This document discusses ultrasound transducers and resolution. It begins by describing how ultrasound is produced and detected using a transducer composed of piezoelectric elements. Over time, transducers have evolved from single elements to arrays with hundreds of individual elements. The key components of a basic transducer are then outlined. The remainder of the document provides detailed explanations of piezoelectric materials, resonance transducers, damping blocks, matching layers, and the properties of transducer arrays including linear arrays and phased arrays. Beam properties such as the near field and far field are also defined.

1. ULTRASOUND
TRANSDUCERS AND
RESOLUTION
Dr V S R Bhupal

2.  Ultrasound
is produced and detected
with a transducer, composed of one or
more ceramic elements with
electromechanical (piezoelectric)
properties.
• The ceramic element converts electrical
energy into mechanical energy to produce
ultrasound and mechanical energy into
electrical energy for ultrasound detection.

3. 
Over the past several decades, the transducer
assembly has evolved considerably in design,
function, and capability, from a single-element
resonance crystal to a broadband transducer
array of hundreds of individual elements.
•
A simple single-element, plane-piston source
transducer has major components including the
• piezoelectric material,
• matching layer,
• backing block,
• acoustic absorber,
• insulating cover,
• sensor electrodes, and
• transducer housing.

5. Piezoelectric Materials
A
piezoelectric material (often a crystal
or ceramic) is the functional component
of the transducer.
• It converts electrical energy into mechanical
(sound) energy by physical deformation of the
crystal structure.

6.  ConverseIy,
mechanical pressure
applied to its surface creates electrical
energy.
• Piezoelectric materials are characterized by a
well-defined molecular arrangement of
electrical dipoles.

7.  An
electrical dipole is a molecular entity
containing positive and negative electric
charges that has no net charge.
• When mechanically compressed by an
externally applied pressure, the alignment of
the dipoles is disturbed from the equilibrium
position to cause an imbalance of the charge
distribution.

8. A
potential difference (voltage) is created
across the element with one surface
maintaining a net positive charge and
one surface a net negative charge.
• Surface electrodes measure the voltage,
which is proportional to the incident
mechanical pressure amplitude.

9.  Conversely,
application of an external
voltage through conductors attached to
the surface electrodes induces the
mechanical expansion and contraction of
the transducer element.

10.  There
are natural and synthetic
piezoelectric materials.
• An example of a natural piezoelectric material
is quartz crystal, commonly used in watches
and other time pieces to provide a mechanical
vibration source at 32.768 kHz for interval
timing.
• This is one of several oscillation frequencies of
quartz, determined by the crystal cut and
machining properties.

11. 
Ultrasound transducers for medical imaging
applications employ a synthetic piezoelectric
ceramic, most often lead-zirconate-titanate
(PZT).
•
The piezoelectric attributes are attained after a
process of
• Molecular synthesis,
• Heating,
• Orientation of internal dipole structures with an applied
external voltage,
• Cooling to permanently maintain the dipole orientation,
and
• Cutting into a specific shape.

12. 
For PZT in its natural state, no piezoelectric
properties are exhibited; however, heating the
material past its “Curie temperature” (i.e., 3280
C to 3650 C) and applying an external voltage
causes the dipoles to align in the ceramic.
•
The external voltage is maintained until the material
has cooled to below its Curie temperature.
• Once the material has cooled, the dipoles retain their
alignment.

13.  At
equilibrium, there is no net charge on
ceramic surfaces.
• When compressed, an imbalance of charge
produces a voltage between the surfaces.
• Similarly, when a voltage is applied between
electrodes attached to both surfaces, mechanical
deformation occurs.

14.  The
piezoelectric element is composed
of aligned molecular dipoles.

15.  Under
the influence of mechanical
pressure from an adjacent medium
(e.g., an ultrasound echo), the element
thickness
• Contracts (at the peak pressure amplitude),
• Achieves equilibrium (with no pressure) or
• Expands (at the peak rarefactional pressure),
• This causes realignment of the electrical dipoles to
produce positive and negative surface charge.

17.  Surface
electrodes measure the
voltage as a function of time.

18.  An
external voltage source applied to the
element surfaces causes compression or
expansion from equilibrium by
realignment of the dipoles in response to
the electrical attraction or repulsion
force.

