USG Physics basic

1. Physics of Diagnostic Ultrasound
Presented by:
-Lalit karki,Resident 1st
year
Moderator:
-Dr.Sagar Khadka, Lecturer

2. Session Aims:
• Introduction to sound waves and their characteristics
• Define Ultrasound
• Basic principles of image formation
• Describe the four main types of ultrasound interactions with matter
•Construction and operation of the ultrasound transducer
• Ultrasound instrumentation
• Ultrasound safety

3. Wave Motion
• Waves transfer energy from one location to another
• Waves can be broadly described as either “Transverse” or “Longitudinal”
Sound Waves and Their Characteristics

4. Propagation of Sound
• Sound waves are mechanical pressure waves (longitudinal) which propagate
through a medium by compression and rarefaction of the particles
• As a sound pressure wave propagates through the medium, particles in regions of
high pressure will be pushed together (compression) and particles in regions of
low pressure will be pulled apart (rarefaction)

5. Propagation of Sound
• Rarefaction follows compression as the compressed particles transfer their energy
to adjacent particles
• The compression (and subsequent rarefaction) continues to travel forward through
the medium as the energy is transferred between particles

6. Power and Intensity
• A sound wave transports Energy through a medium from a source. Energy is
measured in joules (J)
• The Power, P, produce by a source of sound is the rate at which it produces energy.
Power is measured in watts (W) where 1 W = 1 J/s
• The Intensity, I, associated with a sound wave is the power per unit area. Intensity
is measured in W/m2
• The power and intensity associated with a wave increase with the pressure
amplitude, p
Intensity, I  p2
Power, P  p

7. Wavelength, Frequency and Speed
• Waves are characterised by their
wavelength, frequency and speed
• The Wavelength,  , is the distance
between consecutive peaks or other
similar points on the wave.
• The Frequency, f, is the number of
oscillations per second
• Frequency is measured in Hertz (Hz)
where 1 Hz is one oscillation per second.

8. Wavelength, Frequency and Speed
• The Speed of sound, c , is the distance travelled by
the wave per unit time and is equal to the
wavelength multiplied by the frequency
• The speed of sound is dependent on the medium
through which it travels and varies greatly in
different materials
• The speed of the wave is determined by the bulk
modulus, B, (measure of stiffness) and the density, ,
(mass per unit volume) of the medium
• Highly compressible media (low B), such as air, has a
low speed of sound – 330 m/s
• Less compressible media, such as bone, has a higher
speed of sound – 4080 m/s
c =  f
c =  B / 

9. Wavelength, Frequency and Speed
Material Density (Kg/m3) c (m/s)
Air 1.2 330
Fat 924 1450
Water 1000 1480
Kidney 1041 1565
“Average Tissue” 1050 1540
Muscle 1068 1600
Bone 1912 4080

10. Wavelength, Frequency and Speed
• The frequency of a sound wave is unaffected by
changes in the speed of the wave as it propagates
through different media
• Therefore, the wavelength changes as the wave
travels through different media
• Wavelength increases with an increase in wave
speed
• Higher frequency sound waves have a shorter
wavelength
 = c / f

11. What is Ultrasound?

12. Ultrasound
• The term “Ultrasound” refers to sound waves of such a high frequency that they
are inaudible to humans
• Ultrasound is defined as sound waves with a frequency above 20 kHz
• Ultrasound frequencies in the range 3-15 MHz are typically used for diagnostic
imaging purposes
• Medical diagnostic ultrasound uses ultrasound waves and the acoustic properties
of the tissues in the body to produce an image

13. Ultrasound
• The use of ultrasound in medicine began shortly after the
2nd World War
• Dr. Karl Theodore Dussik’s work on transmission
ultrasound investigation of the brain in 1942 (Austria) was
the first published work on medical ultrasound
• Ultrasound was first developed for clinical purposes in
1956 in Glasgow
• Obstetrician Ian Donald and engineer Tom Brown
developed the first prototype systems based on an
instrument used to detect industrial flaws in ships
• They perfected its clinical use, and by the end of the
1950s, ultrasound was routinely used in Glasgow
hospitals
• Commercial systems became available in the mid-1960’s

14. Ultrasound

15. Basic Principals of Image Formation

16. Pulse Echo Principal
• A short ultrasound pulse is delivered to the tissues, and where there are changes
in the acoustic properties of the tissue, a fraction of the pulse is reflected (an
echo) an returns to the source (pulse-echo principal)
• Collection of the echoes and analysis of their amplitudes provides information
about the tissues along the path of travel
Tissue 1 Tissue 2 Tissue 3
Transducer
US Pulse
Reflected Echoes

