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IAEAInternational Atomic Energy Agency

Slide set of 148 slides based on the chapter authored by

J.C. Lacefield

of the IAEA publication (ISBN 978-92-0-131010-1):

Diagnostic Radiology Physics:

A Handbook for Teachers and Students

Objective:

To familiarize the student with practical issues associated

with ultrasound imaging.

Chapter 13: Ultrasound Imaging

Slide set prepared

by E. Berry (Leeds, UK and

The Open University in

London)

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CHAPTER 13 TABLE OF CONTENTS

13.1. Introduction

13.2. Array System Principles

13.3. B-Mode Instrumentation and Signal Processing

13.4. Modern Imaging Methods

13.5. Colour Flow Imaging

13.6. Image Artifacts and Quality Assurance

Diagnostic Radiology Physics: A Handbook for Teachers and Students – 13. Slide 1 (02/148)

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13.1 INTRODUCTION13.1

Diagnostic Radiology Physics: A Handbook for Teachers and Students – 13.1 Slide 1 (03/148)

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Introduction – conventional ultrasonography

� Images are acquired in reflection, or pulse-echo, mode

� An array of small piezoelectric elements transmits a

focused pulse along a specified line of sight known as a

scan line

� Echoes returning from the tissue are received by the same

array

• focused via the delay-and-sum beamforming process

• demodulated to obtain the magnitude, or envelope, of the echo

signal


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Introduction – A-mode

� The scanner measures the arrival time of the echoes

relative to the time the pulse was transmitted

• maps arrival time to distance from the array using an assumed

speed of sound

� One-dimensional A-mode (amplitude mode) format

• result of a single-pulse acquisition

• plot echo magnitude as a function of distance


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Introduction – B-mode

� Two-dimensional (2D) or three-dimensional (3D) B-mode

(brightness mode) image

• acquired by performing a large number of pulse-echo acquisitions

• incrementing the scan-line direction between each pulse-echo

operation

• this sweeps out a 2D or 3D field of view

� B-mode imaging because

• the echo magnitude from each point in the field of view is mapped

to the gray level, or Brightness, of the corresponding pixel in the

image


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13.2 ARRAY SYSTEM

PRINCIPLES13.2.1 ELECTRONIC FOCUSING AND BEAMSTEERING

Diagnostic Radiology Physics: A Handbook for Teachers and Students – 13.2.1 Slide 1 (07/148)

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13.2 ARRAY SYSTEM PRINCIPLES13.2.1 Electronic Focusing and Beamsteering

Beamsteering

� The array transducers employed by modern ultrasound

systems enable the use of high-speed electronic focusing

and beamsteering methods

• the basis of the high frame rates achieved by ultrasound

� Beamsteering increments the direction of the scan line to

sweep out the B-mode field of view

� Details of the beamsteering process differ slightly

depending which of the three major types of array is used:

• linear array

• curvilinear array

• phased array


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Ultrasound array and B-mode image geometries.

(a) Linear array, (b) Curvilinear array, (c) Phased array


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Beamsteering with linear array

(1 of 2)

� A linear array consists of up to

256 elements in a single row

� Each pulse-echo operation is

performed by selecting a small

subaperture of adjacent element

� The scan line is always directed

perpendicular to the array from

the centre of the active

subaperture

• along the axial dimension of the

image

� Scan line is stepped across the

face of the array between each

pulse-echo acquisition

Linear array



� Active subaperture: dark grey

� Inactive portion of array: light

grey

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Beamsteering with linear array

(2 of 2)

� Stepping of scan line achieved

by

• deactivating an element at one

end of the subaperture

• activating a new element at the

opposite end of the subaperture

� This procedure yields a

rectangular field of view whose

lateral width is comparable to the

length of the array

Linear array





grey

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Beamsteering with curvilinear

array

� Curvilinear array functions in a

manner analogous to a linear

array, except the face of the

array is convex rather than

planar

� As the subaperture is stepped

across the array, the scan line

both translates laterally and

rotates in azimuth angle

� This implementation produces a

circular sector image with a wide

field of view at all depths

Curvilinear array





grey

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Beamsteering with phased

array (1 of 4)

� Elements of a phased array are

also co-planar but

• smaller, less numerous, and

more closely spaced than the

elements of a linear array

� Each scan line originates from

the centre of the array and is

acquired using all of the

elements

� The azimuth angle of the scan

line, θ, is incremented by altering

the relative timing of the pulses

transmitted by each element

Phased array





grey

� Azimuth angle, θ

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Beamsteering with phased

array (2 of 4)

� Circular sector field of view

� Spans as much as 90° in

azimuth angle

� Converges to a narrow field of

view at shallow depths, near

the top of an image

Phased array





grey

� Azimuth angle, θ

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Beamsteering with phased array (3 of 4)

� All array types are laterally focused: beam of a phased

array is steered using delay-and-sum beamformation

� During transmission

• ensure that the pulses transmitted from each element arrive

simultaneously at the focal point

• so result is constructive interference

� Relative firing time of element to firing time at centre of

aperture determined by

• computing the distance from each element to the focal point using

elementary trigonometry

• assume the speed of sound, c = 1540 ms-1 (average in soft tissue)


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Beamsteering with phased array (4 of 4)

� During reception

• similar delays are applied to temporally align the echo signals

received by each array element from the intended focal point

• the delayed signals are then summed

• echoes received from the focus add constructively

• this yields a beamformed receive signal


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13.2 ARRAY SYSTEM

PRINCIPLES13.2.2 ARRAY BEAM CHARACTERISTICS


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13.2 ARRAY SYSTEM PRINCIPLES13.2.2 Array Beam Characteristics

Modelling Array Beam Characteristics

� The lateral beam patterns produced by a linear or phased

array can be modelled

• apply linear systems analysis to the Fourier transform relationship

between aperture and beam pattern that arises from the

Fraunhofer diffraction integral

� A single element can be viewed as a narrow rectangular

aperture of width d, represented mathematically by

rect(ξ/d)

• where ξ is the lateral dimension within the aperture plane

• the orthogonal dimension (η) is ignored


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Aperture function

� To model an infinite array with element spacing s.

