ULTRASOUND TRANSDUCERS AND

RESOLUTION

Dr V S R Bhupal

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.

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.

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.

ConverseIy, mechanical pressure applied to its surface creates electrical energy. • Piezoelectric materials are characterized by a

well-defined molecular arrangement of electrical dipoles.

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.

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.

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

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.

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.

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.

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.

The piezoelectric element is composed of aligned molecular dipoles.

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.

Surface electrodes measure the voltage as a function of time.

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.

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 sec) 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.

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

mmmetersm

f

c80.0108

sec/105

sec/4000 46

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.

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.”

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).

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.

The “Q factor” describes the bandwidth of the sound emanating from a transducer as

where fo is the center frequency and the bandwidth is the width of the frequency distribution.

Bandwidth

fQ o

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.

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 narrow-band transducer response in order to preserve velocity information encoded by changes in the echo frequency relative to the incident frequency.

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.

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.

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 ¼ = 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.

Nonresonance (Broad-Bandwidth) “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.

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.

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.

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.

“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.

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

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.

Two modes of activation are used to produce a beam. • These are the “linear”

(sequential) and “phased” activation/receive modes.

Linear Arrays

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

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.

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.

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

view is produced.

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.

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.

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).

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.

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.

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.

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.

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.

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.

The near field length is dependent on the transducer frequency and diameter:

• where d is the transducer diameter, r is the transducer radius, and is the wavelength of ultrasound in the propagation medium.

22

4

rdlengthfieldNear

In soft tissue, = 1.54mm/f(MHz), and the near field length can be expressed as a function of frequency:

mm

MHzmmdlengthfieldNear

22

54.14

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

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.

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.

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.

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.

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

• where d is the effective diameter of the transducer and is the wavelength; both must have the same units of distance.

d

22.1sin

Less beam divergence occurs with high-frequency, 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.

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.

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.

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.

COMMON TRANSDUCERS USED IN CLINICAL SETTING

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.

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 ultrasound-guided procedures.

Vascular access

Evaluate for deep venous thrombosis

Skin and soft tissue for abscess, foreign body

Musculoskeletal—tendons, bones, muscles

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.

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.

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

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.

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

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.

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.

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.

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.

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

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.

The minimal required separation distance between two reflectors is one-half 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.

Objects spaced closer than ½ SPL will not be resolved.

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).

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.

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.

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.

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

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.

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.

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.

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.

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.

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.

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.

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.

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