17. Pulse Echo Principal
Tissue 1 Tissue 2 Tissue 3
Transducer
US Pulse
Reflected Echoes
The ultrasound pulse will travel at the speed of sound and the time between
the pulse emission and echo return will be known.
Therefore, the depth, d, at which the echo was generated can be determined
and spatially encoded in the depth direction.
Distance (D) = speed (c) x time (t)  2d = c t

18. Tomographic Imaging
Repeating this process many times with incremental changes in pulse
direction allow a volume to be sampled and a tomographic image to be
formed.
i.e. A tomographic image is formed from a large number of image lines, where
each line in the image is produced by a pulse echo sequence
Transducer

19. B-Mode Image
• B-mode = Brightness mode
•A B-mode image is a cross-sectional image representing tissues and organ boundaries
within the body
• Constructed from echoes which are generated by reflection of US waves at tissue
boundaries, and scattering from small irregularities within tissues
• Each echo is displayed at a point in the image which corresponds to the relative
position of its origin within the body
• The brightness of the image at each point is related to the strength (amplitude) of
the echo

20. B-Mode Image – How Long Does it Take?
1. Minimum time for one line = (2 x depth) / speed of sound = 2D / c seconds
2. Each frame of image contains N lines
3. Time for one frame = 2ND / c seconds
E.g. D = 12 cm, c = 1540 m/s, Frame rate = 20 frames per second
Frame rate = c / 2ND
N = c / (2D x Frame rate) = 320 lines (poor - approx half of standard TV)
Additional interpolated lines are inserted between image lines to boost image quality
to the human eye
4. Time is very important!!!

21. Time Gain Compensation (TGC)
• The deeper the source of echo  Smaller signal intensity
• Due signal attenuation in tissue and reduction of the initial US beam intensity by
reflections
• Operator can TGC use to artificially ‘boost’ the signals from deeper tissues to
compensate for this (like a graphic equaliser)

22. M-Mode Image
• Can be used to observe the motion of tissues (e.g. Echocardiography)
• Image the same position (one image line) repeatly.
• One direction of display is used to represent time rather than space
Transducer at fixed point Time
Depth

23. Basic Principles of Image Formation
M-Mode Image of Mitral Valve

24. Interactions with Matter

25. Interactions with Matter
• Ultrasound interactions with matter are determined by the acoustic properties of
the media through which it propagates
• As Ultrasound energy propagates through a medium, interactions include:
– Reflection
– Refraction
– Scatter
– Attenuation / Absorption

26. Reflection
• Reflection (specular reflection) occurs at
tissue boundaries where there is a
difference in the acoustic impedance, Z, of
the two tissues
• When the incident ultrasound wave is
perpendicular to the boundary, a fraction
of it’s energy is reflected (an echo) directly
back towards the source
• The remaining energy is transmitted into
the second tissue and continues in the
initial direction
Z1
Z2
Incident
Reflectio
n (echo)
Transmission

27. Reflection – Acoustic Impedance
• The acoustic impedance of a material is a measure of the response of the particles
of the medium to a wave of given pressure (e.g. resistance)
• The acoustic impedance of a medium is again determined by the bulk modulus, B,
(measure of stiffness) and the density, , (mass per unit volume) of the medium
• Consider a row of masses (molecules) linked by springs (bonds)
• A sound wave can be propagated along the row of masses by giving the first mass
a momentary “push” to the right
• This movement is coupled to the second mass by the spring
m m m m
B B B
Sound wave

28. Reflection – Acoustic Impedance
• Small masses (m) model a material of low density linked by weak springs of low
stiffness (b)
• A given pressure is applied momentarily to the first small mass m
• The small mass is easily accelerated to the right and its movement encounters little
resistance from the weak spring b
• This material has a low acoustic impedance, as particle movements are relatively
large in response to a given applied pressure
m m m m
b b b
Sound wave

29. Reflection – Acoustic Impedance
• Large masses (M) model a material of high density linked by springs of high
stiffness (B)
• In this case, the larger masses M accelerate less in response to the applied
pressure
• Their movements are further resisted by the stiff springs B
• This material has a high acoustic impedance, as particle movements are relatively
small in response to a given applied pressure
M M M M
B B B
Sound wave