20. Resonance Transducers

Resonance transducers for pulse echo
ultrasound imaging are manufactured to
operate in a “resonance” mode, whereby a
voItage (commonly 150 V) of very short
duration (a voltage spike of ≈1 µsec) is
applied, causing the piezoelectric material to
initially contract, and subsequently vibrate at a
natural resonance frequency.
•
This frequency is selected by the “thickness cut,” due
to the preferential emission of ultrasound waves
whose wavelength is twice the thickness of the
piezoelectric material.

21.  The
operating frequency is determined
from the speed of sound in, and the
thickness of, the piezoelectric material.
• For example, a 5-MHz transducer will have a
wavelength in PZT (speed of sound in PZT is
≈ 4,000 m/sec) of
c 4000 m / sec
λ= =
= 8 × 10 −4 meters = 0.80 mm
f
5 × 106 / sec

22. 
A short duration
voltage spike causes
the resonance
piezoelectric element
to vibrate at its
natural frequency, fo,
which is determined
by the thickness of
the transducer equal
to 1/A.

23.  To
achieve the 5-MHz resonance
frequency, a transducer element
thickness of ½ X 0.8 mm = 0.4 mm is
required.
• Higher frequencies are achieved with thinner
elements, and lower frequencies with thicker
elements.
• Resonance transducers transmit and receive
preferentially at a single “center frequency.”

24. Damping Block

The damping block, layered on the back of the
piezoelectric element, absorbs the backward
directed ultrasound energy and attenuates
stray ultrasound signals from the housing.
•
This component also dampens the transducer
vibration to create an ultrasound pulse width and short
spatial pulse length, which is necessary to preserve
detail along he beam axis (axial resolution).

27.  Dampening
of the vibration (also known
as “ring-down”) lessens the purity of the
resonance frequency and introduces a
broadband frequency spectrum.
• With ring-down, an increase in the bandwidth
(range of frequencies) of the ultrasound pulse
occurs by introducing higher and lower
frequencies above and below the center
(resonance) frequency.

28. 
The “Q factor” describes the bandwidth of the
sound emanating from a transducer as
fo
Q=
Bandwidth

where fo is the center frequency and the
bandwidth is the width of the frequency
distribution.

29. A
“high Q” transducer has a narrow
bandwidth (i.e., very little damping) and a
corresponding long spatial pulse length.
• A “low Q” transducer has a wide bandwidth
and short spatial pulse length.

30.  Imaging
applications require a broad
bandwidth transducer in order to achieve
high spatial resolution along the direction
of beam travel.
• Blood velocity measurements by Doppler
instrumentation require a relatively narrowband transducer response in order to
preserve velocity information encoded by
changes in the echo frequency relative to the
incident frequency.

31.  Continuous-wave
ultrasound transducers
have a very high Q characteristic.
• While the Q factor is derived from the term
quality factor, a transducer with a low Q does
not imply poor quality in the signal.

32. Matching Layer

The matching layer provides the interface
between the transducer element and the tissue
and minimizes the acoustic impedance
differences between the transducer and the
patient.
•
It consists of layers of materials with acoustic
impedances that are intermediate to those of soft
tissue and the transducer material.
• The thickness of each layer is equal to one-fourth the
wavelength, determined from the center operating
frequency of the transducer and speed of sound in the
matching layer.

33.  For
example, the wavelength of sound in
a matching layer with a speed of sound
of 2,000 m/sec for a 5-MHz ultrasound
beam is 0.4 mm.
• The optimal matching layer thickness is equal
to ¼λ = ¼ x 0.4 mm = 0. 1 mm.
• In addition to the matching layers, acoustic
coupling gel (with acoustic impedance similar to
soft tissue) is used between the transducer and the
skin of the patient to eliminate air pockets that
could attenuate and reflect the ultrasound beam.

34. Nonresonance (BroadBandwidth) “Multifrequency”
Transducers

Modern transducer design coupled with digital
signal processing enables “multifrequency or
“multihertz” transducer operation, whereby rhe
center frequency can be adjusted in he
transmit mode.
•
Unlike the resonance transducer design, the
piezoelectric element is intricately machined into a
large number of small “rods,” and then filled with an
epoxy resin to create a smooth surface.

37.  The
acoustic properties are closer to
tissue than a pure PZT material, and
thus provide a greater transmission
efficiency of the ultrasound beam without
resorting to multiple matching layers.
• Multifrequency transducers have bandwidths
that exceed 80% of the center frequency.

38.  Excitation
of the multifrequency
transducer is accomplished with a short
square wave burst of 150 V with one to
three cycles, unlike the voltage spike
used for resonance transducers.
• This allows the center frequency to be
selected within the limits of the transducer
bandwidth.