• convolve rect(ξ/d) with a sampling, or comb, function, comb(ξ/s)

� To produce a linear-systems description of a 1-D array

extending from ξ = −L/2 to ξ = L/2

• multiply by a broad rect function, rect(ξ/L)

� Result is A(ξ), known as the aperture function


( )

∗

=

LsdA

ξξξξ rectcombrect

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Lateral beam profile in focal plane

� Apply the Fraunhofer diffraction integral

� The lateral beam profile in the focal plane, U(x), is

proportional to the Fourier transform of A(ξ)

where sinc(a) = sin(πa)/(πa) and constant terms are omitted in the

interest of compact notation


( )

∗

∝

F

dxd

F

sxs

F

LxLxU

λλλsinccombsinc

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Lateral point spread function

� The same aperture function is usually employed for

transmission and reception

� So lateral point spread function PSF(x) is given by square

of the lateral beam profile in the focal plane, U(x)

� Can plot normalized point spread function |PSF(x)|/|PSF(0)|


( )2

sinccombsinc

∗

∝

F

dxd

F

sxs

F

LxLxPSF

λλλ

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Primary beam and main lobe

Lateral point-spread function,

PSF(x), of an unapodised linear

array



� f-number, F/L=1

� LR: lateral resolution

� PSL: peak side-lobe level

� Decibel scale

� The primary beam arises from

the sinc(Lx/λF) term in PSF

equation

� The largest central peak of

the primary beam is called the

main lobe of the PSF

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Lateral resolution



array



� f-number, F/L=1

� LR: lateral resolution

� PSL: peak side-lobe level

� Decibel scale

� Under the Rayleigh resolution

criterion

• lateral resolution of the

imaging system, LR, is equal

to the distance from the

maximum to the first zero of

the main lobe

� Setting sinc(Lx/λF) = 0 and

solving for x

� as previously found for a single-

element rectangular aperture

L

FLR

λ=

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Contrast resolution Lateral point-spread function,


array



� Side lobes: Peaks

flanking the main lobe in

the primary beam

� PSL: peak side lobe level

• expressed in decibels

• indication of the contrast

resolution of the imaging

system

( )( )( )( )

=

mainlobe within max

mainlobe outside maxlog20 10

xPSF

xPSFPSL

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Contrast resolution Lateral point-spread function,


array



� The image contrast expected for

a small anechoic target

� For the PSF shown here

• the largest side lobe arises from

the first off-axis peak of the

primary beam

• PSL = −26.6 dB

( )( )( )( )

=

mainlobe within max


xPSF

xPSFPSL

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Apodisation of aperture

� Reduces the side lobe

levels

� Weight magnitudes of

the transmit and receive

signals

• on the outer elements

• during delay-and-sum

beamformation

• with a function that

decreases more gradually

than rect(ξ/L)

Lateral point-spread function of

a linear array with f-number

equal to one and Hamming

window apodisation



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� E.g. Hamming window is

applied for apodisation

during both transmit and

receive beamforming

• peak side-lobe level is

reduced by about 56 dB

• the lateral resolution is

about twice as coarse as

the resolution of the

uniformly weighted

aperture

No apodisation



Hamming window apodisation

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� Array-based ultrasound systems almost always employ

some form of apodisation because contrast resolution is a

crucial design consideration for diagnostic tasks such as

lesion detection


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Grating lobes


PSF(x), showing grating lobes



� f-number, F/L=1

� Unapodised

� Decibel scale

� Convolution of the sinc(Lx/λF)

term with the sampling

function produces copies of

the primary beam called

grating lobes

� Spaced by λF/s along the

lateral dimension of the focal

plane

( )2

sinccombsinc

∗

∝

F

dxd

F

sxs

F

LxLxPSF

λλλ

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Grating lobes

� Under the paraxial approximation

� The azimuth angle, θm, of the mth grating lobe away from

the main lobe peak is

� Grating lobes produce echoes from off-axis targets that

appear as artifacts in images

• See section 13.6

≈ −

s

mm

λθ 1sin

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Grating lobes and undersampling

� Mathematically analogous to the copies of spectral peaks

that are observed in the discrete-time frequency spectrum

of an undersampled signal

� In an array system, the aperture function is undersampled

if the element spacing, s, is too large

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Grating lobes: linear array design rule

� For a linear array

• beam is focused but not steered

• grating lobes are avoided if the azimuth angle of the first-order

grating lobe is greater than π/2 radians

� If m = 1, solving for s yields the linear-array design rule for

element spacing: s ≤ λ

2sin 1 πλ

θ ≥

≈ −

s

mm

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Grating lobes: phased array design rule

� For a phased array

• beam is steered

• most challenging scenario for grating lobe suppression arises

when the beam is steered as far as possible in azimuth angle

• i.e. parallel to the face of the array

� It can be shown that s ≤ λ/2� This is the phased-array design rule for element spacing

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Elevation (out-of-plane) dimension of beam

� Analyse in a similar fashion to the lateral dimension

� Conventional linear and phased arrays possess only one

row of elements

• electronic focusing and beamsteering techniques, including

apodisation, cannot be applied in the elevation dimension

� Elevational focusing is performed by an acoustic lens

positioned at the face of the array

� Since the height of an element is substantially smaller than

the lateral aperture length

• spatial resolution in the elevation dimension is relatively coarse (of

the order of several millimetres)

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Depth of field in elevation dimension

� Depth of field is the interval over which the beam remains

in focus

� Depth of field is proportional to the wavelength multiplied

by the square of the f-number

� In view of the fixed focal distance of the lens, one benefit

of the weaker focusing in elevation is that the depth of field

is larger in the elevation dimension than the lateral

dimension

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Electronic focusing in elevation dimension

� Electronic focusing in the elevation dimension is possible

using multirow linear arrays

• technology that emerged in the late 1990s

• elements of these arrays are divided into a small number of rows

(typically 5 to 9) in the elevation dimension

• permits electronic focusing, but not beamsteering, to be performed

in the elevation dimension

� Compared to a one-dimensional array with an elevation

lens, multirow arrays enable the depth of the elevational

focus to be changed

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Electronic focusing in elevation dimension: pros and cons

� Valuable capability for the multi-focal imaging methods

(section 13.2.3)

� However, use of multirow arrays also increases the system

complexity

• arrays with smaller elements are more difficult to manufacture

• the scanner must include additional channels of transmitter and

receiver electronics

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13.2 ARRAY SYSTEM

PRINCIPLES13.2.3 MULTI-FOCAL IMAGING METHODS


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13.2 ARRAY SYSTEM PRINCIPLES13.2.3 Multi-Focal Imaging Methods

Multi-focal imaging

� The control panel of a modern clinical scanner provides

the user with some flexibility for selecting the number and

depths of transmit foci

� Imaging with multiple transmit focal zones is achieved by

• acquiring each scan line repeatedly with the transmit beam

focused at a different depth along the scan line

• constructing a composite scan line which consists, at each point

along the line, of the pulse-echo sample acquired using the nearest

transmit focus

• if a multirow array is used, the elevation as well as the lateral focus

position can be changed for each pulse echo acquisition


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Multi-focal imaging: Pros and cons

� Imaging with multiple transmit zones produces an image

with more consistent resolution throughout the field of view

� But this improvement comes at the expense of a reduced

frame rate


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Dynamic receive focusing (1 of 2)

