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CT Scan

By: Asadinezhad, MohsenPhD in Medical Physics

Ver:1-96

Contents Early History Tomography CT History CT Generations Spiral CT MultiSlice CT Digital Image CT Number Windowing

2

Contents Major Componemts

X-Ray Tube Beam Filtration Collimators Detectors

Image Reconstruction Image Quality

Spatial Resolution Contrast Resolution Noise

CT Radiation Dose Artifacts 3

References1. Euclid Seeram, Computed Tomography:

Principles, clinical applications and quality control, W.B.SANDERS Company

ني و كنترل كامپيوتري اصول فيزيكي، موارد استفاده بالي توموگرافيبهزادنيا، تانازياميرحسين قاسمي مهر و مترجمين، سيرام اوكليدكيفي،

جهانتابانتشارات 2. Thomas S Curry, James E Dowdey, Robert C

Murrey, Christensen's physics of diagnostic radiology

3. Matthias Hofer, CT teaching manual, THIEME4

Limitations of Radiography

Structures superimposed on film

Must view structure of interest through underlying / overlying structures

Patient

X-rayBeam

Film

5


Multiple views often required to adequately visualize a structure.

6


Optical density dictated by total attenuation encountered by beam

Thin highly-attenuating objects appear to be same density as thicker low-attenuating object.

Patient

X-rayBeam

Film

Thin denseobject

Thick lessdense object

7 8

Early Solution: Conventional Tomography

Sectional imaging methods first developed in 1920’s

9

Tomography

Body section radiography Planigraphy Stratigraphy Laminography Tomography Tomography (by ICRU in 1962)

10

Early History:Conventional Tomography

First used in 1935 Image produced on film Image plane oriented parallel to film Anatomy in plane of fulcrum stays in focus Anatomy outside of fulcrum plane mechanically blurred

11

Conventional Tomography BlurringConventional Tomography Blurring

Image produced on film

Objects above or below fulcrum plane change position on film & thus blur

Conventional Tomography Blurring

b c

12

13

Tomographic LayerTomographic Layer

⇒

S ≪ ,

2 2

2 2

S

U

2tt

b

a

V

θ

2tθ° 3 1320 6 2310 9 127 1 cm4 3132

14

Tube trajectories in Tomography

Linear Elliptical Circular Spiral Hypocycloidal Figure 8

Limitations of Conventional Tomography

Overlying / underlying structures blurred, not removed

5-10% subject contrast difference required for objects to appear different many anatomic systems do not have this subject

contrast

15

fracture of the base of the odontoid process with anterior displacement of the atlas

16

AP Projection Tomogram

TMJ

17

CT Advantages View anatomy without looking

through underlying / overlying structures improves contrast

Uses tightly collimated beam minimizes scattered radiation improves contrast

Demonstrates very small contrast differences reliable & repeatedly

CT X-rayBeam

ConventionalX-ray Beam

18

Film as a Radiation Detector

Analog not quantitative

Not sensitive enough to distinguish small differences in incident radiation

19

CT Detectors

electronic / quantitative extremely sensitive

small radiation input differences reliably & repeatedly measured & discerned

output digitized & sent to computer

20

21

How Did We Go From…

22

Conventional vs Axial Tomography

Conventional Cut

CT Axial Cut

23

Radiography vs. CT Imaging Limits of radiography / fluoroscopy

3D structures are collapsed into 2D image (obscuring of details, loss of one dimension)

Low soft-tissue contrast Not quantitative

Features of x-ray CT X-ray imaging modality (same principles of

generation, interaction, detection) Generation of a sliced view of body interior (“T”,

Tomography from Greek tomos = slice) Computational intensive image reconstruction (“C”)

24

CT Image Not produced on film Mathematically reconstructed from many

projection measurements of radiation intensity Digital Image calculated

Compu-ter

Digital Image

25

Basic principles

Mathematical principles of CT were first developed in 1917 by the Austrian mathematician Johann Radon (1887-1956)

Proved that an image of an unknown object could be produced if one had an infinite number of projections through the object

26

Basic principles

Idea popularized by a physicist (Allan Cormack) at Tufts Univ. (1963)

Allan MacLeod Cormack (1924–1998) shortly after the official announcement of the Nobel Prizes for medicine in 1979

27

CT HistoryCT History

First test images in 1967 First clinical images ~ 1971 First commercial scanner 1972

28

CT HistoryCT History CT made possible by high speed minicomputer

29

CT Computers Old mainframe computers too expensive & bulky

to be dedicated to CT

30

Godfrey Hounsfield, the English engineer developed the first CT scanner (1972) received the Nobel Prize in medicine in 1979 together with the physicist A.M. Cormack

31

Computerized Axial Tomography (CAT) Computerized Trans Axial Tomography (CTAT) Computerized Reconstruction Tomography (CRT) Digital Axial Tomography (DAT) Computed Tomography (CT)

32

Before Hounsfield and Cormack ....projection radiography, skull

33

Hounsfield: brain scan produced in 1974 with an 80 x 80 image matrix (a) and sagittal reconstruction generated from single scans taken with a spacing of 13 mm

Whole Body Scans

34

Topogram, Scout View, Scanogramor Pilot View.

35

CT Generations

37 38

First Generation (1970), Translation/Rotation, Pencil Beam

EMI Mark I (Hounsfield), pencil beam scanner (highly collimated source) excellent scatter rejection, now outdated

2 detectors 160 measurements during translation/ 180 - 240 rotation angle in steps of ~1 Used for the head (water bag fit tightly around head, Original computer

software couldn’t deal with transition from skull to air) 5-min scan time, 20-min reconstruction Original resolution: 80 80 pixels (ea. 3 3 mm2), 13-mm slice

X-ray Tube

Detector

39 40

Second Generation (1972), Translation/Rotation, Narrow Fan Beam

Narrow Fan beam (10˚) Linear detector array (~5-30 detectors) 180 ˚ rotation angle in steps of 5-10˚ Reduced number of view angles scan time ~20-30 s Slightly more complicated reconstruction algorithms because

of fan-beam projection

41

3rd Generation Geometry

Patient

42

Third Generation (1976), Rotation/Rotation, Wide Fan Beam

Wide fan beam (30-60˚) covers entire object 30-900 detectors (ionization chamber or scintillation detector) No translation required scan time ~seconds (reduced dose,

motion artifacts) Reconstruction time ~seconds Pulsed source (reduces heat load, radiation dose) Ring Artifact

43 44

4th Generation CT4th Generation CT

45

Fourth Generation (1978), Rotation/Stationary, Wide Fan Beam

Wide fan beam (30-60˚) covers entire object Stationary detector ring (600 – 4800 scintillation detectors) Rotating x-ray tube (inside or outside detector ring) Scan time, reconstruction time ~1 second Source either inside detector ring or outside (rocking, nutating

detectors)

46

Comparison of 3rd and 4th Generation

Both designs currently employed, neither can be considered superior

3rd Generation (GE, Siemens):Fewer detectors (better match, cheaper)Good scatter rejection with focused septa

4th Generation (Picker, Toshiba):Less moving partsDetectors calibrated twice per rotation

