Learning Objectives
Optimization of Image Acquisition and
Reconstruction in Multi-slice CT
•
Clinical implementations of reconstruction
algorithms on multimulti-slice CT scanners.
Lifeng Yu, Cynthia H. McCollough,
Shuai Leng, James M. Kofler
•
How CT acquisition parameters affect image
reconstruction and image quality.
Department of Radiology, Mayo Clinic, Rochester
•
How reconstruction parameters affect image quality.
From SingleSingle-slice to MultiMulti-slice CT
Wide
cone beam
Narrow
cone beam
Fan beam
I. Clinical Implementations of Image
Reconstruction
1x5 mm
0.75 sec
1995
ConeCone-beam Effect
16o
320-slice CT
64x0.6 mm
0.33 sec
2004
>128x0.6 mm
0.27 sec
2008 - Present
ConeCone-beam Effect
~40 mm <20 mm
4o
64-slice CT
16x0.75 mm
0.42 sec
2001
z
160 mm
<2o
4, 8, 16-slice CT
4x1 mm
0.5 sec
1998
2D FBP
2D13
FBP
cm off center
AMPR
AMPR
13
cm off center
Courtesy D. Platten et al. ImPACT (RSNA 2003)
1
Analytical Reconstruction Methods
Fan-beam, parallel- beam FBP, rebinning
FBP, 180LI, 360LI
FDK, generalized FDK, Wedge
FBP
PI, PI-Slant, ASSR, nPI, AMPR
Various 3D weighted FBP or FDK-based
approximate methods
Analytical
Katsevich’s exact method
BPF
Minimum data exact reconstruction
Bontus et al, EnPiT: Filtered Back-Projection Algorithm for
Helical CT Using an n-Pi Acquisition, TMI, 2005
Clinical Implementations of Image
Reconstruction
• All current analytical reconstructions on clinical
scanners are approximate methods
• Goal is to maintain a good balance between
• computationally efficiency
• accuracy
• dose efficiency – fully utilize redundant data
Vendor Variations on Analytical
Reconstructions
• GE – 3D Weighted FBP (Tang et al, PMB, 2006)
• Philips – COBRA, nPI (Kohler et al, Med Phys, 2002;
Bontus et al, TMI, 2005)
• Siemens – AMPR (Flohr et al, Med Phys 2003) and
3D weighted FBP (Stierstorfer et al, PMB, 2004)
• Toshiba – TCOT, modified FDK (Taguchi et al, Med
Phys, 2003)
What can users control on image reconstruction?
II. Acquisition parameters
• Scan mode: Axial, helical, cine, shuttle, ECGECG-gated
• Rotation time (e.g., 0.28 sec, 0.35 sec, 0.5 sec)
• X-ray beam filter (e.g., flat filter, bowtie filter)
• X-ray beam collimation
• Detector configuration (e.g., 64x0.625 mm, 24x1.2 mm, 128x0.6 mm)
• Helical pitch (in helical mode); Table feed (in axial mode)
• Tube potential: kV
• Tube current: mA
• mAs:
mAs: Tube current (mA
(mA)) x Rotation time (s)
• Effective mAs:
mAs: mAs/pitch
mAs/pitch
• Automatic exposure control (AEC)
2
Scan Mode
„
„
Axial (or sequential)
Helical Pitch
Cine
Continuous
axial scan,
no table
move
„
Helical (or spiral)
„
Pitch =
Shuttle or 4D spiral
Non-cardiac spiral MDCT: pitch=1
Gantry rotation angle (deg)
270
180
90
1
2
3
Non-cardiac spiral MDCT: pitch=0.75
Pitch and Image Quality
270
• Higher pitch – higher scan speed
• For nonnon-cardiac exams, a higher pitch
180
90
0
4
0
Gantry rotation angle (deg)
180
90
1
2
3
3
4
360
270
• Downside of higher pitch
• Lower dose capacity
• More severe helical artifacts
• More severe conecone-beam artifacts If
90
combined with wide collimation
0
4
scan has a higher temporal resolution.