� All modern scanners also employ dynamic receive

focusing

� The arrival time of echoes at the receive aperture

corresponds to the depth of the scatterers that produced

the echoes

� So the receive focusing delays are updated in real time

such that the receive focus tracks the pulse as it

propagates along each scan line


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Dynamic receive focusing (2 of 2)

� The receiver delays can be updated continuously as a

scan line is acquired

� So dynamic receive focusing means that the lateral width

of the receive beam (and its elevation width if a multirow

array is used) should be minimised everywhere in the field

of view

� Use of dynamic receive focusing has no effect on frame

rate since the receive focus is produced computationally

after a pulse-echo signal has been acquired


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Frame rate

� The frame rate of an ultrasound system is determined by

• the lateral resolution

• the field of view

• the number of transmit focal zones

� In general, reduce the frame rate by imaging

• over a larger field of view

• with higher spatial resolution

• with a greater number of transmit focal zones


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Frame rate dependencies

� The total time to acquire one scan line can be

approximated by the round-trip pulse travel time to and

from the maximum depth in the image multiplied by the

number of transmit focal zones

� The lateral spacing of adjacent scan lines should be no

greater than one half of the lateral resolution at the focus

to ensure adequate spatial sampling

� The number of scan lines in one frame therefore depends

on the lateral field of view and the lateral resolution

� The time to acquire one frame is the product of the number

of scan lines and the total time to acquire one line Diagnostic Radiology Physics: A Handbook for Teachers and Students – 13.2.3 Slide 7 (44/148)

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13.3 B-MODE INSTRUMENTATION

AND SIGNAL PROCESSING13.3


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13.3 B-MODE INSTRUMENTATION AND SIGNAL

PROCESSING 13.3

Block diagram of a B-mode ultrasound imaging system


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PROCESSING 13.3

B-mode ultrasound imaging (1 of 3)

� High-voltage excitation pulses are applied to the array by

the transmitter electronics to fire a pulse

� The transmit/receive switch then disconnects the array

from the transmitter electronics and connects it to the

receiver electronics, thereby isolating the receiver

electronics from the high-voltage excitation pulses

� The returning echo signals are immediately amplified,

digitized, and then combined via delay-and-sum

beamforming to produce a beamformed radio-frequency

(RF) signal


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PROCESSING 13.3


� Time-gain compensation is applied to the beamformed RF

signal to compensate for attenuation of echoes arriving

from deeper targets

� Envelope detection is performed to obtain the magnitude

signal

� Signal is logarithmically compressed to maximize the

dynamic range of the image


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PROCESSING 13.3


� Scan conversion is performed to map

• each sample of the magnitude signal to its 2D or 3D position in the

image

• the log-scaled magnitude values to gray levels

� Once the scan lines covering the entire field of view have

been acquired, the resulting image frame is displayed on

the scanner’s video monitor


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PROCESSING 13.3

Transmitter electronics

� Design requirements

• High bandwidth to enable use of short excitation pulses to obtain

high axial resolution

• An ability to operate at high power to drive the cable connecting

the scanner and transducer with 50−100 V excitation signals

� Some modern imaging methods employ coded transmit

pulses (section 13.4.3)

• require more sophisticated transmit electronics to support the use

of programmable excitation waveforms

• in conventional B-mode imaging a simple excitation waveform is

used and the spectral characteristics of the transmitted pulse are

determined primarily by the frequency response of the transducer


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PROCESSING 13.3

Receiver electronics

� Design requirements for digital processing

• sampling at sufficiently high frequency to sample the highest

frequency components of the RF signal

• e.g. 25-50 MHz depending on the pulse spectrum

• quantization using at least 12, and preferably 16, bits to represent

the RF waveform accurately

� Analogue-to-digital conversion is preceded by an anti-

aliasing low-pass filter

• the cut-off frequency of the anti-aliasing filter is often a significant

determinant of the ultimate axial resolution of the image


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PROCESSING 13.3

Beamforming

� Beamforming delays must be applied with high precision

to obtain a sharp focus

� The typical procedure is to interpolate the digitized RF

signals to a sampling frequency as high as 1 GHz

• which would enable focusing delays to be applied with 1 ns

precision

� Then downsample the time-shifted signals to the original

sampling frequency for subsequent processing

� Relatively simple interpolation methods must be used to

enable this step to be performed in real time for dynamic

receive focusing


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PROCESSING 13.3

Amplification

� Three types of amplification are applied to the

beamformed RF signal

� Constant gain

� Time-gain compensation (TGC)

� Lateral gain compensation


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PROCESSING 13.3

Constant gain

� A constant gain is applied to enable the user to adjust the

overall brightness of the image

� Often simply labelled “Gain” on the scanner’s control panel


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PROCESSING 13.3

Time-gain compensation (TGC)

� Time-dependent amplification applied to each scan line to

offset the effects of attenuation

� The reduction in signal intensity from attenuation decays

as an exponential function of propagation distance

• so the TGC should be an approximately exponential function of the

echo arrival time (hence the T in TGC)

� TGC slope is set separately in several depth bands

covering the field of view by adjusting the TGC to obtain

visually consistent brightness throughout the image


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PROCESSING 13.3

Lateral gain compensation

� Lateral gain compensation is applied to compensate for

shadowing or enhancement artifacts (see section 13.6)

that cause image brightness at a given depth to vary as a

function of lateral position

� The lateral gain is adjusted manually by the user in a

manner similar to the TGC


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PROCESSING 13.3

Computation of magnitude signal

� Modern scanners typically compute the magnitude signal

using the Hilbert transform

• applies an exact π/2 phase shift to the RF signal

• thereby estimating its quadrature component

� The instantaneous magnitude at each time sample is then

obtained by adding the original and Hilbert-transformed

signals in quadrature

� Prior to widespread use of digital processing

• envelope detection via analogue amplitude demodulation

• e.g. full-wave rectification followed by low-pass filtering of the

beamformed RF signal


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PROCESSING 13.3

Dynamic range of magnitude signal

� The magnitude signal is logarithmically compressed

� The dynamic range of the pulse-echo data can be greater

than 80 dB

• much greater than the 48 dB (= 20log10(256)) dynamic range of a

standard 256-gray-level display

� The scanner’s electronics could map the data to greater

than 256 gray levels

• little benefit to this approach as the human visual system has

limited sensitivity to subtle differences in brightness


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PROCESSING 13.3

Compression of magnitude signal

� An effective algorithm for logarithmic compression is to

• use the mean of the magnitude signal (averaged over the entire

field of view) as the reference magnitude

• convert each of the linearly scaled magnitude samples to decibels

with respect the reference value

� The dB values are linearly converted to gray levels such

that

• a magnitude of –X/2 dB maps to gray level 0 (black)

