Ultrasound

Ultrasound - Chapter 16 (Lecture 1) 9 February 2006

© UW and Brent K Stewart, PhD, DABMP 1

© UW and Brent K. Stewart PhD, DABMP 1

Ultrasound

Brent K. Stewart, PhD, DABMPProfessor, Radiology and Medical Education

Director, Diagnostic Physics

a copy of this lecture may be found at:http://courses.washington.edu/radxphys/PhysicsCourse05-06.html

Including slides from :

© UW and Brent K. Stewart PhD, DABMP 2

PROPERTIES AND GENERATION OF ULTRASOUND

THE GENERATION OF A DIAGNOSTIC ULTRASOUND BEAM

INTERACTIONS OF ULTRASOUND WITH THE BODY

© UW and Brent K. Stewart PhD, DABMP3

Generation of ultrasound

Areas of compression and rarefaction created


Generation of ultrasound

� Wavelength ( λλλλ) distance between consecutive crests (compression) or troughs (rarefaction)

� US: frequency > 20 kHz� Imaging US:

f = 1 – 20 MHz� f depends on source


Characteristics of sound

� Sound (X-ray) beam transfers energy� X-ray beam passes through vacuum� Sound beam requires medium

� Sound waves – longitudinal Particles move parallel to propagation direction


Ultrasound velocity

� velocity of sound in human tissue- independent of frequency- depends on physical constitution of medium

� Compressibility:- v inversely proportional to compressibility

� Density:- slow progression due to greater inertia

� v = f λλλλ� v is constant in a medium




Sound intensity

� Loudness determined by oscillation length of particles transmitting the wave

� Unit = watt / cm 2 (absolute unit)

� Intensity ↑↑↑↑ as amplitude of oscillation ↑↑↑↑� Intensity level - decibel (relative unit)

II o

LogdB = 10


INTERACTIONS OF ULTRASOUND WITH THE BODY

� Reflection� Refraction� Absorption

� Reflected beam used for image formation


Interactions

Reflectedbeam

Refractedbeam

Incident

beam


Reflection

% reflection at tissue interface depends on- acoustic impedance of tissue- angle of incidence of beam

Acoustic impedance (Z = ρρρρ x v)� v constant for wide range of frequencies� Z also constant

(Reflection at tissue interface determined by impedance difference between tissues)

R = (Z1-Z2)2 / (Z1-Z2)2


Reflection

� Reflection ↑↑↑↑ when Z difference ↑↑↑↑� Big ∆∆∆∆Z between soft tissue and bone� Big ∆∆∆∆Z between soft tissue and air

(Gel coupling with skin imperative)� Angle of incidence θθθθi determines reflection� As θθθθi ↑↑↑↑, amount of reflected sound ↓↓↓↓� θθθθi < critical angle →→→→ entire beam reflected� θθθθi = θθθθr →→→→ θθθθi > 3o reflected beam not detected


Refraction

� At interface the velocity of the US changes� Wavelength also changes →→→→ new direction� Bending of wave →→→→ refraction� θθθθt governed by Snell’s law




Refraction

2

1

sin

sin

v

v

t

i =θθ

θθθθt

θi

v1

v2

Snell’s law


Absorption

� Energy loss due to frictional forces causing heat� Amount of absorption determined by:

- frequency of US beam- viscosity ( ηηηη) of conducting medium- relaxation time of medium


Viscosity

� Increase in viscosity:- particle freedom decreases- internal friction (resistance) increases

� ↑↑↑↑ resistance, ↓↓↓↓ US intensity →→→→ heat� Unimportant in low ηηηη materials →→→→ liquids� More absorption in soft tissue →→→→ ηηηη ↑↑↑↑� High absorption in bone


Relaxation time

� Recovery time (t r) after displacement� Constant t r for specific material� Short t r →→→→ recover before next wave� Long t r →→→→ molecules move in one direction,

compression wave opposite� More energy needed →→→→ heat produced


Frequency

� Absorption αααα frequency� ↑↑↑↑ f 2x →→→→ ↑↑↑↑ absorption 2x →→→→ ½ intensity� Ideal frequency compromise between:

- good resolution- transmission of E to deeper interfaces

� Affects absorption due to viscosity- as f ↑↑↑↑ →→→→ motion slower due to ηηηη

� Affects absorption due to relaxation time- at low f →→→→ molecules can relax between cycles


Summary

� Reflection� acoustic impedance� angle of incidence

� Refraction� Snell’s law

� Absorption� frequency� viscosity� relaxation time




Ultrasound

� Frequency > 20 kHz, generally 1-20 MHz


Ultrasound


Production of piezoelectric crystals

� Dipoles arranged in precise geometric configuration� Crystal heated in strong electric field

(dipoles free to move into desired alignment)� Crystal gradually cooled under high voltage� Room temp. →→→→ dipoles fixed →→→→ piezoelectric properties� Curie temperature →→→→ polarisation lost

(transducers should not be autoclaved)


Transducers


Transducer

� Good contact between crystal and electrodes

� Electrodes produce electric field

� Crystal under stress and deforms

� Backing block absorbs US transmitted to back of transducer


Resonant frequency

� Crystal thickness determines resonant frequency� Thick crystal →→→→ low frequency US� Crystal thickness = ½ λλλλ of US required

(fundamental resonant frequency)

� In sonography crystal subjected to single voltage p ulse →→→→vibrates at natural frequency




Resonance

Fundamental frequency

Forced vibration


16.3 Ultrasonic Transducers (6)

• One characteristic of a transducer is the frequency ‘purity’ and the length of time the resonance persists, or ‘ring down time’

• This characteristic is named the “Q” factor: Q = operating frequency (MHz)/bandwidth = f0/BW

• High Q transducer → long wave train with a narrow frequency range

• Low Q transducer → short wave packet with a wide frequency range

c.f. Bushberg, et al. The Essential Physics of Medical Imaging, 2nd ed., p. 487.


16.3 Ultrasonic Transducers (7)

• A block of damping material is placed behind the PC in the transducer to reduce the ring down time and produces a spatial pulse length (SPL) of about 3λλλλ to achieve reasonable depth (axial) resolution in the final ultrasound image

c.f. Bushberg, et al. The Essential Physics of Medical Imaging, 2nd ed., p. 487. © UW and Brent K. Stewart PhD, DABMP

28

16.3 Ultrasonic Transducers:Broad-bandwidth ‘Multifrequency’ Transducers



Wave pattern of sound beam


Wave pattern of sound beam

� Waves reinforce or cancel each other

� Synchronisation depends on wavelength

� Short λλλλ →→→→ front close to transducer surface

� US beam separated into two components




Ultrasound beam

Fresnel Fraunhofer

Length = d2 / 4λλλλ

d

Only Fresnel zone is used


Fresnel zone (Length = d2 / 4λ)


Fresnel zone(Length = d2 / 4λ)

� Longest →→→→ large transducer, high frequency US

� Shortest →→→→ small transducer, low frequency US


High frequencies

Advantages� Longer Fresnel zone� Better depth resolutionDisadvantages� Tissue absorption (poor penetration)� Low f - penetrates deeperLow frequency, larger transducer, an alternative,

however poor lateral resolution


Comparing beams

(Length = d2 / 4λλλλ)

(low frequency, longer λλλλ)

(high frequency, shorter λλλλ)


Low vs high frequency




16.4 Ultrasound Beam PropertiesNear Field and Far Field

� Near (parallel) Field “Fresnel zone”� Is adjacent to the transducer

face and has a converging beam profile

� Convergence occurs because of multiple constructive and destructive interference patterns of the ultrasound waves (pebble dropped in a quiet pond)

� Near Zone length = d2/4λ = r2/λ



16.4. Ultrasound Beam PropertiesNear Field and Far Field

� The far field or Fraunhofer zone is where the beam diverges� Angle of divergence for

non-focused transducer is given by

� sin(θ) = 1.22λλλλ/d� Less beam divergence

occurs with high-frequency, large-diameter transducers

c.f. Bushberg, et al. The Essential Physics of Medical Imaging, 2nd

ed., p. 491.