30. Reflection – Acoustic Impedance
The acoustic impedance, Z, of a material is given by
Recall that the speed of sound, c =  B / 
B =  c2
Therefore,
z =   B
 = density (kg/m3)
B = bulk modulus (kg/m-s2)
z =  c
 = density (kg/m3)
c = speed of sound (m/s)

31. Reflection – Acoustic Impedance
Material Acoustic Impedance (Kg/m2s)
Air 0.0004 x 106
Fat 1.34 x 106
Water 1.48 x 106
Kidney 1.63 x 106
Muscle 1.71 x 106
Bone 7.80 x 106
There are
relatively small
differences in
acoustic
impedance for
“soft tissues”

32. Reflection
• The fraction of ultrasound intensity
reflected at an interface is given by the
intensity reflection coefficient, R
• The fraction of ultrasound energy
reflected depends on the difference
between the Z values of the two materials
• R increases rapidly as the difference in Z
increases
Z1
Z2
Ii
Ir
It
R =
Z2 – Z1
Z1 + Z2
Ii
Ir
= ( )
2

33. Reflection
• The fraction of ultrasound intensity
transmitted at an interface is given by the
intensity transmission coefficient, T
• Ultrasound imaging is only possible when
the wave propagates through materials
with similar acoustic impedances – only a
small fraction of energy is reflected and
the rest is transmitted
Z1
Z2
Ii
Ir
It
T =
Ii
It
= 1 - R

34. Reflection
Tissue Interface R T
Liver – Fat 0.01 0.99
Fat – Muscle 0.02 0.98
Muscle - Bone 0.41 0.59
Muscle - Air 0.99 0.01
At soft tissue – soft tissue
interfaces 1-2% of the
ultrasound intensity is
reflected
At soft tissue – air
interfaces 99% of the
incident intensity is
reflected
• At soft tissue – air or soft tissue – bone interfaces, a large proportion of the
incident intensity is reflected, making anatomy beyond such interfaces
unobservable
• Acoustic coupling gel is used between the face of the ultrasound transducer
and skin to eliminate air pockets

35. Reflection
• When the wave is not incident
perpendicular to the interface, non-normal
incidence, the reflected angle is equal to
the incident angle (i.e. θr = θi)
• Echoes are directed away from the source
of ultrasound and may be undetected
• The transmitted wave does not continue in
the incident direction (i.e. θt ≠ θi)
• The change in direction is described by
Refraction
Z1
Z2
Incident Reflection
(echo)
Transmission
(refraction)
θi θr
θt

36. Refraction
• Refraction describes the change in direction of the
transmitted ultrasound wave at a tissue interface
when the wave is not incident perpendicular to the
interface
• The angle of refraction, θt , is determined by the
speed of sound change that occurs as the wave
crosses the boundary
• The angle of refraction is related to the angle of
incidence by Snell’s law:
Z1 = 1 c1
Incident Reflection
(echo)
Transmission
(refraction)
θi θr
θt
Z2 = 2 c2
sin θt
sin θi
=
c2
c1

37. Refraction

38. Refraction
When c2 < c1 the angle of transmission is
less than the angle of incidence
c2 < c1
Incident Reflection
(echo)
Transmission
(refraction)
θi θr
θt
When c2 > c1 the angle of transmission is
greater than the angle of incidence
c2 > c1
Incident Reflection
(echo)
Transmission
(refraction)
θi θr
θt
c1 c1
c2 c2

39. Refraction
• A condition known as total reflection occurs
when c2 > c1 and the angle of incidence exceeds
an angle called the critical angle , θc
• When θi = θc the sound wave does not continue
into the second medium but travels along the
boundary
• The critical angle is calculated by setting θt = 90o
in Snell’s law, giving sinθc = c1/c2
c1
Incident Reflection
(echo)
Transmission
(refraction)
θi θr
θt
c2
c2 > c1

40. Refraction
• Refraction does not occur when the speed of sound is the same in the two media, or
when a sound wave is incident perpendicular to the interface
• This “straight-line” propagation is assumed by the ultrasound system during signal
processing
• When refraction does occur, this can result in image artefacts due to the
misplacement of anatomy in the image
Anatomical feature here
Displayed in image here

41. Scattering
• Reflection occurs at large tissue interfaces, such as those between organs, where
there is a change in acoustic impedance
• These large specular reflectors represent a “smooth” boundary where the size of the
boundary is much larger than the wavelength of the incident ultrasound wave
• Within most tissues and organs there are many small-scale variations in acoustic
properties which constitute small-scale reflecting particles that are similar in size or
smaller than the wavelength of the ultrasound
• These small non-specular reflectors represent a “rough” surface and give rise to
acoustic scattering within the insonated tissues