47

3rd & 4th Generation (Non-spiral) CT

Tube rotates once around patient Table stationary data for one slice collected

Table increments one slice thickness Repeat

Tube rotates opposite direction

48

Spiral CT

Patient

49

Spiral CTSpiral CT Continuous rotation of gantry & linear motion of patient

table Patient moves slowly through gantry Cables of old scanners allowed only 360o rotation (or

just a little more) Tube had to stop and reverse direction No imaging done during this time

No delay between slices Dynamic studies now limited only by tube heating

considerations Increased coverage volume / rotation

50

Slip rings - spiral CT

51

Spiral CT

table increment during one 360° rotation Pitch factor = -------------------------------------------------

slice thickness

52

Pitch factor = 1

table motion during one 360° rotation Slice Pitch = ---------------------------------------------

slice thickness

Pitch factor = 1 means slices touch each other

53

Pitch factor >1


slice thickness

Pitch factor > 1 means gap in slices

54

Pitch factor <1


slice thickness

Pitch factor < 1 means overlap in slices Can improve visualization of objects

55

Pitch factor = 1

equivalent dose to non-spiral

56

Pitch factor >1

lower dose for spiral if table increment per rotation > one slice thickness

57

Pitch factor <1

higher dose for spiral if table increment per rotation < one slice thickness

58

59

5th Generation: Scanners for CV Imaging – “Imatron”

No moving parts Electromagnetically swept electron beam 50 ms scan time imaging of beating heart Developed 1983 Multi slice capability

60

Cine CT (Imatron)Cine CT (Imatron) four tungsten target rings that makes a 210° arc around the patient

replaces conventional x-ray tube electron beam sweeps over each annular target ring

can be done at electronic speeds 2 detector rings with arcs of 216°

One arc with 432 detector, another with 864 detector (higher resoution) Cadmium tungstate crystal (CdWO4)

maximum scan rate 24 frames per second

Electron-beam CT, also known as fifth-generation CT

Wolbarst A B , Hendee W R Radiology 2006;238:16-39

61 62

Multi-slice CT or MultiDetector CT (MDCT) 1991

Multi-slice CT or MultiDetector CT (MDCT) 1991

Multiple rows of fan beam detectors Wider fan beam in axial direction Table moves much faster Substantially greater throughput

63

Multi-slice CTMulti-slice CT Multi-slice CT

64

Multiple detector arrays

Set of several linear detector arrays, tightly abutted

Use solid-state detector arrays Slice width is determined by the detectors, not

by the collimator (although collimator does limit the beam to the total slice thickness)

65 66

Multiple detector arrays (cont.)

3rd generation multiple detector array with 16 detectors in the slice thickness dimension and 750 detectors along each array uses 12,000 individual detector elements

4th generation scanner would require roughly 6 times as many detector elements; consequently currently planned systems use 3rd generation geometry

67

Slice thickness:single detector array scanners

Determined by the physical collimation of the incident x-ray beam with two lead jaws

Width of the detectors places an upper limit on slice thickness

For scans performed at the same kV and mAs, the number of detected x-ray photons increases linearly with slice thickness

Larger slice thicknesses yield better contrast resolution (higher SNR), but the spatial resolution in the slice thickness dimension is reduced

68

Slice thickness:multiple detector array scanners

In axial scanning (i.e., with no table movement) where, for example, four detector arrays are used, the width of the two center detector arrays almost completely dictates the thickness of the slices

For the two slices at the edges of the scan, the inner side of the slice is determined by the edge of the detector, but the outer edge is determined either by the outer edge of the detector or by the collimator penumbra, depending on collimator adjustment

69 70

Slice thickness: MDA (cont.)

In helical mode, each detector array contributes to every reconstructed image Slice sensitivity profile for each detector array needs to be

similar to reduce artifacts Typical to adjust the collimation so that the focal spot –

collimator blade penumbra falls outside the edge detectors Causes radiation dose to be a bit higher (especially for small slice

widths) Reduces artifacts by equalizing the slice sensitivity profiles

between the detector arrays

71

Detector pitch/collimator pitch

Pitch is a parameter that comes into play when helical scan protocols are used

In a helical scanner with one detector array, the pitch is determined by the collimator

Collimator pitch = table movement (mm) per 360-degree rotation of gantry / collimator width (mm) at isocenter

Pitch may range from 0.75 (overscanning) to 1.5 (faster scan time, possibly smaller volume of contrast agent)

72

Pitch (cont.)

For scanners with multiple detector arrays, collimator pitch is still valid

Detector pitch = table movement (mm) per 360-degree rotation of gantry / detector width (mm)

For a multiple detector array scanner with N detector arrays, collimator pitch = detector pitch / N

For scanners with four detector arrays, detector pitches running from 3 to 6 are used

73

Multi-detector planes

74

Multi-detector planesNew Technology

GE QXi (multi-detector CT) acquires four interweaving helices simultaneously.e.g., 4 x 5 mm slice = 20 mm total scan width

4-slice in one rotation

76

Definitions of Pitch

Old definition: Table travel per rotation

P = slice thickness

New definition:

Table travel per rotationP’=

Total nominal scan width

77

GE QXi High Quality (HQ) vs High Speed (HS)

Pitch = 15mm/20 mm =0.75

Pitch = 30mm/20 mm =1.5

20 mm

15 mm table travel

30 mm table travel

20 mm

78

79

Typical characteristics of CT

1972 1980 1990 2000

Minimum scan time 300 s 5-10 s 1-2 s 0.3-1s

Data acquired per 360° 57.6 kB 1 MB 2MB 42 MB

Data per spiral sequence - - 24-48 MB 200-500 MB

Image matrix 802 2562 5122 5122

Power (generator) 2 kW 10 kW 40 kW 60 kW

Slice thickness 13 mm 2-10 mm 1-10 mm 0.5-5 mm

Toshiba Aquilion ONE CT

80

320-slice (320 x 0.5 mm), 16cm gantry rotation, Year product introduced: 2007, 7.5 MHU, more than a $1 million

Toshiba Aquilion ONE Vision Edition

81

640-slice. 0.275 sec rotation, 16cm gantry rotation, Year product introduced: 2012

Micro CT A miniaturized design The X-rayed measuring field, usually as small as 2cm3

for material testing and analysis, medical applications are on their way to taking center stage (analysis of trabecular structures in bones)

82

Dual Energy CT Single Source or Dual Source

83

Dual Energy CT

84

Dual Source CT

85

Dual Source CT

86

Dual Source CTDetector 2 x Stellar detectorNumber of slices 2 x 128Rotation time 0.28 s‐1

Temporal resolution 75 ms-1, heart-rate independentGenerator power 200 kW (2 x 100 kW)kV steps 70, 80, 100, 120, 140 kVIsotropic resolution 0.33 mmCross-plane resolution 0.30 mmMax. scan speed 458 mm/s1 with Flash SpiralTable load up to 307 kgGantry opening 78 cm