39.  Likewise,
the broad bandwidth response
permits the reception of echoes within a
wide range of frequencies.
• For instance, ultrasound pulses can be
produced at a low frequency, and the echoes
received at higher frequency.

40.  “Harmonic
imaging” is a recently
introduced technique that uses this
ability;
• lower frequency ultrasound is transmitted into
the patient, and the higher frequency
harmonics (e.g., two times the transmitted
center frequency) created from the interaction
with contrast agents and tissues, are received
as echoes.

41.  Native
tissue harmonic imaging has
certain advantages including greater
depth of penetration, noise and clutter
removal, and improved lateral spatial
resolution.

42. Transducer Arrays
 The
majority of ultrasound systems
employ transducers with many individual
rectangular piezoelectric elements
arranged in linear or curvilinear arrays.
• Typically, 128 to 512 individual rectangular
elements compose the transducer assembly.
• Each element has a width typically less than half
the wavelength and a length of several millimeters.

43. 
Two modes of
activation are used
to produce a beam.
•
These are the “linear”
(sequential) and
“phased”
activation/receive
modes.

44. Linear Arrays
 Linear
array transducers typically contain
256 to 512 elements; physically these
are the largest transducer assemblies.

45.  In
operation, the simultaneous firing of’ a
small group of ≈ 20 adjacent elements
produces the ultrasound beam.
• The simultaneous activation produces a
synthetic aperture (effetive transducer width)
defined by the number of active elements.

46.  Echoes
are detected in the receive mode
by acquiring signals from most of the
transducer elements.
• Subsequent “A-line” acquisition occurs by
firing another group of transducer elements
displaced by one or two elements.

47. A
rectangular field of view is produced
with this transducer arrangement.
• For a curvilinear array, a trapezoidal field of
view is produced.

48. Phased Arrays
A
phased-array transducer is usually
composed of 64 to 128 individual
elements in a smaller package than a
linear array transducer.
• All transducer elements are activated nearly
(but not exactly) simultaneously to produce a
single ultrasound beam.

49. 
By using time delays in the electrical activarion
of the discrete elements across the face of the
transducer, the ultrasound beam can be
steered and focused electronically without
moving the transducer.
•
During ultrasound signal reception, all of the
transducer elements detect the returning echoes from
the beam path, and sophisticated algorithms
synthesize the image from the detected data.

50. BEAM PROPERTIES
 The
ultrasound beam propagates as a
longitudinal wave from the transducer
surface into the propagation medium,
and exhibits two distinct beam patterns:
• a slightly converging beam out to a distance
•
specified by the geometry and frequency of
the transducer (the near field), and
a diverging beam beyond that point (the far
field).

51. 
For an unfocused,
single-element
transducer, the
length of the near
field is determined
by the transducer
diameter and the
frequency of the
transmitted sound.

52.  For
multiple transducer element arrays,
an “effective” transducer diameter is
determined by the excitation of a group
of’ transducer elements.
• Because of the interactions of each of the
individual beams and the ability to focus
and steer the overall beam, the formulas
for a single-element, unfocused transducer
are not directly applicable.

53. The Near Field
 The
near field, also known as the
Fresnel zone, is adjacent to the
transducer face and has a converging
beam profile.
• Beam convergence in the near field occurs
because of multiple constructive and
destructive interference patterns of the
ultrasound waves from the transducer
surface.

54.  Huygen’s
principle describes a large
transducer surface as an infinite
number of point sources of sound
energy where each point is
characterized as a radial emitter.
• By analogy, a pebble dropped in a quiet pond
creates a radial wave pattern.

55. 
As individual wave
patterns interact, the
peaks and troughs from
adjacent sources
constructively and
destructively interfere,
causing the beam profile
to be tightly collimated in
the near field.

56.  The
ultrasound beam path is thus largely
confined to the dimensions of the active
portion of the transducer surface, with
the beam diameter converging to
approximately half the transducer
diameter at the end of the near field.

57.  The
near field length is dependent on the
transducer frequency and diameter:
d 2 r2
Near field length =
=
4λ λ
• where d is the transducer diameter, r is the
transducer radius, and λ is the wavelength of
ultrasound in the propagation medium.

58. soft tissue, λ = 1.54mm/f(MHz), and
the near field length can be expressed
as a function of frequency:
 In
(
)
d2
mm 2 ( MHz )
Near field length =
( mm)
4 × 1.54

59. 
A higher transducer
frequency (shorter
wavelength) will
result in a longer
near field, as will a
larger diameter
element.