42. Scattering
• Scattering from non-specular reflectors reflects
sound in all directions
• Scattering is a weak interaction in that the
amplitude of the returning echoes are
significantly weaker than those from tissue
boundaries
• Intensities of the returning echoes from non-
specular reflectors within the tissue are not
greatly dependent on beam direction, unlike
specular reflectors
• The scattering pattern is characteristic of the
particle size and gives rise to tissue or organ
signatures that lead to a specific speckle or
textured appearance in the ultrasound image

43. Scattering
• Tissue boundary interactions can also give rise to
scatter
• Specular reflection assumes a “smooth” interface,
where the wavelength of the ultrasound is much
greater than the structural variations of the
interface
• With higher frequency ultrasound waves, the
wavelength becomes smaller and the interface no
longer appears “smooth”
• Returning echoes are diffusely scattered (non-
specular reflection) and only a fraction of the
reflected intensity returns to the transducer
• Scattering from non-specular reflectors increases
with ultrasound frequency, but specular reflection is
relatively independent
Z1
Z2
Incident Non-
specular
Reflection
Transmission
(refraction)

44. Attenuation
• As an ultrasound wave propagates through a tissue, the energy of the wave reduces
with the distance travelled
• Attenuation describes the reduction in beam intensity with distance travelled and is
primarily caused by scattering and tissue absorption of the incident beam
• The attenuation coefficient, , (in units dB/cm) is the relative intensity loss per cm
of travel for a given tissue
• The attenuation coefficient varies widely between different tissues and media
• The attenuation coefficient for a given tissue varies with ultrasound frequency;
Attenuation increases linearly with increasing frequency
• For “soft tissue”, the attenuation coefficient can be approximated as 0.5
(dB/cm)/MHz

45. Attenuation
Tissue Attenuation Coefficient
(1 MHz Beam, dB/cm)
Water 0.0002
Blood 0.18
Brain 0.3 – 0.5
Liver 0.4 – 0.7
Fat 0.5 – 1.8
Muscle 0.2 – 0.6
Bone 13 - 26
Lung 40

46. Attenuation
• Ultrasound beam intensity reduces
exponentially due to attenuation,
according to:
Distance travelled, d
Relative
Intensity,
I
Low frequency
High frequency
I = Ioe- d
Io = Initial intensity
1
0.5

47. Attenuation
• The ultrasound half-value thickness
(HVT) is the thickness of tissue
necessary to attenuate the incident
intensity by 50% (or 3 dB)
• The HVT decreases as the frequency
increases
• When penetration to deeper
structures is important, lower
frequency ultrasound transducers
are required
Distance travelled, d
Relative
Intensity,
I
Low frequency
High frequency
1
0.5

48. Attenuation
• In soft tissues a significant proportion of energy loss (attenuation) is due to tissue
absorption
• Absorption is the process by which ultrasound energy is converted into heat energy in
tissue
• Energy lost through absorption does not contribute to image formation
• Ultrasound attenuation is usually expressed in terms of decibels (dB)
Decibel Notation
Relative Intensity (dB) = 10 log10 (I2 / I1)
Where I1 = initial intensity, I2 = final intensity

49. Construction and Operation of the Ultrasound
Transducer

50. Ultrasound Transducer
• The transducer is the device that converts electrical transmission pulses into
ultrasonic pulses, and ultrasonic echo pulses into electrical signals
• A transducer produces ultrasound pulses and detects echo signals using the
piezoelectric effect
• The piezoelectric effect describes the interconversion of electrical and mechanical
energy in certain materials
• If a voltage pulse is applied to a piezoelectric material, the material will expand or
contract (depending on the polarity of the voltage)
• If a force is applied to a piezoelectric material which causes it to expand or
contract (e.g. pressure wave), a voltage will be induced in the material

51. Ultrasound Transducer

52. Ultrasound Transducer
• A piezoelectric material called PZT is commonly used in transducers
• A transducer only generates a useful ultrasound beam at one given frequency
• This frequency corresponds to a wavelength in the transducer equal to twice the
thickness of the piezoelectric disk – This is due to a process known as Resonance!
• Choice of frequency is important – remember that attenuation increases with
increasing frequency
• Image resolution increases with frequency
• Therefore, there is a trade-off between scan depth and resolution for any
particular application