87

SPECT-CT

88

SPECT-CT

89

PET-CT

90

PET-CT

4D PET-CT Image

91

PE

T-C

T

92

93

Measure Intensity of a Pencil Beam

X-Ray Source

Radiation Detector

94

Principle of X-Ray CT

In one plane, obtain set of line integrals for multiple view angles

Reconstruct cross-sectional views

Detector

Linear scan

Angular scan

Object

95

Pixels & Voxels

96

Digital ImageDigital Image

2-dimensional array of individual image points calculated

each point called a pixel picture element

each pixel has a value value represents x-ray

transmission (attenuation)

97

Pixels & Voxels

Pixel is 2D component of an image

Voxel is 3D cube of anatomyVolume Element

CT reconstruction calculates attenuation coefficients of Voxels

CT displays CT numbers of Pixels as gray shades

98

Pixel & Voxel Size

Voxel depth same as slice thickness

Pixel dimension field of view / matrix size

FOV = 30 cm256 pixels 30 cmPixel size = ------------

256 pixels

Pixel size = 0.117 cm = 1.17 mm

99

Attenuation Equation forMono-energetic Photon Beams

I = Ioe-x

I = Exiting beam intensityIo = Incident beam intensitye = constant (2.718…) = linear attenuation coefficient

•property of•absorber material•beam energy

x = absorber thickness

MaterialIo

I

x

For photons which are neither absorbed nor scattered

100

Example Beam Attenuation

Using equation to calculate beam intensity for various absorber thicknesses ( = .223)

1cm100 80

I = Ioe-x

100*e-(0.223)(1) = 80-20%

101



1cm 1cm100 80 64

I = Ioe-x

100*e-(0.223)(2) = 64

-20% -20%

102



1cm 1cm 1cm100 80 64 51

I = Ioe-x

100*e-(0.223)(3) = 51

-20% -20% -20%

103



1cm 1cm 1cm 1cm100 80 64 51 41

I = Ioe-x

100*e-(0.223)(4) = 41

-20% -20% -20% -20%

104

More Realistic CT Example Beam Attenuation for non-uniform Material 4 different materials 4 different attenuation coefficients

#1 #2 #3 #4

1 2 4

Io I

x

I = Ioe-(+++)x

105 106

Reconstruction:Solve for ’s

16 22 11 1017

22

12

10

15

13

11 12 13 14

21 22 23 24

31 32 33 34

41 42 43 44

107

Real Problem Slightly More Complex

11 12 13 14

21 22 23 24

31 32 33 34

41 42 43 44

24 13 15 22 16

35

13

22

9

14512 values

512values

108

Effect of Beam Energy on Attenuation

Low energy photons more easily absorbed High energy photons more penetrating All materials attenuate a larger fraction of low

than high energy photons

Material100 80

Higher-energymono-energeticbeam

30Material

Lower-energymono-energeticbeam

100

109

Mono vs. Poly-energetic X-ray Beam Equations below assume Mono-energetic x-

ray beam

#1 #2 #3 #4

1 2 4

Io I

x

I = Ioe-(+++)xI = Ioe-x110

Mono-energetic X-ray Beams

Available from radionuclide sources Not used in CT because beam intensity much

lower than that of an x-ray tube

111

X-ray Tube Beam High intensity Produces poly-energetic beam

#1 #2 #3 #4

1 2 4

Io I

x

I = Ioe-(+++)xMono-energetic beam equation!

112

Beam Hardening Complication Attenuation coefficients n depend on beam energy!!! Beam energy incident on each block unknown Four ’s, each for a different & unknown energy

1 2 4

1cm 1cm 1cm 1cm

I = Ioe-(+++)x

113

Beam Hardening Complication

Beam quality changes as it travels through absorber greater fraction of low-energy photons removed from

beam Average beam energy increases

1cm 1cm 1cm 1cm

Fewer PhotonsBut higher avg

kV than A


kV than B

A B


kV than C

C D


kV than D

E

114

Reconstruction

Scanner measures “I” for thousands of pencil beam projections

Computer calculates tens of thousands of attenuation coefficients one for each pixel

Computer must correct for beam hardening effect of increase in average beam energy from one side of

projection to other

I = Ioe-(++++)x

115

CT Number (The Hounsfield Unit)

Calculated from reconstructed pixel attenuation coefficient

t - W)HU= CT # = 1000 ------------

W

Where:t = linear attenuation coefficient for tissue in pixelW = linear attenuation coefficient for water

Caculate CT # for Water. Answer: 0Caculate CT # for Air. Answer: -1000

116

CT Numbers for Special Stuff

Bone: +1000 Water: 0 Air: -1000

t - W)CT # = 1000 ------------

W

117

The Hounsfield scale

118

Digital Image MatrixDigital Image Matrix

125 25 311 111 182 222 176

199 192 85 69 133 149 112

77 103 118 139 154 125 120

145 301 256 223 287 256 225

178 322 325 299 353 333 300

119

Numbers / Gray ShadesNumbers / Gray Shades

Each number of a digital image corresponds to a gray shade for one pixel

120

Digital to Analog Conversion(D to A)

Computer reconstructs digital image set of numbers

Computer displays analog image

125 25 311 111 182 222 176

199 192 85 69 133 149 112

77 103 118 139 154 125 120

145 301 256 223 287 256 225

178 322 325 299 353 333 300

121

Analog vs. Digital Images

Analog continuous gray

shade information Digital

Discrete gray shade information

122

Digital Image FormationDigital Image Formation

Clinical ImageScreen Wire Mesh

123

Digital Image Formation:Sampling


Place mesh over image

Assign each square (pixel) a value based on density

Pixel values form the digital image

120

-10

-650

124



Each pixel assigned a value

Value averages entire pixelAny spatial variation

within a pixel is lostThe larger the pixel,

the more variation120

-10

-650

125

Digital Image FormationDigital Image Formation The finer the mesh (sampling), the more accurate the

digital rendering

126

What is this?What is this?

12 X 9 Matrix

127

Same object, smaller squaresSame object, smaller squares

24 X 18 Matrix128


48 X 36 Matrix

129


96 X 72 Matrix 130


192 X 144 Matrix

131

Image Reconstruction

AcmeMini-

Computer

Projection(raw)Data

Pixel(calculated)

Data

X-Ray Source

Radiation Detector

132

Data AcquisitionData Acquisition

cross sectional image reconstructed from many straight line transmission measurements made in different directions

Tube

Detector

133

Projection MeasurementsProjection Measurements

Radiation detector generates a voltage proportional to radiation intensity

134

Image Reconstruction Minicomputer does its thing

Analog to Digital (A to D) conversion

135

Digital Image MatrixDigital Image Matrix

125 25 311 111 182 222 176

199 192 85 69 133 149 112

77 103 118 139 154 125 120

145 301 256 223 287 256 225

178 322 325 299 353 333 300

Digital Matrix contains many numbers which may be Displayed on CRT Manipulated Stored

136


One of these equations for every projection line

IA = Ioe-(++++)xProjection #A

IC = Ioe-(C+C+C+C+)xProjection #C

Projection #B

IB = Ioe-(++++)x

137


IA = Ioe-(++++)x

IB = Ioe-(++++)x

IC = Ioe-(C+C+C+C+)x

Projection #A

Projection #B

Projection #C

IA, IB, IC, ...What We Measure:

A1, A2, A3, ...