• 0 dB (i.e., the mean magnitude) maps to gray level 128

• X/2 dB maps to gray level 255 (white)

� X is the displayed dynamic range, typically 60 or 80 dB


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PROCESSING 13.3

Interpolation – Linear array

� For the rectangular images produced by a linear array

system

• spatial sampling of the original echo data is much coarser in the

lateral dimension compared to the axial dimension due to the

anisotropic spatial resolution of an ultrasound scanner

• if the scan lines are displayed directly, the rectangular aspect ratio

of the pixels produces visible banding artifacts in the images

� Images are laterally interpolated before they are displayed

to obtain a uniform pixel width in all dimensions

• a relatively simple one-dimensional interpolation, e.g. a cubic

spline, yields a large improvement in the appearance of the image


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PROCESSING 13.3

Interpolation – Phased array (1 of 2)

� Sector images are produced by a phased array system

• scan lines acquired in a polar format must be displayed on a

rectangular pixel grid

� Adjacent scan lines will pass through the same pixel near

the origin of the sector

• a single gray level can be assigned by averaging the magnitude

signals or use the maximum of the overlapping magnitude samples

� Adjacent scan lines may be separated by greater than the

pixel width near the base of the sector

• gaps between scan lines at deeper ranges are typically filled by

two-dimensional interpolation within the image plane


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PROCESSING 13.3

Interpolation – Phased array (2 of 2)

� The interpolation process also compensates for the fact

that locations of most of the samples in the original echo

data do not correspond exactly to centre of any pixel in the

rectangular display grid


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13.4 MODERN IMAGING

METHODS13.4


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13.4 MODERN IMAGING METHODS13.4

Modern Imaging Methods

� 13.4.1 Contrast-Enhanced Imaging

� 13.4.2 Tissue Harmonic Imaging

� 13.4.3 Coded Excitation Imaging

� 13.4.4 Three- and Four-Dimensional Imaging


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13.4 MODERN IMAGING

METHODS13.4.1 CONTRAST-ENHANCED IMAGING


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13.4 MODERN IMAGING METHODS13.4.1 Contrast-Enhanced Imaging

Contrast-Enhanced Imaging

� Microbubbles are used as blood-pool contrast agents in

diagnostic ultrasound

• gas-filled

• encapsulated

• sizes typically ranging from 1 to 4 mm diameter

� A microbubble scatters ultrasound strongly despite its

small size

• because of the over three orders of magnitude difference between

the acoustic impedances of a gas and soft tissue


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Size of microbubbles

� 1 -4 mm diameter microbubbles are smaller than red blood

cells

• so they are not trapped in capillary beds but are still too large to

extravasate

• their fundamental resonant frequency is in the 2-4 MHz range,

which is convenient for ultrasound imaging


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Encapsulation of microbubbles

� Microbubbles are encapsulated by a shell to prevent the

gas from dissolving into the blood

• the original “first-generation” microbubble formulation consisted of

air bubbles encapsulated in a relatively stiff shell material such as

albumin

• most contrast-enhanced exams are now performed with “second-

generation” microbubbles consisting of an inert perfluorocarbon

gas core, such as perfluoropropane (C3F8), encapsulated in a

phospholipid shell

� Circulating half-life of second-generation agents increased

• perfluorocarbon gas is insoluble in blood

• more deformable shell


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Harmonics and microbubbles

� The key to forming contrast-enhanced ultrasound images

is to differentiate echoes produced by microbubbles from

echoes produced by tissue

• exploit the nonlinear scattering characteristics of microbubbles

� A microbubble’s spherical symmetry yields strong

scattering resonances

• at a fundamental frequency determined by the bubble’s radius and

shell properties

• resonances at harmonics (integer multiples) and subharmonics

(integer fractions) of the fundamental frequency

� So, seek echoes with harmonic spectral characteristics


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Non-linear ultrasound propagation

� During the initial development stages of contrast-

enhanced ultrasound, it was assumed that echo signals

containing harmonic spectral characteristics would

uniquely identify echoes from microbubbles

� Soft tissue was unexpectedly discovered to possess

nonlinear ultrasonic properties of its own

� This led to the development of tissue harmonic imaging

(section 13.4.2) and also motivated further research into

contrast-agent detection techniques


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Pulse Inversion (1 of 2)

� Contrast-enhanced imaging may use a variation on a

multi-pulse imaging method such as pulse inversion

� Contrast-enhanced images produced in this fashion reveal

abnormalities in vascular function via regional differences

in the timing and spatial patterns of contrast enhancement

� Each scan line is acquired twice in close succession

• first by transmitting a standard pulse

• then by transmitting an inverted pulse (e.g. the original pulse

multiplied by –1)

� Resulting echo signals are added


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Pulse Inversion (2 of 2)

� Tissue component of the second echo signal is

approximately an inverted copy of the first echo signal

• the echoes from tissue cancel when the two received signals are

summed

� Harmonic components of the microbubble echoes are not

inverted in the second echo signal

� The summation step

• cancels the fundamental frequency components of the microbubble

echoes

• maintains the harmonic components


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Destruction-replenishment imaging (1 of 2)

� Another important approach to contrast-enhanced imaging

� A sequence of high-mechanical-index (MI) transmit pulses

is used to fragment all microbubbles in the field of view

� Contrast enhancement kinetics are measured as new

microbubbles flow into the region of interest

• contrast replenishment can be imaged using a low MI technique,

such as pulse inversion, that does not destroy the agents

� Biophysical models used to estimate parameters such as

• blood volume

• transit time through the region of interest

• perfusion


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Destruction-replenishment imaging (2 of 2)

� Destruction-replenishment imaging is arguably the

contrast-enhanced ultrasound method with the greatest

biomedical value because the functional parameters

estimated with this technique are most similar to the

perfusion parameters measured using modalities such as

dynamic contrast-enhanced CT or MRI


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13.4 MODERN IMAGING

METHODS13.4.2 TISSUE HARMONIC IMAGING


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13.4 MODERN IMAGING METHODS13.4.2 Tissue Harmonic Imaging

Tissue harmonic imaging or native tissue harmonic imaging

� Harmonic spectral components are generated as an

ultrasound pulse propagates through tissue

• from modulation of the sound speed of the tissue by pressure wave

� Substantial harmonics are only produced at the transmit

focus of a diagnostic imaging system, where the pulse

intensity is greatest

� The received signal can be bandpass filtered to construct

an image showing only the harmonic component of the

echoes

• need sufficiently high bandwidth transducer to detect echoes at the

second harmonic of the transmit frequency


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Tissue harmonic imaging – resolution and penetration

� Partially offsets the trade-off between spatial resolution

and penetration depth encountered in conventional B-

mode imaging

• lower frequency transmit pulse experiences less severe frequency-

dependent attenuation than the returning second-harmonic echoes

• the lateral width of the receiver focus is narrower than the transmit

beam.