(windmill)
180
0
z-position in units of the total nominal beam width
1
2
3
4
z-position in units of the total nominal beam width
Primak A et al, RadGraphics, 2006
Detector Configuration
GE
64 x 0.6
20 mm
Sens-16
4 x 1.5
4 x 1.2
Def or 64
4 x 1.2
20 mm
Z-flying focal spot
SIEMENS
AS+ or
Flash
LS-8
8x1.25 or 8x2.5
LS-16
4 x 1.25
LS-64
4 x 1.25 16 x 0.625
0
2
Non-cardiac spiral MDCT: pitch=1.5
270
64 x 0.625
Gantry rotation angle (deg)
Non-cardiac spiral MDCT: pitch=0.5
360
0
1
z-position in units of the total nominal beam width
z-position in units of the total nominal beam width
16 x 0.75
0
Pitch > 1
4 x 1.5
0
Pitch = 1
360
32 x 0.6
Gantry rotation angle (deg)
Pitch < 1
360
Table translation per rotation
Detector collimation
24 mm
28.8 mm
40 mm
38.4 mm
Flohr et al, Image reconstruction and image quality evaluation for a 64-slice
CT scanner with z-flying focal spot. Med Phys 32, 2536, 2005.
3
Effect of Z FFS
NO Z FFS
Detector configuration and electronic noise
With Z FFS
Flohr et al, Image reconstruction and image quality evaluation for a 64-slice
CT scanner with z-flying focal spot. Med Phys 32, 2536, 2005.
Wider detector collimation
• Routinely used for faster speed and better use of tube
power (64x0.625 mm, 128x0.6* mm, …)
128 x 0.6* mm
Equal CTDIvol
32 x 1.2 mm
ConeCone-beam Artifacts in Wider
Collimation (4 cm)
• Tends to generate more severe artifacts if not wellwellcorrected
• Helical and conecone-beam artifacts, particularly for highhighcontrast object at high pitch
• Some applications, narrower collimation is better
• e.g., 32x0.625 mm is often a better choice for head CT
than 64x0.625 mm
• Use slower pitch for wide collimation if there is no
scan time concern.
Pitch=0.983
Pitch=0.516
Narrower Collimation (2 cm)
Pitch=1.375
X-ray Beam Bowtie Filter
X-ray source
bowtie filter
flat filter
• Small bowtie
• Ped Head, Cardiac Small,
Small Head, Ped Body, Small
Body
• Medium bowtie
• Head, Cardiac Medium,
Medium Body
Patient
Pitch=0.562
Pitch=0.969
• Large bowtie
• Large body, Cardiac Large
Pitch=1.375
Detector
4
mA,
mA, mAs,
mAs, effective mAs,
mAs, mAs/slice
mAs/slice
• mAs:
mAs: Tube current (mA
(mA)) x Rotation time (s)
• Effective mAs = mAs/pitch
mAs/pitch
• Effective mAs is proportional to the radiation
output (usually expressed as CTDIvol)
• mA or mAs cannot be related to CTDIvol without
knowing the pitch
Effective mAs 70
Effective mAs 450
mAs vs. Noise in Cardiac and NonNoncardiac Exams
• In nonnon-cardiac exams, image noise is directly
Noise: 22.3
240 effective mAs
Noise: 25.3
180 effective mAs
related to effective mAs (mAs/pitch).
mAs/pitch).
• Redundant data is used when pitch is less than 1
• In cardiac exams, image noise is directly
related to mAs.
mAs.
• Independent of Pitch
• Each reconstruction uses data from a halfhalf-scan
range.
• No redundant data is used even when pitch is
much less than 1.