Reconstruction Calculates:

B1, B2, B3, ...C1, C2, C3, ...

Etc. 138

Display & WindowingDisplay & Windowing

Gray shade assigned to each pixel value (CT #)

Windowing Assignment of display brightness to pixel

values does not disturb original pixel values in

memory Operator controllable

Window Width Window Level

47

93

139

Display & Display Matrix:Resolution

CT images usually 512 X 512 pixels Display resolution better

often 1024 X 1024 can be as high as 2048 X 2048

$$$

140

Display & Display Matrix:Contrast

CT #range -1000 to 3000

Monitor can display far fewer gray shades Eye can discern few gray shades Purpose of Window & Leveling

display only portion of CT # values Emphasize only those CT #’s display of CT #’s above & below window all black OR all white

141

Pixel Values & Gray Shades

# of valid pixel values depends on bit depth 1 bit: 2 values 2 bits: 4 values 3 bits: 8 values 8 bits: 256 values 10 bits: 1024 values n bits: 2n values

142

Pixel Values & Gray Shades

CT can discern ~ 4000 gray shades Typical bit depth: 10 bits = 1024 gray shades Single gray shade represents range of pixel

values

143

Window Width & Level

Window width range of CT #’s imaged determines maximum # of gray

shades which could be displayed on CRT

Window level center or midpoint of CT # range

144


Pixels outside of window displayed as Black or White

145


>200

151-200

101-150

51-100

1-50

(-49)-0

(-99)-(-50)

(-149)-(-100)

(-199)-(-150)

<(-199)

3000

0

1000

2000

Window: 400Level: 0

-1000146

Small Window Width200

-200

0

1000

-1000

Window: 400Level: 0

� Short gray scale� Small block of CT #’s

assigned gray levels� Small transition zone

of white to black

147

Small Window Width200

-200

0

1000

-1000

Window: 400Level: 0

� Used to display soft tissues within structures containing different tissues of similar densities

� Level centered near average CT # of organ of interest

148

Large Window Width

0

1000

Window: 2000

Level: 0

-1000

� Long gray scale� Large block of CT #’s

assigned gray levels� Large transition zone of

white to black

149

Large Window Width

0

1000

Window: 2000

Level: 0

-1000

� Used where large latitude required

� Used to simultaneously display tissues of greatly differing attenuation

150

Window Example

WL =0WW = 200

All pixels with CT #’s > 0 +(200/2) = 100: White

All pixels with CT #’s < 0 -(200/2) = -100: Black

100

-100

200 0

151

Another Window Example

WL = 40WW = 200

All pixels with CT #’s > 40 + (200/2) = 140: White

All pixels with CT #’s < 40 - (200/2) = -60: Black

140

-60

200 40

152

Still Another Window Example

WL = 0WW = 400

All pixels with CT #’s > 0 + (400/2) = 200: White

All pixels with CT #’s < 0 - (400/2) = -200: Black

200

-200

400 0

153

Larger Window Means Obscuring Small Differences in Tissue Attenuation

One gray shade encompasses larger range of CT #’s

200

-200-100

100

20 - 40 40 - 8020 40

WW=200 WW=400

Range

154

Windowing procedures to display CT images. The diagnostically relevant range of CT values is selected by choosing the center and width (C/W) of the window.

155

Window Width & Contrast

As WW increases contrast decreases latitude (range of CT #’s

imaged) increases As WW decreases

contrast increases latitude decreases

Clinical goal: Largest available contrast at

the latitude required by study

156

Window Width & Image Contrast

Large window width different structures more likely to have same

gray shade Narrow window width

Gray shade differences more likely visible between structures

Very narrow window width Small differences in attenuation seen as black

& white

157

Preset Window & Level

Available for all commercial CT initial WW and WL pre-sets for specific study

types Can be overridden by operator

158

Silly CT # Display Example:10 Gray Shades

>700

651-700

601-650

551-600

501-550

451-500

401-450

351-400

301-350

<301

159

CT # Level Change

Darks lighterlights lighter

Decreased LevelConstant Window

160

CT # Level Change

>700

651-700

601-650

551-600

501-550

451-500

401-450

351-400

301-350

<301

>200

151-200

101-150

51-100

1-50

(-49)-0

(-99)-(-50)

(-149)-(-100)

(-199)-(-150)

<(-199)

161

CT # Level Change

>700

651-700

601-650

551-600

501-550

451-500

401-450

351-400

301-350

<301

>200

151-200

101-150

51-100

1-50

(-49)-0

(-99)-(-50)

(-149)-(-100)

(-199)-(-150)

<(-199)

3000

-1000

0

1000

2000

Window: 400Level: 500

Window: 400Level: 0

162

CT # Window Change

Darks lighter,lights darker

163

CT # Window Change

>700

651-700

601-650

551-600

501-550

451-500

401-450

351-400

301-350

<301

>900

801-900

701-800

601-700

501-600

401-500

301-400

201-300

101-200

<101164

CT # Window Change

>700

651-700

601-650

551-600

501-550

451-500

401-450

351-400

301-350

<301

3000

-1000

0

1000

2000

>900

801-900

701-800

601-700

501-600

401-500

301-400

201-300

101-200

<101



165 166

Major Components

Scanner Room Imaging system Generator (?)

Electronics Room Power Computer (?) Generator (?)

Operator’s Area Display / recording / storage Computer (?)

167

Major Components

X-Ray Production

X-Ray Detection

Computer Systems

Reconstruction

X-Ray Tube

Detectors

A - D Conversion

Display & Format

Printing

Archiving

Generator

168

Major Components

The Gantry X Ray Tube, Detectors, H.V Generator

The Operating Console Operator Console

kV, mA, Slice Thickness

Physician Console The Computer

169

X-Ray System Components

X-Ray Generator X-Ray Tube Beam Filter Collimators

170

X-Ray Generator

3 phase originally used Most vendors now use high frequency generators

relatively small small enough to rotate with x-ray tube can fit inside gantry

lower ripple than 3 phase more efficient production of x-rays

171

X-Ray Tube

Must provide sufficient intensity of transmitted radiation to detectors

Radiation incident on detector depends upon beam intensity from tube patient attenuation

beam’s energy spectrum patient

thickness atomic # Density

500,000 to 2,000,000 HU172

173

Patient

CT Beam Filtration

Hardens beam preferentially removes low-

energy radiation Removes greater fraction of low-

energy photons than high energy photons

reduces patient exposure Attempts to produce uniform

intensity & beam hardening across beam cross section

Filter

CT Beam Filtration

174

175

Filter

Collimators

Source Detector

Pre-Collimator Post-Collimator

Patient

Scattering

176

Tube, collimator & detector

177

Pre-Collimation

Constrains size of beam Reduces amount of scatter produced Designed to minimize beam divergence Often consists of several stages or sets of jaws

Tube

Detector

Pre-collimator

178

Post-Collimation

Helps define slice (beam) thickness Reduces scatter radiation reaching detector

Tube

Detector

Post-collimator

179

CT Detector Technology:Desirable Characteristics

High efficiency Quick response time High dynamic range Stability

180

CT Detector EfficiencyDefinition

Ability to absorb & convert x-ray photons to electrical signals

181

Efficiency Components

Capture efficiency fraction of beam incident on active detector

Absorption efficiency fraction of photons incident on the detector which are

absorbed Conversion efficiency

fraction of absorbed energy which produce signal

182

Overall Detector Efficiency

Overall detector efficiency =

capture effi. × absorption effi. × conversion effi.