� The resulting image exhibits lateral resolution and

penetration intermediate between those observed in

• conventional imaging at the fundamental frequency and

• conventional imaging at twice that frequency


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Tissue harmonic imaging – pros and cons

� Tissue harmonic imaging has also been shown

experimentally to reduce clutter artifacts compared to

conventional B-mode imaging

� Benefits are achieved at the expense of reduced axial

resolution

• the passband of the transducer’s frequency response must be

divided into fundamental and harmonic segments rather than using

the entire bandwidth to generate the shortest possible transmit

pulse


IAEA

13.4 MODERN IMAGING

METHODS13.4.3 CODED EXCITATION IMAGING


IAEA

13.4 MODERN IMAGING METHODS13.4.3 Coded Excitation Imaging

Coded Excitation Imaging

� Technique was developed to increase the penetration

depth of ultrasound imaging systems

� Also enables high-frame-rate (e.g., several hundred

frames per second) imaging

� Transmit a relatively long duration signal such as

• a chirp (a sinusoid with increasing or decreasing instantaneous

frequency)

• a pulse modulated code, in which a sinusoid is switched on and off

in a specific temporal sequence to create a binary code


IAEA


Coded Excitation Imaging and penetration depth

� Use of a pulse code spreads the transmitted energy over

the pulse duration

� The signal-to-noise ratio and hence the penetration depth

can be increased without exceeding patient exposure

regulatory requirements

� Matched filtering techniques are used during reception to

deconvolve the transmitted code from the echo signal

� Axial resolution is close to that achieved by conventional

B-mode imaging


IAEA


Coded Excitation Imaging – high-frame-rate

� Simultaneously transmit multiple orthogonal pulse codes

steered to different azimuth angles

� The echoes produced by each of these transmit beams will

interfere at the receive aperture

� Matched filtering by the receiver discriminates among the

echoes produced by each pulse code, enabling multiple

scan lines to be acquired simultaneously


IAEA

13.4 MODERN IMAGING

METHODS13.4.4 THREE- AND FOUR-DIMENSIONAL IMAGING


IAEA

13.4 MODERN IMAGING METHODS13.4.4 Three- and Four-Dimensional Imaging

Three- and Four-Dimensional Imaging

� Advances in ultrasound technology have enabled

conventional 2D B-mode imaging to be supplemented, and

in some applications supplanted, by

• 3D (or volumetric) imaging

• 4D imaging, in which a temporal sequence of 3D images is

presented as a ciné loop

� There are three primary methods of acquiring three-

dimensional images that present differing combinations of

advantages, disadvantages, and technical complexity

• mechanical scanning, freehand scanning and real-time imaging


IAEA


Mechanical scanning

� Mechanically translate or rotate a linear or curvilinear array

transducer with the motion directed out of the 2D image

plane

� A sequence of 2D images is acquired at regular linear or

rotational increments

� A 3D image volume can readily be reconstructed from the

set of 2D images because the spatial relationship among

the images is known in advance


IAEA


Mechanical scanning – pros and cons

� Relatively uncomplicated to implement

� For stationary structures, provides the highest image

quality

• large-aperture or higher frequency linear arrays can be used

• the transducer motion is not operator dependent

� One 3D image volume can require several seconds to

acquire

• the translation or rotation speed of the transducer should be slow

compared to the 2-D frame rate

• extensive respiratory and cardiac gating required if it is used to

image dynamic structures such as the beating heart


IAEA


Freehand scanning

� A conventional linear or curvilinear array is translated or

rotated manually by the sonographer

� Transducer motion

• measured by an external electromagnetic or optical tracking

system that is synchronized with the 2D image acquisition, and the

transducer position measurements are used to position each 2D

image plane within the reconstructed 3D volume, or

• estimated directly from the image sequence using cross-correlation

to estimate motion within the 2D plane and using the rate of

decorrelation of consecutive frames to estimate out-of-plane

motion


IAEA


Freehand scanning – pros and cons

� The quality of the reconstructed image will depend on the

ability of the sonographer to sweep the transducer in a

regular, smooth pattern so that the 3D image volume is

uniformly filled in by the constituent 2D images


IAEA


Real-time imaging

� The most technologically sophisticated method of 3D

imaging

� Two approaches

• 2D phased array or

• 2D array of slightly larger elements


IAEA


Real-time imaging with 2D phased array

� Phased array consists of a matrix of small (ideally < λ/2 on

each side) square elements

� This enables phased array beamsteering in both the

azimuth and elevation angles

� The resulting image volume is a 3D pyramidal sector that

can be considered an extension of the circular sector

produced by a conventional 1D phased array


IAEA


Real-time imaging with 2D array

� 2D array of slightly larger elements than in phased array

� Acquire images in a manner comparable to a conventional

linear array system

� Each scan line is acquired using a 2D subaperture of

active elements that are focused straight ahead

� The subaperture is stepped in both dimensions across the

face of the array to build up a 3D rectilinear image volume


IAEA

13.5 COLOUR FLOW IMAGING 13.5


IAEA

13.5 COLOUR FLOW IMAGING 13.5

Colour Flow Imaging

� Doppler-based methods for blood-flow imaging can be

viewed as extensions of the pulsed-wave Doppler method

in which Doppler processing is applied to a large number

of sample volumes to produce 2D or 3D images of blood

flow

� 13.5.1 Flow Imaging Modes

• Colour Doppler

• Power Doppler

• Duplex / Triplex Doppler

� 13.5.2 Tissue Doppler Imaging


IAEA

13.5 COLOUR FLOW IMAGING 13.5.1 FLOW IMAGING MODES


IAEA

13.5 COLOUR FLOW IMAGING 13.5.1 Flow Imaging Modes

Colour Doppler display

� One of two principal formats for displaying the resulting

flow images

� Colour Doppler uses a red and blue colour scale to

represent the mean axial velocity in each sample volume

• flow towards the transducer shown in red

• flow away from the transducer shown in blue

• the velocity magnitude mapped to the colour intensity

� The colour pixels are superimposed on a B-mode image

such that the echo magnitude from tissue volumes

containing no detectable flow is displayed in gray scale


IAEA


Colour Doppler

� Used when imaging the heart and major blood vessels

• in applications for which mean flow velocity is a diagnostically

useful parameter

� Acquisition of the B-mode image must be interleaved with

acquisition of the Doppler data

• so the colour Doppler frame rate is always less than the frame rate

of conventional B-mode imaging

• and the velocity estimates in a colour Doppler image are prone to

aliasing because the Doppler pulse repetition frequency is limited

by the need to also acquire B-mode scan lines

• the spectral wrap-around produced by aliasing appears as sudden