Noise: 30.8
120 effective mAs
Noise: 46.4
60 effective mAs
Photon Starvation
B10
Adaptive Filter in Reconstruction
B18
5
Automatic Exposure Control
(AEC)
X-ray attenuation
• Varies over body region and with projection angle
• Image noise is primarily determined by noisiest projections
(thick body parts)
• More photons (dose) to thinner body parts is unnecessary
radiation dose
Automatic tube current modulation
Automatic Tube Current Modulation
• Goal of tube current modulation is to dynamically adjust
tube current according to the attenuation level.
⎛ A(d ) ⎞
⎟⎟
General strategy I (d ) = I (d 0 ) ⋅ ⎜⎜
⎝ A(d 0 ) ⎠
x=0.5
x=1
– Higher attenuation, higher tube current
– Lower attenuation, lower tube current
x=0.4
x=0.3
x
x=0
A( d ) = exp( µd )
where
d0 is the reference attenuation path length. x is the modulation strength
If
If
x=0
x>0
x =1
then
then
I ( d ) = I (d 0 )
Constant tube current
Attenuation
decreases
tube current is modulated according to attenuation
Attenuation
increases
equal noise for each view independent of attenuation
M. Gies, W. A. Kalender, H. Wolf and C. Suess, "Dose reduction in CT by anatomically
adapted tube current modulation. I. Simulation studies," Med Phys. 26, 2235-47 (1999)
Topogram Evaluation: a.p. and lateral
Optimal mA: a.p. and lateral
4000
a.p. (measured)
a.p.
3500
lateral (calculated)
max
2000
1500
1000
atten u atio n I_0 / I
2500
tub e c u rr e nt in m A
lateral
3000
500
0
600
500
400
300
table position
200
100
0
600
500
400
300
200
100
0
table position
6
402 mA
544 mA
On-line tube current modulation
400
4000
Attenuation
3500
tube current
300
3000
250
2500
200
2000
150
1500
100
1000
50
500
0
attenuation I_0 / I
tube current
350
0
600
500
400
300
200
table position in m m
100
0
678 mA
Streaking Artifact Reduction
Different Versions of AEC
Manufacturer
Constant Tube Current
AEC
452 mA
AEC Trade name
Image Quality
reference
Goal
GE
Auto mA, Smart
mA
Noise Index
Maintain a constant noise level
(defined in noise index), using tube
currents within prescribed
minimum and maximum values.
Toshiba
SureExposure
Standard deviation
(high quality,
standard, low dose)
Maintain a constant noise level
(defined in standard deviation
values for each protocol), using
tube currents within preset
minimum and maximum values.
Siemens
CARE Dose4D
Quality Reference mAs
Maintain the same image quality
(varying noise target for different
attenuation level) with reference to
a target effective mAs level for a
standard-sized patient.
Philips
DoseRight
Reference Image
Keep the same image quality as in the
saved reference image.
3.00E+05
80 kVp
100 kVp
120 kVp
140 kVp
Tube Potential (kV)
x-ray intensity
2.50E+05
2.00E+05
1.50E+05
1.00E+05
5.00E+04
0.00E+00
0
50
100
150
En ergy (keV)
7
Iodine Contrast
LowerLower-kV Benefit – Increased Iodine Contrast
Iodine Contrast vs. kVp
120 kV, CTDIvol=5.18 mGy
300
100 kV, CTDIvol=3.98 mGy
250
Contrast
200
Small
Medium
150
Large
100
50
0
60
80
100
120
140
The same patient scanned with a protocol at 120 kV and a protocol at 100 kV. Note that
improved contrast and visualization of mural stratification of the 100 kV image despite a
25% radiation dose reduction.