183

Absorption Efficiency

Depends upon detector’s atomic # density size thickness

Depends on beam spectrum

capture efficiency×

absorption efficiency×

conversion efficiency

184

Response Time

Minimum time after detection of 1st event when detector can detect 2nd event

If time between events shorter than response time, second event may not be detected

Shorter response time better

185

Stability

“Steadiness” of detector system

Consistency of detector signal over time

The less stable, the more frequently calibration required

186

Dynamic Range

Ability to faithfully detect large range of intensities

Ratio of largest to smallest signal which can be faithfully detected

Typical dynamic range: 1,000,000:1 much better than film

Detector Types: Gas Ionization Measurement of conductivity induced in a gas volume by the

ionizing effect of x-rays. X-rays ionize gas molecules Ions are drawn to electrodes by electric field

Number of ion pairs N produced x-ray intensity

187

++ + +-- - -

+ -

Ammeter

Anode

Cathode

Collimator

Gas Ionization Chambers CharacteristicsTo optimize efficiency usually filled with Xenon (high Z) under pressure (up to 30 atm) Cheap Excellent stability Large dynamic range High spatial resolution Low efficiency

188

189

Detector Types: Scintillation

X-ray energy converted to light Light converted to electrical signal

X-Rays

Photomultiplier Tube

Light

ElectricalSignal

ScintillationCrystal

Scintillation materials: NaI(Tl), BGO PMTCdWO4, CsI, Rare Earth Oxides (Gd2O2S) PD

•Scintillation material thick enough to provide quantum efficiency ~ 100% 190

Photomultiplier Tubes

Light incident on Photocathode of PM tube Photocathode releases electrons

X-Rays Light

ScintillationCrystal PM

TubePhotocathode

-+

Dynodes

191

Photomultiplier Tubes

Electrons attracted to series of dynodes each dynode slightly more positive than last one

X-Rays Light

ScintillationCrystal PM

TubePhotocathode

-+

+

+

+

+

Dynodes

192

Photodiode

Made of two types of materials p-type n-type

Lens focuses light from crystal onto junction of p & n type materials

pn

Lens JunctionX-Rays Light

193

Photodiode

Light controls resistance of junction Semiconductor current proportional to light

falling on junction

pn

LensJunctionX-Rays Light

194

Solid State Detectors Output electrical signal amplified Fast response time Large dynamic range Almost 100% conversion & photon capture

efficiency

Factors Affecting Detector Signal

195

kV: high kV x-rays more penetrating mA: high tube current gives more intense x-

ray beam Scan time: long scan time more x-rays to

detectors Slice thickness: wide slice more x-rays Patient composition: small patients less

attenuating


196

“It All Adds Up” Puzzle

2 5

6 4 0

9

2 1

9 23 15 17 14

22

16

19

7

17

8

7

31

0 3

6

7

197

“It All Adds Up” Puzzle

5 1

8

6 7

5 0 2

9 26 14 17 18

13

21

18

14

23

198

This is what your CT Scanner must solve!

16 22 11 10 17

22

12

10

15

13

199

Reconstruction:Solve for ’s

16 22 11 10 17

22

12

10

15

13

11 12 13 14

21 22 23 24

31 32 33 34

41 42 43 44

200

Real Problem Slightly More Complex

11 12 13 14

21 22 23 24

31 32 33 34

41 42 43 44

24 13 15 22 16

35

13

22

9

14512 values

512values

201

Real Reconstruction Problem

� Intensity (transmission) measured

� Rays transmitted through multiple pixels

� Find individual pixel values from transmission data (question marks)

? ? ? ?? ?

? ? ? ?? ?

? ? ? ?? ?

? ? ? ?? ?

? ? ? ?? ?

? ? ? ?? ? 534

417

364

555

501

355

255 712199

202

Raw Data� unanalyzed data;

data not yet subjected to analysis

� Intensity (transmission) measurements for each ray for each projection

534

417

364

501

255 712199

203

Image Data� Individual pixel

values (question marks)

? ? ? ?? ?

? ? ? ?? ?

? ? ? ?? ?

? ? ? ?? ?

? ? ? ?? ?

? ? ? ?? ?

204

Algorithm� Set of rules for getting a specific output (answer)

from a specific input� Basic Reconstruction algorithm methods

1. Back Projection پس نمايش 2. Iterative method روش تكرار شونده 3. Analytical Method روشهاي تحليلي

205

� ALSO CALLED SUMMATION METHOD OR LINEAR SUPERPOSITION METHOD

� Reverse the process of measurement of projection data to reconstruct an image

Back Projection ReconstructionBack Projection Reconstruction

206


� Back Projection• for given projection,

assume equal attenuation for each pixel

• repeat for each projection adding results

9999999 9

207


� Reconstruction Problem• converting transmission data for

individual projections into attenuation data for each pixel

??????? 9

208


� Assume actual image has 1 hot spot (attenuator)

� Each ray passing through spot will have attenuation back-projected along entire line

� Each ray missing spot will have 0’s back-projected along entire line

Hot Spot

9999999 9

0000000 0209


� Each ray missing spot stays blank� Each ray through spot shares some density

• Location of spot appears brightest

Hot Spot

9999999 9

0000000 0

210


HotSpot

Star Artifact Spokes

� Streaks appears radially from spot• star artifact

211


212


:معايب اين روش�.تصوير هر نقطه به شكل ستاره در مي آيد1..دقت تصوير بسيار پائين است2..تصوير مقدار زيادي نويز زمينه يا مه آلودگي دارد3..كنتراست تصوير پائين است4.

213

Iterative methods روشهاي تكرار شونده Based on a series of estimates and corrections for errorsITERATIVE TECHNIQUES ARE NOT USED IN TODAY’S

COMMERCIAL SCANNERS1. Simultaneous Correction تصحيح همزماني

or Iterative Least Square Technique (ILST)روش بازسازي تصوير با حداقل كردن خطاهاي مربعها

2. Ray by Ray Correction تصحيح اشعه به اشعهor Algebra Reconstruction Technique (ART)روش بازسازي جبري

3. Point by Point Correction تصحيح نقطه به نقطهor Simultaneous Iterative Reconstruction Technique (SIRT)روش بازسازي تصحيح همزمان

214

Iterative ReconstructionIterative Reconstruction� Start with measured data

? ? ?

? ? ?

? ? ?

24 12 12

Measurements

17

19

12

12

159

215

Iterative ReconstructionIterative Reconstruction� Make initial guess for first projections

by assuming equal attenuation for each pixel in a projection

� Similar to back projection

8 4 4

8 4 4

8 4 4

24 12 12

Initial guess basedupon vertical projections

Measurements

? ? ?

? ? ?

? ? ?