16.4. Ultrasound Beam PropertiesNear Field and Far Field

� A focused single element transducer uses either a curved element or an acoustic lens:� Reduce beam width� All diagnostic transducers are

focused� Focal zone is the region over

which the beam is focused� A focal zone describes the

region of best lateral resolution

� Near Zone length = d2/4λ = r2/λd=transducer diameter, r=radius


ed., p. 492. 1/23/2014 © UW and Brent K. Stewart PhD, DABMP

Image formation with ultrasound


A-mode data display

Depth Time


Gray scale imaging

Display the amplitude of echo in intensities of gray (television scan converter)

� Analog� gray scale drift - poor image quality� unacceptable flicker

� Digital� 16, 32 or more gray scales

� no drift� rapid read and write (no flicker)




B-mode data display

Position Position


M - mode data display


M mode display


B Mode display


Real time imaging


Real time imaging

Observe motion in B-mode

Linear array of transducers




Electronic focus/steering


16.3 Ultrasonic Transducers:Phased Array Transducers

� 64 to 128 elements� All active during imaging� Using time delays can steer and focus beam electronically



16.3 Ultrasonic Transducers:Linear or Curvilinear Array Transducers

� 256 to 512 elements� Simultaneous firing of a small group of approx. 20 elements produces the

ultrasound beam� Rectangular field of view produced for linear and trapezoidal for curvilinear

array transducers

c.f. Bushberg, et al. The Essential Physics of Medical Imaging, 2nd ed., p. 489. © UW and Brent K. Stewart PhD, DABMP

52

Example 1


Example 2


Example 3




16.4. Spatial Resolution

� In ultrasound, the major factor that limits the spatial resolution and visibility of detail is the volume of the acoustic pulse

� The axial, lateral, and elevational (slice thickness) dimensions determine the minimal volume element



16.4. Spatial Resolution - Axial

� Axial resolution (linear, range, longitudinal or depth resolution) is the ability to separate two objects lying along the axis of the beam

� Achieving good axial resolution requires that the returning echoes be distinct without overlap

� The minimal required separation distance between two boundaries is ½ SPL (about ½ λ) to avoid overlap of returning echoes

� SPL = number of cycles emitted per pulse by the transducer x λ

� Objects spaced closer than ½ SPL will not be resolved



16.4. Spatial Resolution - Axial

� Typical axial resolution is 0.5 mm� Higher frequencies reduce SPL,

improving axial resolution however, increases attenuation

� Axial resolution remains constant with depth



16.4. Spatial Resolution - Lateral

� Lateral (azimuthal) resolution - the ability to resolve adjacent objects perpendicular to the beam direction and is determined by the beam width (diameter)

� Typical lateral resolution (unfocused) is 2 - 5 mm, and is depth dependent

� Single focused transducers restrain the beam to within narrow lateral dimensions at a specified depth using lenses at the transducer face



16.4. Spatial Resolution - Slice thickness (Elevational)

� Elevational resolution is dependent on the transducer element height

� Perpendicular to the image plane

� Use of a fixed focal length lens across the entire surface of the array provides improved elevational resolution at the focal distance, however partial volume effects before and after focal zone

c.f. Bushberg, et al. The Essential Physics of Medical Imaging, 2nd ed., p. 497, 500.


16.4. Ultrasound Beam Properties - Side Lobes

� Side lobes are unwanted emissions of ultrasound energy directed away from the main pulse

� Caused by the radial expansion and contraction of the transducer element during thickness contraction and expansion

� Lobes get larger with transducer size� Echoes received from side lobes are mapped into the main beam, causing artifacts





16.4. Ultrasound Beam Properties - Side Lobes

� For multielement arrays, side lobes are forward directed� Grating lobes result when ultrasound energy is emitted far off-axis by multielement

arrays, and are a consequence of the noncontinuous transducer surface of the discrete elements� results in appearance of highly reflective, off-axis objects in the main beam



16.8 Ultrasound Artifacts

� Artifacts arise from the incorrect display of anatomy or noise during imaging

� Refraction causes misplaced anatomic position in the image

� Shadowing and Enhancement� Shadowing is the reduced echo

intensity behind a highly attenuating or reflecting object, such as a stone creating a “shadow”

� Enhancement is the increased echo intensity behind a minimally attenuating object such as a fluid filled cyst