160
kVp
Phantom size: small 30x13 cm, medium 30x21 cm, large, 36x27 cm
Limitation of Lower kV – increased noise
Noise
80kV CTDIvol=6.4 mGy
Noise vs. kVp (CTDIvol=23mGy)
120kV CTDIvol=6.5 mGy
45
40
35
Noise (HU)
30
Small
25
Medium
20
Large
15
10
5
0
60
80
100
120
140
160
kVp
Optimal kV
• Optimal kV is the kV that uses the
minimum radiation dose to achieve the
desired image quality
• Dependent on patient size
• Dependent on diagnostic task
A General Strategy for Optimal kV Selection
• Iodine CNR with a noise constraint α:
CNR (Ω d , kV , D ) ≥ CNR (Ω d , kVref , Dref )
and
σ (Ω d , kV , D ) ≤ α ⋅ σ (Ω d , kVref , Dref )
• The corresponding relative dose factor
(RDF) at each kV is given by
⎧ C (Ω d , kVref ) 1 ⎫ k (Ω d , kV )
RDF (Ω d , kV ) = max ⎨
, ⎬⋅
⎩ C (Ω d , kV ) α ⎭ k (Ω d , kVref )
L. Yu, H. Li, J. Fletcher, C. McCollough, Automatic selection of tube potential for
radiation dose reduction in CT: A general strategy. Med Phys, 37:234-43, 2010.
8
Optimal kV for Abdominal CT
Modulation
strength
25 cm
30 cm
35 cm
40 cm
45 cm
50 cm
55 cm
Very weak
120
120
120
120
120
120
140
Weak
100
100
100
120
120
120
140
average
100
100
100
100
120
120
140
Strong
80
80
100
100
100
120
140
Very strong
80
80
80
100
100
120
140
•
•
•
•
•
Very weak:
weak:
average:
strong:
Very strong:
Routine nonnon-contrast exams
Liver/pancreas
Routine contrastcontrast-enhanced exams
CT enterography,
enterography, CTU, stone
CT angiography
Slice Thickness
III. Reconstruction parameters
„
Nominal width of image (z-direction)
Slice sensitivity profile
• Slice thickness/Increment
• Reconstruction kernel
• Reconstruction FOV
• 3D reformat
• Iterative reconstruction or noise
(line spread function along Z)
reduction options
Slice Thickness, Noise, and Dose
Slice Thickness – Partial Volume
Image (mm): 10
Noise ∝
5
2.5
1.25
3.84
5.89
7.82
1
# Photons
Slice thickness (mm):
5
Relative Noise:
100%
Required mAs:
100%
(for = noise)
2.5
141%
200%
1.25
200%
400%
0.625
283%
800%
„
Better z-resolution (less partial vol. averaging)
„
Increased image noise
„
Requires more radiation dose to get the same noise level
Noise (HU): 2.93
All other parameters are identical
„
Thinner slices => less partial volume effect
9
Slice Thickness – Partial Volume
Slice Thickness – Partial Volume
10mm image thickness
5mm image thickness
All other parameters are identical
All other parameters are identical
Slice Thickness – Partial Volume
Slice Thickness – Partial Volume
2mm image thickness
1mm image thickness
All other parameters are identical
All other parameters are identical
Slice Thickness – Partial Volume
Reconstruction kernel
• Reconstruction kernel (filter or algorithm) has a
significant impact on spatial frequency and noise
characteristics of an image
• Smooth kernels reduce high spatial-frequency
information and image noise
• Sharp kernels increase high spatial-frequency
information and image noise
0.6mm image thickness
All other parameters are identical
10
Impact of Reconstruction Kernel on In-plane
spatial resolution
Modulation Transfer Function (MTF)
B10
B20
B30
B40
1
MTF
0.8
0.6
0.4
0.2
0
0
A smoother kernel
A sharper kernel
Soft
CT number: 89.7±2.2
2
4
6
8
10
spatial frequency (1/cm)
Standard
CT number: 89.7±2.6
Detail
Bone
CT number: 90.7±3.5
CT number: 90.6±8.9
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Some Considerations on GE Reconstruction
algorithms
Soft
• Lung and bone plus alter CT # values
• For high res lung use bone
• Standard is most frequently used
• Soft is to recover a really grainy exam
• Use Full recon for getting the slice width you
Standard
Detail
Lung
Bone
ask for
• Plus mode can help smooth noisy images or
Edge
Bone Plus
Siemens Reconstruction Kernels
• B10 Æ B90 Body
• H10 Æ H90 Head
• U30 Æ U90 Ultra High Resolution
• T20 Æ T81 Topogram
• Lower number smoother
• Higher number sharper
• Multiples of 10 are the “basic”
basic” kernels
• In between values are “special”
special” kernels
reduce some artifacts, but is up to 20% wider
slice thickness
Reconstruction FOV & Pixel Size
0.78 mm
0.78 mm
40 cm
40 cm
0.39 mm
0.39 mm
20 cm
20 cm
Pixel size & System Spatial Resolution
• Pixel size should not be confused with system spatial resolution
• Pixel size should be small enough in order not to limit the display
display of
3D Reformat
system resolution.