changes in pixel colour, and thus the apparent direction of flow


IAEA


Power Doppler display

� The second of the two principal formats for displaying the

resulting flow images

� Power Doppler uses a red-to-orange-to-yellow colour

scale to represent the total power in the Doppler spectrum

at each sample volume

• the lowest powers are displayed in red

• the highest powers are displayed in yellow

� As in colour Doppler, the B-mode signal is displayed in

gray scale for tissue volumes with no detectable flow


IAEA


Power Doppler

� Doppler power is in theory proportional to the

concentration of moving blood cells in the sample volume

� Power Doppler is typically used for applications where

blood volume is a diagnostically useful parameter

• such as tumour imaging


IAEA


Power Doppler – pros and cons

� Power Doppler images contain no information about flow

direction or velocity, but

� Doppler power does not depend on the Doppler angle

• so power Doppler provides a more continuous display of tortuous

vessels than colour Doppler

� Aliasing does not affect the total power in the Doppler

spectrum

• so power Doppler images are not susceptible to aliasing artifacts

� Integrated power is less affected by low SNR than is the

mean Doppler frequency

• so power Doppler better for imaging small, slow-flow vessels


IAEA


Triplex Doppler

� Colour Doppler and pulsed-wave Doppler are sometimes

combined in a mode known as triplex Doppler

• the B-mode information included in the colour Doppler image

represents the third component of the triplex

� In triplex Doppler, acquisition of a pulsed-wave Doppler

spectrum from a user selected sample volume is

interleaved with acquisition of a colour Doppler image

� The colour flow image and the Doppler spectrum are

displayed side-by-side


IAEA


Triplex Doppler – the advantage

� This mode provides both

• the 2D spatial information of a colour flow image and,

• for the pulsed-wave sample volume, the higher maximum velocity

and velocity resolution of a conventional pulsed Doppler exam

� Triplex Doppler can be considered an extension of duplex

Doppler, a scanning mode that predates colour flow

imaging in which a B-mode image and a pulsed Doppler

spectrum are acquired and displayed simultaneously


IAEA

13.5 COLOUR FLOW IMAGING 13.5.2 TISSUE DOPPLER IMAGING


IAEA

13.5 COLOUR FLOW IMAGING 13.5.2 Tissue Doppler Imaging

Tissue Doppler Imaging – low-velocity motion

� Conventional Doppler systems discriminate blood flow

from soft tissue motion based on velocity

• including continuous-wave Doppler, pulsed-wave Doppler and

colour flow imaging systems

• a Doppler system assumes blood flow is concentrated at

intermediate and high velocities while scatterers moving at low

velocities correspond to soft tissue

• discrimination is achieved by applying a high-pass wall filter, also

known as a clutter filter, to eliminate low-Doppler-frequency

components from the Doppler signal

� In Tissue Doppler, the wall filter is replaced with a low-

pass filter so only low-velocity motion is displayed


IAEA

13.5 COLOUR FLOW IMAGING 13.5.2 Tissue Doppler Imaging

Tissue Doppler Imaging

� Useful in applications that require measurements of soft

tissue motion

• diagnosis of regional abnormalities in ventricular wall motion, e.g.

following a myocardial infarction or in heart failure patients

� Tissue Doppler can be performed at a single sample

volume in a manner analogous to pulsed-wave Doppler

• the Doppler spectrogram attributed to tissue motion is displayed

� Or, tissue Doppler measurements can be performed over

a 2D region of interest in a manner analogous to colour

Doppler

• to produce an image of mean Doppler frequency or axial velocity


IAEA

13.6 IMAGE ARTIFACTS AND

QUALITY ASSURANCE 13.6


IAEA

13.6 IMAGE ARTIFACTS AND QUALITY ASSURANCE 13.6

IMAGE ARTIFACTS AND QUALITY ASSURANCE

� 13.6.1 B-mode Image Artifacts

� 13.6.2 Speckle

� 13.6.3 Quality Assurance Phantoms and Methods


IAEA


QUALITY ASSURANCE 13.6.1 B-MODE IMAGE ARTIFACTS


IAEA




� 13.6.2 Speckle



IAEA

13.6 IMAGE ARTIFACTS AND QUALITY ASSURANCE 13.6.1 B-Mode Image Artifacts

B-Mode Image Artifacts

Many common B-mode image artifacts can be understood in

terms of fundamental physical concepts

� Reverberation artifact (comet-tail artifact)

� Refraction artifact

� Mirror image artifact

� Shadowing and enhancement artifacts

� Sound-speed artifact

� Side-lobe artifact

� Grating-lobe artifact


IAEA


Reverberation artifact

� Transmitted pulse bounces back-and-forth between two

strongly reflecting interfaces

• each time the reverberating pulse returns to the top interface, a

portion of the acoustic intensity is transmitted through the interface

and continues to the transducer

• these pulses appear in the image as copies of the deeper

boundary separated by a distance equal to the thickness of the

object causing the reverberation

� The image intensity of reverberations decays as a function

of depth

• due to the cumulative multiplication of the pulse intensity by the

reflection and transmission coefficients at the object’s boundaries


IAEA


Reverberation artifact in practice

� Reduction of image intensity with depth means that the

artifact usually obscures only a limited region below the

reverberating object

� Reverberation artifacts are often produced by large

calcifications or metallic foreign bodies

� Sonographers describe the sequence of bright echoes

below a reverberating object as a comet-tail artifact and

use the artifact as an aid to indentifying such hard

inclusions


IAEA


Refraction artifact

� Produced by transmission through specular interfaces

• refraction occurs when a pulse is obliquely incident on a boundary

between two tissues with differing sound speed

• the change in propagation direction of the pulse at the boundary

deflects the pulse away from the intended direction of the scan line

• if the refracted pulse subsequently encounters a strong reflector,

the resulting echo can be refracted back toward the transducer at

the overlying specular boundary

� The refracted echo will be displayed at a point along the

intended direction of the scan line

• the reflecting object will be displayed at an incorrect lateral position

in the image


IAEA


Mirror image artifact

� Produced at a very strongly reflecting (i.e., R→ 1)

specular interface

� Consider a transmit pulse that is redirected by a specular

reflector as described by Snell’s law

• echoes from the redirected pulse are scattered back to the

specular reflector, which reflects them back to the transducer

� Those echoes will be displayed

• along the direction of the original scan line

• behind the specular reflector due to their longer round-trip travel

time


IAEA


Mirror image artifact in practice

� The most striking mirror image artifacts are produced by

the diaphragm when imaging the liver with the scan plane

oriented cranially

� The redirected echoes will cause features to appear in the

image above the diaphragm in the lungs

• at locations from which no signal would be expected due to the

difficulty of coupling ultrasound into an air-filled organ


IAEA


Shadowing and enhancement artifacts (1 of 2)