� Reverberation artifacts commonly occurs between two strong reflectors, such as an air pocket and the transducer array at the skin surface� The echoes bounce back and

forth between the two boundaries and produce equally spaced signals of diminishing amplitude in the image

� This is often called a “comet-tail” artifact




� Speed displacement artifacts are caused by the variability of the speed of sound in various tissues� In the case of fatty tissues, the

slower speed of sound in fat (1,450 m/sec) results in a displacement of the returning echoes from distal anatomy by about 6% of the distance traveled through the mass




� Side Lobes Grating Lobes

� Side lobe energy emissions in transducer arrays can cause anatomy outside the main beam to be mapped into the main beam

� For a curved boundary, such as the gallbladder, side lobe interactions can be remapped and produce findings such as “pseudo-sludge” that is not apparent with other scanning angles

� Grating Lobes

� Create ghost images of off-axis high-contrast objects




� A mirror image artifact arises from multiple beam reflections between a mass and a strong reflector, such as a diaphragm� Multiple echoes result in the

creation of a mirror image beyond the diaphragm of the mass

� Speckle is a textured appearance that results from small, closely-spaced structures that are too small to resolve as seen on images of solid organs





Artefacts

� Speckle pattern� Multiple scatter by parenchime (random)

� Reverberation� Multiple reflection between surfaces� Appear as if coming from deeper position

� Shadowing� Causes by high reflection or absorption

� Low intensity


Artefacts

� Enhancement� Less attenuation than adjacent tissue� Streak with increased intensity

� Edge shadows� Curved walls reflect sound

� Lack of echoes beyond curvature


Artefacts


Other images







Take Aways: Five Things You should be able to Describe after Today’s Ultrasound Lecture

� The basic mechanism underlying the formation of ultrasonic images

� The various characteristics of ultrasonic waves� The various types of interactions of ultrasonic waves with

matter � How medical ultrasound transducers generate and

receive ultrasonic waves to form the basis of an image� The various components of a medical ultrasound

transducer and their function

Ultrasound – Chapter 16 Bushberg Diagnostic Imaging Physics Course9 February – 16 March 2006

Kalpana M. Kanal, Ph.D. 1

Ultrasound – Chapter 16 – Lecture 2

Kalpana Kanal, Ph.D., DABRAssistant Professor, Diagnostic Physics

Dept. of RadiologyUW Medicine

a copy of this lecture may be found at:http://courses.washington.edu/radxphys/PhysicsCourse05-06.html

Kanal 2

16.5 Image Data Acquisition


Kanal 3

16.5 Image Data Acquisition

c.f. Bushberg, et al. The Essential Physics of Medical Imaging, 2nd ed., p. 505-507. Kanal 4

16.5 Image Data AcquisitionEcho Display Modes


� A-mode “amplitude” mode: displays echo amplitude vs. time (depth)

� One “A-line” of data per pulse repetition

� A-mode used in ophthalmology or when accurate distance measurements are required

Kanal 5


� B-mode (B for brightness) is the electronic conversion of the A-mode and A-line information into brightness-modulated dots on a display screen

� In general, the brightness of the dot is proportional to the echo signal amplitude


Kanal 6


� M-mode (“motion” mode) or T-M mode (“time-motion” mode): displays time evolution vs. depth

� Sequential US pulse lines are displayed adjacent to each other, allowing visualization of interface motion

� M-mode is valuable for studying rapid movement, such as mitral valve leaflets




Kanal 7

16.5 Image Data AcquisitionScan Converter

� The function of the scan converter is to create 2D images from echo formation received and to perform scan conversion to enable image data to be viewed on video display monitors

� Scan conversion is necessary because the image acquisition and display occur in different formats

� Modern scan converters use digital methods for processing and storing data

� For color display, the bit depth is often as much as 24 bits or 3 bytes of primary color

Kanal 8

16.6 2D Image Display and Storage

� Mechanical Scanning and Real-Time Display� Mechanically driven transducers sweep out sections of tissue

repeatedly, so many times per second� Single wobbling or continuously rotated multi-element

transducers are used


Kanal 9


� Electronic Scanning and Real-Time Display� Sequential linear array: 256 to 512 individual transducer elements

producing a rectangular or convex curvilinear array� Sequential scanning uses transducers pulsed in groups, where each

group sends and receives before next group is pulsed� Number of A-lines is approx. equal to the number of transducer

elements


Kanal 10


� Electronic Scanning and Real-Time Display� Electronic phased array transducer� Sector or linear images; 64,128 or 256 elements� Beam sweep is accomplished through pulsing the individual

transducers with small timing delays between transducer elements


Kanal 11

16.5 Image Data Acquisition Pulse Echo Operation

� Diagnostic ultrasound utilizes a pulse-echo format using a single transducer to generate images