60.  For
a 10-mm-diameter transducer, the
near field extends 5.7 cm at 3.5 MHz and
16.2 cm at 10 MHz in soft tissue.
• For a 15-mm-diameter transducer, the
corresponding near field lengths are 12.8 and
36.4 cm, respectively.

61.  Lateral
resolution (the ability of the
system to resolve objects in a direction
perpendicular to the beam direction) is
dependent on the beam diameter and is
best at the end of the near field for a
single-element transducer.
• Lateral resolution is worst in areas close to
and far from the transducer surface.

62.  Pressure
amplitude characteristics in the
near field are very complex, caused by
the constructive and destructive
interference wave patterns of the
ultrasound beam.
• Peak ultrasound pressure occurs at the end
of the near field, corresponding to the
minimum beam diameter for a single-element
transducer.

63.  Pressures
vary rapidly from peak
compression to peak rarefaction several
times during transit through the near
field.
• Only when the far field is reached do the
ultrasound pressure variations decrease
continuously.

64.  The
far field is also known as the
Fraunhofer zone, and is where the beam
diverges.
• For a large-area single-element transducer,
the angle of ultrasound beam divergence, 0,
for the far field is given by
λ
sin θ = 1.22
d
• where d is the effective diameter of the
transducer and λ is the wavelength; both must
have the same units of distance.

65.  Less
beam divergence occurs with highfrequency, large-diameter transducers.
• Unlike the near field, where beam intensity
varies from maximum to minimum to
maximum in a converging beam, ultrasound
intensity in the far field decreases
monotonically with distance.

66. Transducer Array Beam
Formation and Focusing
 In
a transducer array, the narrow
piezoelectric element width (typically less
than one wavelength) produces a
diverging beam at a distance very close
to the transducer face.
• Formation and convergence of the ultrasound
beam occurs with the operation of several or
all of the transducer elements at the same
time.

67. 
Transducer elements in a linear array that are
fired simultaneously produce an effective
transducer width equal to the sum of the
widths of the individual elements.
• Individual beams interact via constructive and
destructive interference to produce a collimated
beam that has properties similar to the properties
of a single transducer of the same size.

68.  With
a phased-array transducer, the
beam is formed by interaction of the
individual wave fronts from each
transducer, each with a slight difference
in excitation time.
• Minor phase differences of adjacent beams
form constructive and destructive wave
summations that steer or focus the beam
profile.

69. COMMON TRANSDUCERS USED
IN CLINICAL SETTING

70. STRAIGHT LINEAR ARRAY
PROBE
The straight linear array probe is designed
for superficial imaging.
The crystals are aligned in a linear fashion
within a flat head and produce sound
waves in a straight line.
The image produced is rectangular in
shape.

71.  This
probe has higher frequencies (5–13
MHz), which provides better resolution
and less penetration.
 Therefore, this probe is ideal for imaging
superficial structures and in ultrasoundguided procedures.

72. Vascular access
Evaluate for deep venous thrombosis
Skin and soft tissue for abscess, foreign
body
Musculoskeletal—tendons, bones,
muscles

74. CURVILINEAR ARRAY
PROBE
 The
curvilinear array or convex probe is
used for scanning deeper structures. The
crystals are aligned along a curved
surface and cause a fanning out of the
beam, which results in a field of view that
is wider than the probe’s footprint.

75.  The
image generated is sector shaped.
These probes have frequencies ranging
between 1 and 8 MHz, which allows for
greater penetration, but less resolution.
These probes are most often used in
abdominal and pelvic applications.
 They are also useful in certain
musculoskeletal evaluations or
procedures when deeper anatomy needs
to be imaged or in obese patients.

76.  Abdominal
aorta
 Biliary/gallbladder/liver/pancreas
 Abdominal portion of FAST exam
 Kidney and bladder evaluation
 Transabdominal pelvic evaluation

78. ENDOCAVITARY PROBE
 The
endocavitary probe also has a
curved face, but a much higher
frequency (8–13 MHz) than the
curvilinear probe.
 This probe’s elongated shape allows it to
be inserted close to the anatomy being
evaluated.

79.  The
curved face creates a wide field of
view of almost 180° and its high
frequencies provide superior resolution .
This probe is used most commonly for
gynecological applications, but can also
be used for intraoral evaluation of
peritonsillar abscesses.
 Transvaginal ultrasound
 Intraoral

81. PHASED ARRAY PROBE
 Phased
array probes (Fig. 4-4a) have
crystals that are grouped closely
together.
 The timing of the electrical pulses that
are applied to the crystals varies and
they are fired in an oscillating manner.