53. Ultrasound Transducer
Linear Array Curvilinear/Convex Array Phased Array
Rectangular FOV
Useful in applications
where there is a need to
image superficial areas at
the same time as organs at
a deeper level
Trapezoidal FOV
Wide FOV near transducer
and even wider FOV at
deeper levels
Sector FOV
Useful for imaging heart
where access in normally
through a narrow acoustic
window between ribs

54. Beam Shape - Diffraction
• Diffraction is the process by which the ultrasound wave diverges (spreads out) as it
moves away from the source
• Divergence is determined by the relationship between the width of the source
(aperture) and the wavelength of the wave
Low Divergence
Aperture large compared to 
High Divergence
Aperture small compared to 

55. Beam Shape - Diffraction
NEAR FIELD FAR FIELD
NFL
a
Near Field Length, NFL = a2 /  a = radius of transducer
 = Wavelength

56. Beam Shape - Diffraction
• In the near field region the beam energy is largely confined to the dimensions of
the transducer
• Need to select a long near field length to achieve good resolution over the depth
you wish to scan too
• Near field length increases with increasing transducer radius, a, and decreasing
wavelength, 
• Short wavelength means high frequency – not very penetrating
• Large transducer radius – Wide beam (poor lateral resolution)
• Trade-off between useful penetration depth and resolution!!

57. Beam Focusing
• An improvement to the overall beam width can be obtained by focusing
• Here the source is designed so that the waves converge towards a point in the
beam, the focus, where the beam achieves its minimum width
• Beyond the focus, the beam diverges again but more rapidly that for an unfocused
beam with the same aperture and frequency
a
W
F
Beam width at focus,
W = F / a
At focal point:
• Maximum ultrasound intensity
• Maximum resolution

58. Beam Focusing
For a single element source, focusing can be achieved in one of two ways:
1) A curved source
A curved source is manufactured with a radius of curvature of F and hence
produces curved wave fronts which converge at a focus F cm from the source
F
Source Focus

59. Beam Focusing
For a single element source, focusing can be achieved in one of two ways:
2) An acoustic lens
An acoustic lens is attached to the face of a flat source and produces curved
wave fronts by refraction at its outer surface (like an optical lens). A convex lens
is made from a material with the lower speed of sound than tissue.
Source Focus
Lens

60. Beam Shape
Single transducer element is very
small.
Beam of one element has very
short near field length followed by
significant divergence.

61. Beam Shape – Overlapping Groups of Elements
Fire elements
1-5 together
And then…
Fire elements
2-6 together
And so on…
Near field length increases as (N)2
Image element
line 3 Image element
line 4

62. Array Focusing
Waves from outer elements 1 and 5 have
greater path lengths than those from other
elements
Therefore signals do not arrive simultaneously
at the target and reflections do not arrive at all
elements at the same time

63. Array Focusing
Introduce time delays to compensate for extra
path length on both transit and receive
Time delays
A large-summed signal is obtained for echoes
from the focal zone
Only a weak-summed signal (noise) results
from echoes elsewhere

64. Multiple Zone Focusing
• Fire transducer several times with different focus to compile better image
• However, more focus points decreases frame rate

65. Image Resolution
Resolution in three planes
Axial Slice Thickness
Lateral

66. Image Resolution
Resolution Depends on Typical Value (mm)
Axial Pulse length 0.2 - 0.5
Lateral Beam width 2 – 5
Slice Thickness Beam height 3 - 8
• Higher frequency improves resolution in all three planes
• Slice thickness is a hot topic for improvement – 2D arrays

67. Ultrasound Instrumentation

68. Instrumentation
Transmitter Clock
TGC Generator Transducer Beam Controller
AD Converter
Signal Processor Image Store
Archive Display
x, y
z

69. Instrumentation
Clock
• Command and control centre
• Sends synchronising pulses around the system
• Each pulse corresponds to a command to send a new pulse from the
transducer
• Determines the pulse repetition frequency (PRF)
PRF = 1 / time per line = c / 2D
Where c is speed of sound and D is maximum scan depth
If there are N lines, then Frame Rate = c / 2ND

70. Instrumentation
Transmitter
• Responds to clock commands by generating high voltage pulses to excite
transducer
Transducer
• Sends out short ultrasound pulses when excited
• Detects returning echoes and presents them as small electrical signals

71. Instrumentation
AD Converter
• Converts analogue echo signals into digital signals for further processing
Needs to:
• Be fast enough to cope with highest frequencies
• Have sufficient levels to create adequate grey scales (e.g. 256 or 512)