24 12 12

Measurements

17

19

12

12

159

216

Iterative ReconstructionIterative Reconstruction

217

Iterative ReconstructionIterative Reconstruction

� changing pixels for one projection alters previously-calculated attenuation for others

� corrections repeated for all projections until no significant change / improvement

218

Iteration Image ReconstructionIteration Image Reconstruction

� operationally slow and cumbersome, even for computers

� not used

219

:معايب اين روش�ه زمان نسبتاً طوالني جهت اجراي هر تكرار و در نتيج1.

بازسازي تصويررائب احتمال تفاوت ضرائب تضعيف محاسبه شده با ض2.

حقيقي

Iteration Image ReconstructionIteration Image Reconstruction

220

Analytical Methods روشهاي تحليليUSED IN MODERN CT SCANNERS Developed to overcome

limitations of back-projection and iterative algorithms1. Filtered Back Projection پس نمايش فيلتر شده2. Fourier Analysis آناليز فوريه

221

Filtered Back ProjectionFiltered Back Projection� enhancement of back projection

technique� filtering function (convolution)

is imposed on transmission data• small negative side lobes placed

on each side of actual positive data

• negative values tend to cancel star artifact

Filtered back

projection

Unfiltered back

projection

*

222

Filtered Back ProjectionFiltered Back ProjectionSTEPS IN FILTERED BACK PROJECTION:� ALL PROJECTION PROFILES ARE OBTAINED � THE LOGARITHM OF DATA IS OBTAINED � LOGARITHMIC VALUES ARE MULTIPLIED BY DIGITAL

FILTER � FILTERED PROFILES ARE BACKPROJECTED � THE FILTERED PROJECTIONS ARE SUMMED AND THE

NEGATIVE AND POSITIVE COMPONENTS ARE CANCELLED

* Filtered Back ProjectionFiltered Back Projection� operationally fast

• reconstruction begins upon reception of first transmission data

� best filter functions found by trial & error

� Most common commercial reconstruction algorithm

224

Filtered Back ProjectionFiltered Back Projection

225

Filtered Back ProjectionFiltered Back Projection

226

Fourier Analysis� USED IN MRI, NOT USED IN CT BECAUSE OF COMPLICATED

MATHEMATICS

� converts data from spatial domain to frequency domain• breaks any signal into frequency component

parts

227

Multi-plane reconstructionMulti-plane reconstruction� Or Multiplanar Reformatting� using data from multiple axial

slices it is possible to obtain• sagittal & coronal planes• oblique & 3D reconstruction

� Non-spiral reconstruction• Poor appearance if slice thickness

>>pixel size • isotropic imaging

� multi-plane reconstructions are computer intensive (Can be slow) 228

Saggital / Coronal Reconstructions

Saggital

Coronal

Axial

229

3D Reconstructions

230

3D Reconstructions� Uses pixel data from multiple slices� Algorithm identifies surfaces & volumes� Display renders surfaces & volumes

• Three Dimensional Shaded Surface• Volume Rendering• Maximum Intensity Projection (MIP)

� Real-time motion» auto-rotation» user-controlled multi-plane rotation

231

Interpolation� Calculating attenuation data for specific slice

from spiral raw data� Table moves continually� As tube rotates table constantly moves

Position at start of rotation

Position at start of rotation

Position of interest 232

Interpolation

� Estimates value of function using known values on either side

When x = 50, y = 311When x = 80, y = 500

What will be the value of y when x=58? (50,311)

(80,500)?

58233

Interpolation� 58 is 8/30ths of the way between points� “y” when x=58 will be 8/30ths of the way

between 311 and 500

(50,311)

(80,500)?

58

?=311+8/30 (500-311)

234

Three Dimensional Shaded Surface1. User selects a threshold range2. The voxels with CT# within the

threshold range are set to the "on" state, The rest of the voxels are set to the "off“ state

3. Project rays through the entire volume4. As the rays pass through the data, they

stop when they identify the first "on" voxel (part of the surface); the other voxels are ignored.

5. This is done for all the rays� No details below the surface 235

Volume Rendering� Displays an entire volume set with control of the opacity or

translucency of selected tissue types� Each voxel has an associated intensity in addition to an

associated opacity value

236

Maximum Intensity Projection (MIP)� easy viewing of vascular structures or air-filled cavities� The rays are cast throughout the volume, and depending on

whether it is maximum intensity projection or minimum intensity projection, maximum or minimum values along the rays are used in the final image display.

� preferred method for many CT angiography applications because visualization of contrast-filled vasculature is fast and easy.

237 238

CT Image Quality Parameters

SpatialResolution

ImageNoise

ContrastResolution

Artifacts

CTاطالعات آناتوميكي مهم در روي تصوير ) تشخيص(قابل رويت بودن

239

Factors InfluencingCT Image Quality

BeamCharacteristics

Patient Dose

SliceThickness

Scatter

DisplayResolution

ReconstructionAlgorithm

SubjectTransmissivity

QUALITY MEASUREMENT METHODS

� PSF- POINT SPREAD FUNCTION � LSF- LINE SPREAD FUNCTION � CTF – CONTRAST SPREAD FUNCTION � MTF- MODULATION TRANSFER

FUNCTION � ERF- Edge Response Function

240

PSF- POINT SPREAD FUNCTION

� Describes the lack of sharpness that results when a point in the object is not reproduced as a “true” point in the image. As a result this causes blurring

241

Full width at half-maximum (FWHM)

LSF- LINE SPREAD FUNCTION

� Also describes the unsharpness of an imaging system when a line or slit object is not reproduced as a line or slit but instead spreads out as a measurable distance.

242

CTF – CONTRAST SPREAD FUNCTION

� Also Known As the contrast response function measures the contrast response of an imaging system. When utilizing a phantom/test pattern it is the resultant difference in density between the adjacent regions of the slits.

243

MTF MODULATION TRANSFER FUNCTION

� The MTF is a combination of LSF, PSF and ERF. It is obtained with the Fourier transform of the LSF, PSF, and ERF.

� MOST COMMONLY USED TO DESCRIBE SPATIAL RESOLUTION IN CT.

244

MTF

� MTF OF 1 – 100% TRANSFER OF OBJECT TO IMAGE

� MTF OF 0 – 0% TRANSFER OF OBJECT TO IMAGE

Spatial frequency

Lp/cm

Spatial Frequency

� LARGE OBJECTS – LOW S.F.

� SMALL OBJECTS – HIGH S.F.

248

SPATIAL FREQUENCY� Smallest resolvable high contrast object� Often expressed as line pairs / cm� “Pair” is one object + one space

OnePair

SPATIAL FREQUENCY

1

3

250

SPATIAL FREQUENCY� Smallest resolvable high contrast object is

half the reciprocal of spatial frequency� Example:

• Limited resolution = 15 line pairs per cm• Pair is 1/15th cm• Object is half of pair

» 1/15th / 2» 1/30th cm» .033 cm» 0.33 mm

1/15th cm

1/30th cm

Resolution in CT

Spatial Resolution Contrast Resolution

252

Spatial Resolution� Quantifies image blurring� “Ability to discriminate objects of varying density a

small distance apart against a uniform background”� Minimum separation required between two high

contrast objects for them to be resolved as twoobjects

254

Factors affectingSpatial Resolution (↑)

� Focal spot size (↓)� Detector aperture width (↓)� FOV (↓) Pixel Size (↓)� Slice Thickness or collimation (↓)

• Less variation likely for thinner slices• attenuation variations within a voxel are averaged

» partial volume effect

� Voxel Size (↓): Pixel Size & Slice Thickness� Patient Movement (↓)� Reconstruction algorithms (Bone filters have the best spatial

resolution, and soft tissue filters have lower spatial resolution)• smoothing or enhancing of edges

� Greater pitches reduce resolution

� CT spatial resolution phantom, consisting of 4–12 line-pairs per centimeter, reconstructed using standard (A) and bone (B, high-resolution) filters.