� Most ultrasound beams are emitted in brief pulses (1-2 µµµµs duration)

� For soft tissue (c = 1540 m/s or 0.154 cm/ µsec), a return time of 13 µµµµs corresponds to a depth of 1 cm (round trip = 2 cm)� c = 2D / time� Time (µµµµsec) = 2D (cm) / c (cm/ µµµµsec) � = 13 µµµµsec x D (cm)� Distance (cm) = 0.077 x Time (µµµµsec)

Kanal 12

16.5 Image Data Acquisition Pulse Repetition Frequency (PRF)

� Maximum imaging depth is dependent on attenuation and the pulse repetition frequency (PRF = the number of times the transducer is pulsed/sec)� Increase in PRF causes a decrease in maximum imaging depth� Increase in transducer frequency causes a decrease in imaging depth due

to signal attenuation of the ultrasound pulse




Kanal 13

16.6 2D Image Display and StorageImage Frame Rate

� A 2D image (a single frame) is created from a number of A-lines, N (typically 100 or more), acquired across the FOV

� Line density is the number of vertical lines per FOV� The frame rates (1/acquisition time per frame) for real time imaging

are typically 15-40 frames/second, which permits motion to be followed

Kanal 14

16.6 2D Image Display and StorageImage Frame Rate

Kanal 15

16.6 2D Image Display and StorageImage Storage

� US images are typically composed of 640 x 480 or 512 x 512 pixels� Pixel depth is typically 8 bit or 1 byte, providing 256 levels of gray

scale� Image storage without compression is approximately ¼ megabytes

(MB) per image� For real time imaging (10-30 frames per sec), this can amount to

hundreds of MBs of data� Color images require 24 bits/pixel for storage

Kanal 16

16.7 Clinical Transducers

� Low frequency transducers have better tissue penetration� Transducers used for abdominal imaging are generally in the 2.5 to 5

MHz range� Specialized high-resolution and shallow-penetration probes (8 to 20

MHz) have been developed for studying the eye� In infants, 3.5 to 7 MHz transducers are used for echoencephalography

� Endovaginal transducers – pelvic region and fetus� Endorectal transducers – prostrate, Transesophageal transducers –

heart, Intravascular transducers – blood vessels


Kanal 17

Doppler effect

Kanal 18

Doppler effect



Kanal 19 Kanal 20

Doppler Ultrasound

� The Doppler ultrasound is based on the shift in frequency in an ultrasound wave caused by a moving reflector (siren on a fire truck)� Objects moving toward the

observer (transducer) appear to have a higher frequency and shorter wavelength

� Objects moving away from the observer (transducer) appear to have a lower frequency and longer wavelength

� If object moving perpendicular to the observer (transducer), no change in the observed frequency or wavelength c.f. Bushberg, et al. The Essential Physics of

Medical Imaging, 2nd ed., p. 532.

Kanal 21

Doppler Frequency Shift� The Doppler shift is the difference

between the incident frequency and reflected frequency

� fd = Doppler frequency shift

� fi = transducer frequency� fr = reflected frequency

� v = blood velocity� ct = speed of sound in tissue� As the angle of incidence increases

with respect to the long axis of the blood vessel, the Doppler shift decreases according to the dependence on the cosine of the angle

� Cos 0 = 1, cos 30 = 0.87, cos 45 = 0.707, cos 60 = 0.5, cos 90 = 0

( )2 cosi r id

t

vf f f f

c

θ= − =

( )2 costd

i

f cv

f θ=


Kanal 22

� Frequency shifts are in the audible range� fi = 5 MHz, v = 35 cm/sec, θ = 45o

� fd = 2 (35 cm/sec)(0.707)(5 MHz)/(154,000 cm/sec) = 1.6 kHz

� Human audible spectrum: 15 Hz – 20 kHz� Preferred Doppler angle is from 30-60 degrees� At >60 deg, minor errors in angle accuracy can

result in large errors in velocity� At <20 deg, refraction and aliasing problems can

occur

( )2 cosi r id

t

vf f f f

c

θ= − =

Doppler Frequency Shift


Kanal 23

Continuous-Wave Doppler Operation

� Continuous Wave Doppler: one transducer continuously transmits and one transducer continuously receives