72. Instrumentation
Signal Processor
Carries out:
• TGC application
• Overall gain
• Signal compression – fits very large dynamic range ultrasound signal on to
limited greyscale display dynamic range
• Demodulation – removal of the carrier (ultrasound) frequency
Grey level
Input Amp
Linear
Liver
Heart

73. Instrumentation
Image Store
• Takes z (brightness) signal from processor
• Positions it in image memory using x (depth) and y (element position)
information from beam controller
• Assembles image for each frame
• Presents assembled image to display
• Typically have capacity to store 100-200 frames to allow cine-loop

74. Ultrasound Safety

75. Hazard and Risk
• There are two main hazards associated with ultrasound:
- Tissue heating
- Cavitation

76. Tissue Heating
• During a scan some of the ultrasound energy is absorbed by the exposed tissue and converted to
heat causing temperature elevation
• Elevated temperature affects normal cell function
• The risk associated with this hazard depends on the:
- Degree of temperature elevation
- Duration of the elevation
- Nature of the exposed tissue
Rate of energy absorption per unit volume
q = 2I
Where  = absorption coefficient,  = frequency, I = intensity

77. Tissue Heating
• Thermal effects in patient are complex
• Temperature increase will be fastest at the focus resulting in a temperature
gradient
• Heat will be lost from focus by thermal conduction
• The transducer itself will heat up and this heat will conduct into tissue
enhancing the temperature rise near the transducer
• The presence of bone in the field will increase the temperature rise
• Blood flow will carry heat away from the exposed tissues
• It is impossible to accurately predict the temperature increase occurring in
the body and a simple approach to estimate the temperature increase is used
to provide some guidance - Thermal Index (TI)

78. Thermal Index (TI)
TI = W / Wdeg
W = Transducer power exposing the tissue
Wdeg = The power required to cause a maximum temperature rise of 1oC
anywhere in the beam
• TI is a rough estimate of the increase in temperature that occurs in the region
of the ultrasound scan
• A TI of 2.0 means that you can expect at temperature rise of about 2oC
• The difficulty with calculating the TI lies mostly in the estimation of Wdeg
• To simplify this problem there are three TIs

79. Thermal Index
Soft-Tissue Thermal Index (TIS)
Soft tissue
Maximum temperature
Bone-at-Focus Thermal Index (TIB)
Soft tissue
Maximum temperature
Bone

80. Ultrasound Safety
Cranial (or Bone-at-Surface) Thermal Index (TIC)
Soft tissue
Maximum temperature
Bone
All three TI values depend linearly on the acoustic power emitted by the
transducer

81. Tissue Heating
Does Temperature Rise Matter?
• Normal core temperature is 36-38oC and a temperature of 42oC is “largely
incompatible with life”
• During an ultrasound examination only a small volume of tissue is exposed and
the human body is quite capable of recovering from such an event
• Some regions are more sensitive such as reproductive cells, unborn fetus, and
the CNS
• Temperature rises of between 3 and 8oC are considered possible under certain
conditions
• There has been no confirmed evidence of damage from diagnostic ultrasound
exposure

82. Cavitation
• Refers to the response of gas bubbles in a
liquid under the influence of an ultrasonic
wave
• Process of considerable complexity
• High peak pressure changes can cause
micro-bubbles in a liquid or near liquid
medium to expand – resonance effect
• A bubble may undergo very large size
variations and violently collapse
• Very high localised pressures and
temperature are predicted that have potential
to cause cellular damage and free radical
generation

83. Cavitation
Micro-bubbles grow by resonance processes
Bubbles have a resonant frequency, fr, depending on their radius, R.
frR  3 Hz m
This suggests that typical diagnostic frequencies (3 MHz and above) cause
resonance in bubbles with radii of the order of 1 micrometer

84. Mechanical Index (MI)
• The onset of cavitation only occurs above a threshold for acoustic pressure
• This has resulted in the formulation of a mechanical index (MI)
• Mechanical index is intended to quantify the likelihood of onset of cavitation
MI = pr / f
• where pr is the peak rarefaction pressure and f is the ultrasound frequency
• For MI  0.7 the physical conditions probably cannot exist to support bubble
growth and collapse
• Exceeding this threshold does not mean there will be automatically be
cavitation
• Cavitation is more likely in the presence of contrast agents and in the
presence of gas bodies such as in the lung and intestine

85. Thank you