255 256

Hi-Resolution CT (HRCT) Technique� Very small slice thicknesses

• 1-2 mm� High spatial frequency algorithms

• increases resolution» increases noise» Noise can be offset by using higher doses

� Optimized window / level settings� Small field of view (FOV)

• Known as “targeting”

257

Contrast Resolution

� Ability of an imaging system to demonstrate small changes in tissue contrast

� The difference in contrast necessary to resolve 2 large areas in image as separate structures

258

Contrast ResolutionContrast Resolution

� Difference in x-ray attenuation required for 2 pixels to be assigned different digital values

89

88

Image Contrast

259

� Contrast = difference in signal = difference in CT number between an object and the surrounding tissue

� = CT # B - CT # A

260

Image quality� C.RCT > C.RR� S.RCT< S.RR� Limiting spatial resolution for screen-film

radiography is about 7 lp/mm; for CT it is about 0.5 – 1.5 lp/mm

� Contrast resolution of screen-film radiography is about 5%; for CT it is about 0.5%

261

Factors affecting Contrast Resolution (↑)

� Noise (↓)� Voxel Size (↑): Pixel Size & Slice Thickness� kVp (↑)� Patient Movement (↓)� Reconstruction algorithms: (Bone filters

produce lower contrast resolution, and soft tissue filters improve contrast resolution)

CT Contrast Resolution Depends on Noise

� (A and B) Comparison of noise from scans using 270 mAs (typical clinical value) and 100 mAs.

� (C) Appearance of image noise is strongly affected by reconstruction filter; sharp filter such as bone also sharpens (enhances) appearance of noise. 262

263

CT Contrast Resolution Depends on NoiseImage Noise

� 120 kVp, 1.25mm, 0.5 sec

� 640 mA 25mA264

265

Small Contrast Difference Harder to Identify in Presence of Noise

266

CT Contrast Resolution

Contrast depends on noise

Noise depends on # photons detected

# photons detected depends on …

267

# of Photons Detected Depends Upon

� photon flux (x-ray technique)� slice thickness� patient size� Detector efficiency� Note:

• Good contrast resolution requires that detector sensitivity be capable of discriminating small differences in intensity

268

CT Image Noise� Fluctuation of CT

#’s in an image of uniform material (water)

� Usually described as standard deviation of pixel values

Image Noise� Variation in CT number in image of a uniform object

269 270

CT Image Noise

� Standard deviation of pixel values

S(xi - xmean)2

Noise (s) = -------------------(n-1)

Xi = individual pixel valueXmean = average of all pixel values in ROIn = total # pixels in ROI

271

Noise Level� Units

• CT numbers (HU’s)or

• % contrast

� Example• CT # range: 1000 HU’s• Standard deviation: 3 HU’s• Noise level is 3 or 3 / 1000 X 100 = 0.3%

272

CT Noise Levels Depend Upon

# detected photonsquantum noise

matrix size (pixel size) slice thickness algorithm electronic noise scattered radiation object size Photon flux to

detectors…

273

Photon Flux to Detectors

� Tube output flux (intensity) depends upon• kVp• mAs• beam filtration

� Flux is combination of beam quality & quantity

� Flux to detectors modified by patient

274

Slice Thickness� Thinner slices mean

• less scatter» better contrast

• less active detector area» less photons detected» More noise

� To achieve equivalent noise with thinner slices, dose (technique factors) must be increased

275

Noise Levels in CT:� Increasing slice thickness degrades spatial

resolution• less uniformity inside a larger pixel• partial volume effect

276

Noise Levels in CT:� When dose increases, noise decreases

• dose increases # detected photons� Doubling spatial resolution (2X lp/mm)

requires an 8X increase in dose for equivalent noise• Smaller voxels mean less radiation per voxel

CT Radiation Doses

277278

CT Radiation Doses

• Well-established relationship among SNR, pixel dimensions (), slice thickness (T), and radiation dose (D):

Artifacts in CT Imagesany systematic discrepancy between the CT numbers in the reconstructed image and the

true attenuation coefficients of the object

Artifacts in CT Images

• Artifacts can seriously degrade the quality of computed tomographic (CT) images, sometimes to the point of making them diagnostically unusable. To optimize image quality, it is necessary to understand why artifacts occur and how they can be prevented or suppressed.

280

Artifacts in CT Images

• CT artifacts originate from a range of sources.

1. Physics-based artifacts2. Patient-based artifacts3. Scanner-based artifacts4. Helical and multi section technique

artifacts

281

1. Physics-based artifacts

result from the physical processes involved in the acquisition of CT data

a. Beam Hardeninga. Cupping Artifactsb. Streaks and Dark Bands

b. Partial Volume

282

Beam Hardening

• As the beam passes through an object, it becomes “harder” , that is to say its mean energy increases, because the lower energy photons are absorbed more rapidly than the higher-energy photons

283

Beam Hardening

Changing energy spectrum of an x-ray beam as it passes through waterThe mean energy increases with depth(The attenuated spectra have been rescaled to be equivalent in size to the unattenuated spectra.)

284

285

Because the attenuation of bone is greater than that of soft tissue, bone causes more beam hardening than an equivalent thickness of soft tissue

Beam Hardening

Two types of artifact can result from this effect:

• cupping artifacts • the appearance of dark bands or

streaks between dense objects in the image

286

Cupping Artifacts• X rays passing through

the middle portion of a uniform cylindrical phantom are hardened more than those passing though the edges because they are passing though more material.

287

Cupping Artifacts

288

CT number profiles obtained across the center of a uniform water phantom

Streaks and Dark Bands• In very heterogeneous cross sections, dark

bands or streaks can appear between two dense objects in an image. They occur because the portion of the beam that passes through one of the objects at certain tube positions is hardened less than when it passes through both objects at other tube positions. This type of artifact can occur both in bony regions of the body and in scans where a contrast medium has been used

289

Streaks and Dark Bands

290

CT image shows streaking artifacts due to the beam hardening effects of contrast medium.

Minimizing Beam Hardening

• Filtration• Higher kVp• calibration correction• beam hardening correction software

291292

Filtration

• A flat piece of attenuating, usually metallic material is used to “pre-harden” the beam by filtering out the lower-energy components before it passes through the patient. An additional “bowtie” filter further hardens the edges of the beam, which will pass through the thinner parts of the patient.

