑠𝑠Page 47/49
Doses associated to imagingprocedures Approximate effective doses for common X-rayimaging proceduresBody section(Procedure)Effective dose (mSv)Planar radiographyCT scanChest0.048.3Abdominal1.57.2BrainLumbar spine1.82.4 Exact dose depends on:1. Imaging system used2. Patient’s sizePage 48/49
CT –vs– planar radiography CT disadvantagesCT much more complex than planar radiographyCT much more expensive than planar radiographyCT delivers higher dose to patients CT advantagesCT allows contrasts down to 1% to Planar radiography allows contrastsbe imaged distinguishes softonly down to 2% to be imaged tissuecannot distinguish soft tissuesCT provides 3D imagesPlanar radiography provides only2D images 3D body structurecollapsed on 2D filmPage 49/49
Webb's Physics of Medical Imaging CRC Press 3. A Del Guerra Ionizing Radiation Detectors for Medical Imaging World Scientific 4. W R Leo Techniques for Nuclear and Particle Physics Experiments Springer-Verlag . X-ray in the body Ref. 1 - Chapter 2, Ref. 2 - Chapter 2
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service i Norge och Finland drivs inom ramen för ett enskilt företag (NRK. 1 och Yleisradio), fin ns det i Sverige tre: Ett för tv (Sveriges Television , SVT ), ett för radio (Sveriges Radio , SR ) och ett för utbildnings program (Sveriges Utbildningsradio, UR, vilket till följd av sin begränsade storlek inte återfinns bland de 25 största
Hotell För hotell anges de tre klasserna A/B, C och D. Det betyder att den "normala" standarden C är acceptabel men att motiven för en högre standard är starka. Ljudklass C motsvarar de tidigare normkraven för hotell, ljudklass A/B motsvarar kraven för moderna hotell med hög standard och ljudklass D kan användas vid
LÄS NOGGRANT FÖLJANDE VILLKOR FÖR APPLE DEVELOPER PROGRAM LICENCE . Apple Developer Program License Agreement Syfte Du vill använda Apple-mjukvara (enligt definitionen nedan) för att utveckla en eller flera Applikationer (enligt definitionen nedan) för Apple-märkta produkter. . Applikationer som utvecklas för iOS-produkter, Apple .