� Produced by localized variations in the attenuation

coefficient

� If the transmitted pulse traverses a feature that attenuates

more strongly than the surrounding tissue

• pulses incident on features located behind that feature will possess

a lower intensity than expected

• the tissue behind the strongly attenuating feature will thus appear

darker than the laterally adjacent tissue in the B-mode image

• this is shadowing


IAEA


Shadowing and enhancement artifacts (2 of 2)

� Produced by localized variations in the attenuation

coefficient

� If the transmitted pulse traverses a feature that attenuates

less than the surrounding tissue

• pulses incident on features located behind that feature will possess

a higher intensity than expected

• the tissue behind the weakly attenuating feature will thus appear

brighter than expected in the B-mode image

• this is enhancement


IAEA


Shadowing and enhancement artifacts in practice

� Neither time-gain compensation nor lateral gain

compensation is effective for eliminating shadowing or

enhancement artifacts

• time-gain compensation curve is the same along all scan lines

• lateral gain compensation is the same at all depths in the image

� In cancer imaging applications, the relative attenuation in a

lesion can be correlated with whether the lesion is benign

or malignant, so sonographers sometimes consider

shadowing or enhancement artifacts to be diagnostically

useful observations


IAEA


Sound-speed artifacts

� Localized variations in the speed of sound can cause

reflectors to be displayed at incorrect depths in the image

� A region with elevated sound speed

• will cause echoes backscattered from behind that inclusion to

arrive at the receiver sooner than echoes from the same depth

along other scan lines

• scatterers behind the high-sound-speed region will appear to be

closer to the transducer than their true position


IAEA


Sound-speed artifacts in practice

� The sound speeds of most soft tissues are clustered near

1540 m/s

• geometric distortion of the image due to sound speed variations is

rarely perceptible

� Sound-speed artifacts are a cause for concern when

ultrasound is employed in image-guided interventional

procedures


IAEA


Side-lobe artifacts

� For a small anechoic feature such as a fluid-filled cyst,

• when acquiring scan lines passing through the cyst, the side lobes

of the beam may extend outside the cyst into surrounding tissue

• a portion of the intensity contained in the side lobes can be

backscattered toward the transducer

• echoes received as a result of scattering from the side lobes will be

displayed along the beam axis of the scan line, i.e., within the cyst

� The cyst thus appears to be a weakly scattering lesion

rather than an anechoic lesion

� Its image contrast is lower than it would be in the absence

of the side-lobe artifact


IAEA


Side-lobe artifacts and contrast resolution

� Side lobes reduce image contrast

� This reasoning is the basis for interpreting the peak side-

lobe level

as an approximate measure of the contrast resolution of

an ultrasound system


( )( )( )( )

=

mainlobe within max


xPSF

xPSFPSL

IAEA


Grating-lobe artifacts (1 of 2)

� Grating lobes, if present, can produce B-mode artifacts in

a manner analogous to a side-lobe artifact

• the main lobe of the beam is propagating through tissue with

moderate scattering strength

• while a grating lobe is incident upon a strongly reflecting feature

� The echo produced by the grating lobe will be displayed

along the axis of the scan line

� Most array systems are designed to prevent the formation

of grating lobes, so grating-lobe artifacts are relatively

uncommon


IAEA


Grating-lobe artifacts (2 of 2)

� The grating lobe pattern is usually symmetric about the

main lobe axis, so grating-lobe artifacts tend to be visually

symmetric as well

• the strong reflector will be displayed brightly in the image at its

correct position

• less intense copies of the reflector will be displayed at equal lateral

distances to the left and right of the reflector

• the artifactual depictions of the reflector occur on scan lines in

which the m = 1 and m = −1 grating lobes are directed toward the

reflector.


≈ −

s

mm

λθ 1sin

IAEA


QUALITY ASSURANCE 13.6.2 SPECKLE


IAEA




� 13.6.2 Speckle



IAEA

13.6 IMAGE ARTIFACTS AND QUALITY ASSURANCE 13.6.2 Speckle

Speckle

� The granular or mottled texture observed in B-mode

images

� Sometimes considered another type of artifact, but speckle

is also a fundamental characteristic of a B-mode image

and therefore merits special consideration

� The formation of speckle can be understood using a

random-walk model that was originally developed in optics


IAEA


Random-walk model (1 of 4)

� At any instant in time, the ultrasound system receives

echoes from multiple unresolved tissue structures located

in a resolution volume defined by the three-dimensional

point-spread function

� In the narrowband limit, each of these echoes can be

represented by a phasor with a distinct magnitude and

phase

� The instantaneous value of the received RF signal is equal

to the coherent sum of those phasors


IAEA



� If all of the scatterers are similar structures

• the magnitude of each phasor can be modelled as a Gaussian

random variable

� If the scatterers are positioned randomly throughout the

resolution volume

• the phase of the individual echoes can be modelled as a uniformly

distributed random variable from –π to π

� The coherent summation can be visualized graphically as

a random walk in the complex plane


IAEA


� The coherent summation can

be visualized graphically as a

random walk in the complex

plane

� The thin black phasors

represent echoes from

individual scatterers

� The thick gray phasor is the

coherent sum, s, of those

echoes.



IAEA


� If there are at least 10

scatterers in the resolution

volume, the magnitude of the

phasor sum, which

corresponds to the envelope-

detected echo signal, follows

the Rayleigh probability

density function

� The Rayleigh scale parameter

equals one in this example,

so the mean value of |s| is

Rayleigh probability density

function for the magnitude of a

coherently summed speckle

signal, |s|



25.12 ≈π

IAEA


Speckle treated as noise

� Speckle in a B-mode image is thus a random process that

is sometimes compared to noise

� If all of the forgoing assumptions are satisfied

• the ratio of the mean to the standard deviation of the envelope-

detected signal, which is termed the point signal-to-noise ratio, is a

constant ≈ 1.91

• the histogram of the signal provides little information about the

tissue beyond its mean scattering strength

• the size of the individual speckle grains, i.e., the spatial

autocorrelation length of the speckle pattern, is entirely determined

by the point-spread function and so also carries little information

about the tissue

• . Diagnostic Radiology Physics: A Handbook for Teachers and Students – 13.6.2 Slide 8 (131/148)