293

Calibration correction

• Manufacturers calibrate their scanners using phantoms in a range of sizes. This allows the detectors to be calibrated with compensation tailored for the beam hardening effects of different parts of the patient.

294

Calibration correction of Cupping Artifacts

295

CT number profiles obtained across the center of a uniform water phantom without calibration correction (a) and with calibration correction (b).

Calibration correction of Cupping Artifacts

• Since patient anatomy never exactly matches a cylindrical calibration phantom, in clinical practice there may be either a slight residual cupping artifact or a slight “capping” artifact, with a higher central CT value due to overcorrection.

296

Beam hardening correction software

• An iterative correction algorithm may be applied when images of bony regions are being reconstructed

297

Beam hardening correction software

298

dark banding that occurs between dense objects when only calibration correction is applied (a) and the reduction in artifacts when iterative beam hardening correction is also applied (b)

Avoidance of Beam Hardening by the Operator

• It is sometimes possible to avoid scanning bony regions, either by means of patient positioningor by tilting the gantry

• It is important to select the appropriate scan field of view to ensure that the scanner uses the correct calibration and beam hardening correction data and, on some systems, the appropriate bowtie filter.

299300

Partial Volume• CT #’s based on linear attenuation coefficient for

tissue voxels• If voxel non-uniform (contains several materials),

detection process will average• Can appear as

incorrect densitiesstreaksbands

• MinimizingUse thinner slicesSoftware compensation

301302

2. Patient-based artifacts

are caused by such factors as patient movement or the presence of metallic materials in or on the patient.

a. Metallic Materialsb. Patient Motionc. Incomplete Projections

303

Metallic Materials

• Can lead to severe streaking artifacts• They occur because the density of the

metal is beyond the normal range that can be handled by the computer, resulting in incomplete attenuation profiles

304

Avoidance of Metal Artifacts by the Operator

• Patients are normally asked to take off removable metal objects such as jewelry

• For non removable items, such as dental fillings, prosthetic devices, and surgical clips, it is sometimes possible to use gantry angulation to exclude the metal inserts from scans of nearby anatomy

• When it is impossible to scan the required anatomy without including metal objects, increasing technique, especially kilovoltage, may help penetrate some objects, and using thin sections will reduce the contribution due to partial volume artifact

305

Software Corrections for Metal Artifacts

• Manufacturers use a variety of interpolation techniques to substitute the overrange values in attenuation profiles

306

Software Corrections for Metal Artifacts

307

CT images of a patient with metal spine implants, reconstructed without any correction (a) and with metal artifact reduction (b)

Patient Motion

• Patient motion can cause misregistration artifacts, which usually appear as shading or streaking in the reconstructed image.

• Steps can be taken to prevent voluntary motion, but some involuntary motion may be unavoidable during body scanning.

308

309

Patient Motion• Small motions cause image blurring• Larger physical displacements produce

artifacts that appear as double images or image ghosting

Patient Motion

310

Avoidance of Motion Artifacts by the Operator

• The use of positioning aids is sufficient to prevent voluntary movement in most patients.

• in some cases (eg, pediatric patients), it may be necessary to immobilize the patient by means of sedation.

• Using as short a scan time as possible helps minimize artifacts when scanning regions prone to movement.

• Respiratory motion can be minimized if patients are able to hold their breath for the duration of the scan.

311

Patient Motion

• The sensitivity of the image to motion artifacts depends on the orientation of the motion. Therefore, it is preferable if the start and end position of the tube is aligned with the primary direction of motion, for example, vertically above or below a patient undergoing a chest scan.

312

Built-in Features for Minimizing Motion Artifacts

Manufacturers minimize motion artifacts by using:

• Overscan and underscan modes• Software correction• Cardiac gating

313

Overscan and underscan modes

• The maximum discrepancy in detector readings occurs between views obtained toward the beginning and end of a 360° scan. Some scanner models use overscan mode for axial body scans, whereby an extra 10% or so is added to the standard 360° rotation. The repeated projections are averaged, which helps reduce the severity of motion artifacts. The use of partial scan mode can also reduce motion artifacts, but this may be at the expense of poorer resolution.

314

Software correction

• Most scanners, when used in body scan mode, automatically apply reduced weighting to the beginning and end views to suppress their contribution to the final image. However, this may lead to more noise in the vertical direction of the resultant image, depending on the shape of the patient. Additional, specialized motion correction is available on some scanners.

315

Software correction

316

CT images of the body created with conventional reconstruction (a) and with motion artifact correction (b).

Cardiac gating• The rapid motion of the heart can lead to

severe artifacts in images. To overcome these difficulties, techniques have been developed to produce images by using data from just a fraction of the cardiac cycle, when there is least cardiac motion. This is achieved by combining electrocardiographic gating techniques with specialized methods of image reconstruction.

317

Incomplete Projections

• If any portion of the patient lies outside the scan field of view, the computer will have incomplete information relating to this portion and streaking or shading artifacts are likely to be generated.

318

Incomplete Projections

319

CT image of the body obtained with the patient’s arms down but outside the scanning field shows streaking artifacts.

avoid artifacts due to incomplete projections

• It is essential to position the patient so that no parts lie outside the scan field. Scanners designed specifically for radiation therapy planning have wider bores and larger scan fields of view than standard scanners and permit greater versatility in patient positioning. They also allow scanning of exceptionally large patients who would not fit within the field of view of standard scanners.

320

3. Scanner-based artifacts

result from imperfections in scanner function.

a. Ring Artifacts

321

Ring Artifacts

• If one of the detectors is out of calibration on a third-generation scanner, the detector will give a consistently erroneous reading at each angular position, resulting in a circular artifact

322

323

Ring Artifacts• Causes

1 or more bad detectorssmall offset or gain

difference of 1 detector compared to neighbors•detector calibration required

• Reason: Rays measured by a given detector are all tangent to same circle

Ring Artifacts

• A scanner with solid-state detectors is in principle more susceptible to ring artifacts than a scanner with gas detectors.

324

Avoidance and Software Corrections

• recalibration the detector gain• repair services of detectors• software that characterizes and corrects

detector variations

325

4. Helical and multi section technique artifacts

• In general, the same artifacts are seen in helical scanning as in sequential scanning. However, there are additional artifacts that can occur in helical scanning due to the helical interpolation and reconstruction process. The artifacts occur when anatomic structures change rapidly in the z direction (eg, at the top of the skull) and are worse for higher pitches.

326

4. Helical and multi section technique artifacts

are produced by the image reconstruction process.

a. Helical Artifacts in the Axial Plane: Single-Section Scanning

b. Helical Artifacts in Multisection Scanningc. Cone Beam Effectd. Multiplanar and Three dimensional

Reformation327

328

New Stuff

CT Angiography CT fluoroscopy CT virtual endoscopy / colonoscopy

329

CT Angiography

This thoracic, abdominal, pelvic and lower-extremity CTA, was acquired in one acquisition of just 58 seconds using 0.5 second rotation time (8 slices/sec) and 2.5 millimeter thin slices for the entire scan.

330

Fluoroscopic CT

CT virtual endoscopy / colonoscopy

331

The end

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contents references - mums.ac.ir · image quality spatial resolution contrast resolution noise ct...

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