• 512 x 512 matrix
Pixel Size
(mm)
Pixel size
Resolution (lp/cm)
50
1.0
5.1
40
0.8
6.4
30
0.6
8.5
20
0.4
12.8
rFOV
(cm)
System spatial
Resolution at 10%
MTF (lp/cm)
B20
5.5
B30 B40
6.2
7.0
Reformat from 1 mm
axial reconstruction
Reformat from 2 mm
axial reconstruction
12
Iterative Reconstruction
• Benefits over analytical reconstruction
• Accurate physical models (imaging system geometry,
Vendor implementations of Iterative
Reconstruction or noise reduction methods
• GE: ASIR (adaptive statistical iterative
reconstruction), MBIR (model based iterative
scattering, spectrum, etc.)
• less bias, improved quantitative accuracy, improved spatial
resolution
• Appropriate statistical models (FBP treats all data equally)
• Lower image noise, lower radiation dose
• Object constraints (e.g., nonnon-negativity)
• Flexible to handle nonnon-standard scanning configurations
• Potential to reconstruct images from sparse or incomplete
data
reconstruction)
• Phillips: iDose
• Siemens: IRIS (iterative reconstruction in image
space), SAFIRE (sinogram affirmed iterative
reconstruction)
• Toshiba: AIDR3 (adaptive iterative dose reduction)
Source - SPIE Medical Imaging Symposium, 2011
Example – Image quality improvement
Standard FBP
Noise Texture Change
MBIR
Images provided by Rendon Nelson, MD, Duke University, to AuntMinnie.com (“MBIR aims to
outshine ASIR for sharpness, CT dose reduction,” May 18, 2010)
LowLow-contrast Detectability
50% dose reduction possible?
Hara, AK et al (2009). "Iterative reconstruction technique for reducing body
radiation dose at CT: feasibility study." AJR Am J Roentgenol 193(3): 764-71.
Special Considerations in Iterative Reconstruction
• Potential for dose reduction ~ 20% - 50% - task dependent
• HighHigh-contrast allows more
• LowLow-contrast allows less or none
• Noise texture could change dramatically
• NonNon-linear regularization term leads to nonnon-linear noisenoiseresolution tradeoff
• Spatial resolution is highly dependent on contrast level
• Could reduce noise substantially without sacrificing highhighcontrast spatial resolution (measured in MTF)
• LowLow-contrast detectability needs to be evaluated more carefully.
• Quantitative evaluation and optimization of iterative
FBP, 25 mGy
ASIR, 12.5 mGy
reconstruction is still an active research field
Hara, AK et al (2009). "Iterative reconstruction technique for reducing body
radiation dose at CT: feasibility study." AJR Am J Roentgenol 193(3): 764-71.
13
Summary
• Implementations of reconstruction algorithms on
clinical scanners
• How data acquisition parameters affect image quality
• How reconstruction parameters affect image quality
• Understanding of these is essential for designing dosedose-
Thank you!
http://mayoresearch.mayo.edu/ctcic
http://mayoresearch.mayo.edu/ctcic//
efficient and quality scanning protocols for various
diagnostic tasks.
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