82.  The
sound waves that are generated
originate from a single point and fan
outward, creating a sector-type image.
This probe has a smaller and flatter
footprint than the curvilinear one, which
allows the user to maneuver more easily
between the ribs and small spaces.
These probes have frequencies between
2 and 8 MHz.

84. IVUS PROBE
 IVUS
is a miniature ultrasound probe
positioned at the tip of a coronary
catheter.
 The probe emits ultrasound frequencies,
typically at 20-45 MHz, and the signal is
reflected from surrounding tissue and
reconstructed into a real-time
tomographic gray-scale image.

87. Spatial Resolution

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

89. 
The axial, lateral,
and elevational (slice
thickness)
dimensions
determine the
minimal volume
element.

90.  Each
dimension has an effect on the
resolvability of objects in the image.

91. Axial Resolution
 Axial
resolution (also known as linear,
range, longitudinal, or depth resolution)
refers to the ability to discern two closely
spaced objects in the direction of the
beam.
• Achieving good axial resolution requires that
the returning echoes be distinct without
overlap.

92.  The
minimal required separation
distance between two reflectors is onehalf of the spatial pulse length (SPL) to
avoid the overlap of returning echoes, as
the distance traveled between two
reflectors is twice the separation
distance.

93. 
Objects spaced
closer than ½ SPL
will not be resolved.

94.  The
SPL is the number of cycles emitted
per pulse by the transducer multiplied by
the wavelength.
• Shorter pulses, producing better axial
resolution, can be achieved with greater
damping of the transducer element (to reduce
the pulse duration and number of cycles) or
with higher frequency (to reduce wavelength).

95.  For
imaging applications, the ultrasound
pulse typically consists of three cycles.
• At 5 MHz (wavelength of 0.31 mm), the SPL
is about 3 x 0.31 0.93 mm, which provides an
axial resolution of /2(0.93 mm) = 0.47 mm.

96.  At
a given frequency, shorter pulse
lengths require heavy damping and low
Q, broad-bandwidth operation.
• For a constant damping factor, higher
frequencies (shorter wavelengths) give better
axial resolution, but the imaging depth is
reduced.
• Axial resolution remains constant with depth.

97. Lateral Resolution
 Lateral
resolution, also known as
azimuthal resolution, refers to the ability
to discern as separate two closely
spaced objects perpendicular to the
beam direction.

98. 
For both single
element transducers
and multielement
array transducers,
the beam diameter
determines the
lateral resolution.

99.  Since
the beam diameter varies with the
distance from the transducer in the near
and far field, the lateral resolution is
depth dependent.
• The best lateral resolution occurs at the near
field—far field face.

100.  At
this depth, the effective beam
diameter is approximately equal to half
the transducer diameter.
• In the far field, the beam diverges and
substantially reduces the lateral resolution.

101.  The
typical lateral resolution for an
unfocused transducer is approximately 2
to 5 mm.
• A focused transducer uses an acoustic lens
(a curved acoustic material analogous to an
optical lens) to decrease the beam diameter
at a specified distance from the transducer.

102.  With
an acoustic lens, lateral resolution
at the near field-far field interface is
traded for better lateral resolution at a
shorter depth, but the far field beam
divergence is substantially increased.
• The lateral resolution of linear and curvilinear
array transducers can be varied.

103. Elevational Resolution
 The
elevational or slice-thickness
dimension of the ultrasound beam is
perpendicular to the image plane.
• Slice thickness plays a significant part in
image resolution, particularly with respect to
volume averaging of acoustic details in the
regions dose to the transducer and in the far
field beyond the focal zone.

104. 
Elevational
resolution is
dependent on the
transducer element
height in much the
same way that the
lateral resolution is
dependent on the
transducer element
width.

105.  Slice
thickness is typically the worst
measure of resolution for array
transducers.
• Use of a fixed focaI length lens across the
entire surface of the array provides improved
elevational resolution at the focal distance.

106.  Unfortunately,
this compromises
resolution due to partial volume
averaging before and after the
elevational focal zone (elevational
resolution quality control phantom image
shows the effects of variable resolution
with depth.

107.  Multiple
linear array transducers with five
to seven rows, known as 1.5dimensional (1.5-D) transducer arrays,
have the ability to steer and focus the
beam in the elevational dimension.