� The frequency of the two signals are subtracted to give the Doppler shift

� Continuous-wave Doppler is inexpensive, does not suffer from aliasing but lacks depth resolution and provides little spatial information� Samples everything along the Doppler line

� Cannot position the Doppler to listen at a specific area along it’s path

� Good for measuring fast flow and assessing deep lying vessels

Kanal 24

Pulsed Doppler Operation

� Pulsed Wave Doppler: allows both velocity and depth information to be obtained (pulse-echo)

� Single transducer used� Doppler information is only provided for a selected area (range gate)

specified by the operator



Kanal 25


� According to sampling theory, a signal can be reconstructed as long as the true frequency (Doppler shift) is less than half the sampling rate

� This the PRF (sampling frequency) must be at least twice the maximal Doppler shift encountered in the measurement

� High blood flow causing the Doppler shift to exceed ½ PRF will result in false (aliased) velocities


Kanal 26


� Thus the maximum velocity which can be displayed is limited

� A 1.6 kHz Doppler shift requires a minimum PRF of 2 x 1.6 = 3.2 kHz

� The most straightforward method to reduce or eliminate aliasing error is for the user to adjust the velocity scale to a wider range as most instruments have the PRF linked to the scale setting

( )max0max

2 cos

2 t

f vPRFf

c

θ∆ = =

( )max04 costv

c PRFf θ

⋅=


Kanal 27

Duplex Scanning

� Combination of 2D B-mode imaging (visual guidance) and pulsed Doppler data acquisition

� The 2D B-mode creates the real-time image to facilitate selecting the Doppler gate window position, and then is switched to the Doppler mode

Figure courtesy: Brent Stewart, Ph.D.

Kanal 28

Duplex Scanning

� Sample volume position (range gate) indicated by a window position cursor and a line cursor for the angle

� Errors in the flow volume may occur� vessel axis might not lie

totally within scanned plane� vessel might be curved


Kanal 29

Laminar Flow

� Layers of flow (normal)� Slowest at vessel wall� Fastest within center of vessel

� Disease states disrupt laminar flow (turbulent flow)

Figure courtesy: Bill Warren, M.D.

Kanal 30

Spectral Waveform

� Doppler produces an audible signal as well as a graphical representation of flow = Spectral Waveform

� The spectral waveform represents the audible signal and provides information about� the direction of the flow� how fast the flow is traveling

(velocity)� the quality of the flow (normal

vs. abnormal)� Plot of Doppler shift frequency

spectrum versus time




Kanal 31

Spectral Waveform

� Flow coming TOWARD the transducer is represented above the baseline

� Flow traveling AWAY from the transducer is represented below the baseline

� The amplitude of the shift frequency is encoded as gray-scale variations

� Two Doppler spectra are shown at 2 discrete points in time with amplitude versus frequency

� A broad spectrum represents turbulent flow while

� A narrow spectrum represents laminar flow within the Doppler gate


Kanal 32

Color Flow Imaging

� Color flow imaging provides a 2D visual display of moving blood in the vessels, superimposed on the conventional gray-scale image

� Blue and Red colors are assigned, depending on motion toward or away from the transducer

� Typically, flow toward the transducer is assigned red and flow away from the transducer blue


Kanal 33

Color Flow Imaging

� Turbulent flow can be displayed as green or yellow

� The color intensity varies with flow intensity

� Color Doppler can detect flow in vessels too small to be seen by imaging alone

� One limitation of color scanning is that clutter of slow-moving solid structures and noise can overwhelm the smaller echoes from moving blood

� Spatial resolution of the color image is much lower than that of gray-scale image


Kanal 34

Color Doppler imaging

Kanal 35

Power Doppler

� Power Doppler permits detection and interpretation of slow blood flow but sacrifices directional and quantitative flow information

� Power Doppler uses the return Doppler signal strength alone� It ignores the direction of frequency shift or phase, as in

conventional color flow imaging� Power Doppler uses the same power levels as those of conventional

color scanning� It is more sensitive than standard color flow imaging

� The image signal does not vary with the direction of flow� Aliasing artifacts do not occur in power Doppler

Kanal 36

Harmonic ImagingOverview

c.f. Ultrasound Technology Update, GE Medical Systems Document, p. 1



Kanal 37

Harmonic ImagingHow Are Harmonics Generated?