IAEA


Speckle treated as noise: speckle reduction

� In applications such as lesion detection, the gray-level

fluctuations due to speckle can obscure low-contrast

lesions

� Considerable effort has been devoted to developing

speckle reduction methods

� Spatial compounding, in which several images of a region

of interest are acquired from different angles and

averaged, is the most successful speckle reduction

method and is now implemented under various trade

names on most modern scanners


IAEA


Speckle treated as signal (1 of 2)

� The comparison of speckle to noise can be misleading

� Speckle is the coherent sum of all of the echoes scattered

from the interior of a tissue structure, so the speckle is the

signal in most of the image

� If the transducer and the tissue are both stationary, the

speckle pattern, in contrast to any electronic noise in the

image, is constant

� Importantly, the conditions concerning the scatterers that

allow it to be treated as noise are met only in simulated

images and phantoms


IAEA


Speckle treated as signal (2 of 2)

� All tissues possess some degree of spatial organization to

be able to perform their biological functions

• the condition of randomly positioned scatterers is strictly met only

in simulated images and tissue-mimicking phantoms

� There may also be

• fewer than scatterers per resolution volume

• two or more populations of scatterers in real tissue

� The first- and second-order statistics and spectral

characteristics of echoes acquired from tissue carry more

information about the tissue than the random-walk model

suggests


IAEA


Speckle for tissue characterisation

� The observations concerning tissue characteristics are the

motivation behind ongoing efforts to develop quantitative

tissue characterization methods that employ ultrasound

imaging


IAEA


QUALITY ASSURANCE 13.6.3 QUALITY ASSURANCE PHANTOMS AND

METHODS


IAEA




� 13.6.2 Speckle



IAEA

13.6 IMAGE ARTIFACTS AND QUALITY ASSURANCE 13.6.3 Quality Assurance Phantoms and Methods

Quality Assurance Phantoms and Methods

� Tissue-mimicking phantoms

� Spatial resolution phantoms

� Contrast resolution phantoms

� Flow phantoms

� Training and testing phantoms


IAEA


Tissue-mimicking phantoms (1 of 2)

� Consist principally of a material designed to reproduce the

bulk acoustic properties of soft tissue

• simple laboratory phantoms can be made from suspensions of

gelatine or agar

• commercially available phantoms use proprietary polymer

materials with greater shelf life than gelatine- or agar-based

materials

� The sound speed and attenuation coefficients of a

phantom is usually carefully controlled

• 1540 ms-1

• 0.5 or 0.75 dB/(cm×MHz)


IAEA


Tissue-mimicking phantoms (2 of 2)

� The phantoms also contain suspensions of small

scatterers to produce visually realistic speckle when the

phantom is imaged

� Fabricated from materials such as graphite, polystyrene, or

collagen


IAEA


Spatial resolution phantoms (1 of 2)

� Background material plus

� Spatial resolution targets

• usually small-diameter metal wires or monofilament fibres

suspended horizontally through the phantom

• produce bright point-like targets when imaged in cross section

� A set of wire targets is positioned at the same depth below

the phantom’s acoustic window

• in a regular pattern

• each pair of adjacent wires presents a progressively decreasing

separation from several millimetres down to 0.5 mm or less


IAEA


Spatial resolution phantoms (2 of 2)

� When the pattern of targets is imaged

• lateral spatial resolution can be estimated by observing which pairs

of targets are resolved in the image

• additional wire targets are oriented in a vertical pattern to enable

evaluation of the axial spatial resolution

� A phantom may include several laterally and axially

oriented groups of wires at different depths to enable

resolution to be assessed throughout the scanner’s field of

view


IAEA


Contrast resolution phantoms (1 of 2)

� Contrast resolution targets are usually

• spherical inclusions

• several millimetres to several centimetres in diameter

• similar to the background material but with a mean backscattering

coefficient greater than or less than the background material

� The targets are designed to produce a specific backscatter

contrast relative to the background material

� The contrast of the inclusions can be varied by changing

• size of scattering particles

• composition of scattering particles

• concentration of scattering particles


IAEA


Contrast resolution phantoms (2 of 2)

� Usually easiest to manipulate contrast by changing the

concentration of scattering particles relative to the

background material

• including a few targets with no scatterers to mimic anechoic lesions

� Targets of differing depth will be distributed throughout the

phantom to enable image contrast to be evaluated as a

function of lesion size and depth


IAEA


Flow phantoms

� Provide a useful means of evaluating the performance of

spectral, colour, and power Doppler systems

� In a typical flow phantom design

• tissue-mimicking material is moulded to include one or more hallow

channels that mimic blood vessels of varying diameter and/or

orientation

• channels are connected via tubing to a calibrated pump that is

used to pump blood-mimicking fluid through the flow channels at

controlled flow rates

� Blood-mimicking fluid, is a suspension of small scatterers

designed to reproduce the acoustic properties of blood


IAEA


Training and testing phantoms

� Phantoms are also available for

• training sonographers for specific applications

• testing emerging ultrasound imaging methods

� Examples of training phantoms include anthropomorphic

breast or prostate phantoms designed for practicing

ultrasound-guided biopsy or brachytherapy procedures

� An example of a phantom intended for testing emerging

imaging methods is the elastography phantom

• similar in design to contrast phantoms, except the emulated lesions

differ from the background material in elastic modulus as well as

backscattering coefficient


IAEA

13. BIBLIOGRAPHY13.

Diagnostic Radiology Physics: A Handbook for Teachers and Students – 13. Bibliography Slide 1 (147/148)

IAEA Diagnostic Radiology Physics: A Handbook for Teachers and Students – 13. Bibliography Slide 2 (148/148)

13. BIBLIOGRAPHY13.

� EVANS, D.H., McDICKEN, W.N., Doppler Ultrasound: Physics,

Instrumentation and Signal Processing, Wiley, New York (2000)

� JENSEN, J.A., Estimation of Blood Velocities Using Ultrasound: A Signal

Processing Approach, Cambridge, UK: Cambridge University Press (1996)

� KREMKAU, F.W., Diagnostic Ultrasound: Principles and Instruments, 7th edn,

Saunders/ Elsevier, St. Louis, MO (2006)

� QIN, S., CASKEY, C.F., FERRARA, K.W., Ultrasound Contrast Microbubbles

in Imaging and Therapy: Physical Principles and Engineering, Phys. Med. Biol.

54:R27-R57 (2009)

� SHUNG, K.K., Diagnostic Ultrasound: Imaging and Blood Flow

Measurements, CRC Press, Boca Raton, FL (2006)

� SZABO, T.L., Diagnostic Ultrasound Imaging: Inside Out, Elsevier Academic

Press, Boston (2004)

� ZAGZEBSKI, J.A., Essentials of Ultrasound Physics, Mosby, St. Louis, MO

(1996)

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