� The harmonics are not generated by the ultrasound scanner itself

� These signals are generated in the body as a result of interactions with the tissue or contrast agents

� Interactions with contrast agents� Patient injected with contrast

agents containing very small bubbles

� A conventional ultrasound pulse is sent into the body


Kanal 38


� When the pulse encounters the bubble, it generates two kinds of responses

� First, echo returns from the bubble as in conventional ultrasound and second, the bubble vibrates in response to the shock from the pulse (bell)

� Vibration generates a second harmonic at twice the frequency of the original ultrasound pulse


Kanal 39


� This kind of imaging benefits from the fact that the only the strong signal returning from the body at twice the fundamental frequency will be the signal that comes back from places where the bubbles are

� By listening only for the ring of the bell, the harmonic signal, the ultrasound system can generate very high contrast ultrasound images that are relatively free from the kind of interference that makes conventional ultrasound imaging difficult


Kanal 40


� Interactions with Tissue� When the sound wave passes through the tissue, it compresses and

and expands the tissue� When the tissue is compressed, the speed of sound is higher and

when it is expanded, the speed of sound is lower


Kanal 41


� Interactions with Tissue� Because the speed of sound is higher when the pressure is higher,

the top of the waveform gets pulled forward as the wave passes through tissue

� This distortion of the tissue causes harmonics to be generated� Different tissues distort the wave in different ways (fat distorts more

then muscle, liver or kidney tissue)� The resultant waveform contains both the fundamental frequency

plus the harmonic frequencies caused by the distortion

c.f. Ultrasound Technology Update, GE Medical Systems Document, p. 3 Kanal 42


� Interactions with Tissue� This ability to create

harmonics in tissue is an effect that is seen in varying degrees through out the ultrasound field of view

� The harmonic imaging effect without contrast agents is most pronounced in the mid field (middle of the ultrasound image)




Kanal 43


Kanal 44


Kanal 45

Potential Advantages of the Harmonic Signal

� Harmonic beams are narrower than their conventional counterparts

� Side lobes are lower as well


Kanal 46

Potential Advantages of the Harmonic Signal

� The result is improved lateral spatial resolution and better contrast resolution, removal of multiple reverberation artifacts caused by anatomy adjacent to the transducer

� Furthermore, since the harmonics are generated inside the body, they only have to pass through the fat layer once (on receive), not twice (transmit and receive)


Kanal 47


Kanal 48




Kanal 49

Pulse Inversion Harmonic ImagingContrast Agents

� Improves sensitivity to microbubble contrast agents

� Reduces signal from surrounding soft tissues

� Disadvantage include motion artifacts from moving tissues that occur between pulses and frame rate penalty (at least 2 times slower than a standard scan)


Kanal 50

Bioeffects

� Isppa and Ispta are important parameters when considering the possibility of inducing bioeffects� The intensity at a specific point during a single pulse is the

spatial peak pulse average intensity (Isppa)� The intensity at a specific point averaged over a long period

(many pulses) is the spatial peak temporal average intensity (Ispta)

� At high power levels, ultrasound can cause:� Cavitation - the creation and collapse of microscopic bubbles� Small-scale fluid motions called microstreaming

� Tissue heating occurs as a result of energy absorption and is the basis of using ultrasound for hyperthermia treatment

� No harmful effects have been reported for diagnostic imaging uses of pulsed ultrasound below 100 mW/cm2 (Ispta)

Kanal 51

Bioeffects

Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 553.

Kanal 52

System Performance and Quality Assurance


ed., p. 545.

Precision multi-purpose grey scale phantom (RMI 403GS LE)

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Ultrasound