Rep. Prog. Phys. 62 (1999) 671–722. Printed in the UK
PII: S0034-4885(99)74694-4
Ultrasonic imaging of the human body
P N T Wells
Department of Medical Physics and Bioengineering and Centre for Physics and Engineering Research in Medicine,
Bristol General Hospital, Bristol BS1 6SY, UK
Received 30 November 1998
Abstract
Ultrasonic imaging is a mature medical technology. It accounts for one in four imaging
studies and this proportion is increasing. Wave propagation, beam formation, the Doppler
effect and the properties of tissues that affect imaging are discussed. The transducer materials
and construction of the probes used in imaging are described, as well as the methods of
measuring the ultrasonic field. The history of ultrasonic imaging is briefly reviewed. The
pulse–echo technique is used for real-time grey-scale imaging and the factors that limit
the spatial and temporal resolutions are considered. The construction and performance
of transducer arrays are discussed, together with the associated beam steering and signal
processing systems. Speckle and scattering by blood are introduced, particularly in the
context of the observation of blood flow by means of the Doppler effect and by time-domain
signal processing. Colour flow imaging, and the colour coding schemes used for velocity and
power imaging, are explained. The acquisition and display of three-dimensional images are
discussed, with particular reference to speed and segmentation. Specialized imaging methods,
including endoluminal scanning, synthetic aperture imaging, computed tomography, elasticity
imaging, microscanning, contrast agents, and tissue harmonic imaging, are reviewed. There
is a discussion of issues relating to safety. Conclusions are drawn and future prospects are
considered.
0034-4885/99/050671+52$59.50
© 1999 IOP Publishing Ltd
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P N T Wells
Contents
1. Introduction
2. Physical foundations
2.1. Wave propagation
2.2. Beam formation
2.3. The Doppler effect
2.4. Radiation force
2.5. Basic ultrasonic properties of biological materials
2.6. The physical dimensions of ultrasonic imaging
2.7. Tissue inhomogeneity
2.8. Nonlinear propagation
2.9. Beam and pulse propagation in real tissue
3. Generation and detection of ultrasound
3.1. Transducer materials
3.2. Probe construction: elementary considerations
3.3. Matching, backing and loading: pulse operation
3.4. Beam studies
4. Image formation
4.1. Principles of pulse–echo ultrasound
4.2. Transducer array scanning
4.3. Signal processing and display for grey-scale pulse–echo imaging
4.4. Resolution
4.5. Speckle
4.6. Examples of real-time grey-scale scanning
4.7. Blood flow and tissue motion imaging
4.8. Three-dimensional image acquisition and display
4.9. Specialized imaging methods
5. Safety considerations
6. Conclusions and future prospects
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1. Introduction
More than one out of every four medical diagnostic imaging studies in the world is now
estimated to be an ultrasound study and the proportion continues to increase (WFUMB 1997).
This situation has come about because of the remarkable advances that have taken place in the
physics and engineering of ultrasonic imaging since the medical applications of ultrasonics
were last reviewed, some 30 years ago, in Reports on Progress in Physics (Wells 1970). In 1970,
the emphasis was on the biological effects, surgical and therapeutic applications of ultrasonics;
little more than 20% of that review was concerned with diagnosis. That was a fair balance
then, because medical imaging was dominated by x-radiography, with a small contribution
from radionuclide studies; neither x-ray computed tomography nor nuclear magnetic resonance
imaging had yet been invented and, except perhaps for applications in obstetrics, gynaecology
and cardiology, ultrasonic imaging was generally regarded only as a laboratory curiosity.
The choice of the best imaging technique in any given clinical situation is based on
considerations such as the resolution, contrast mechanism, speed, convenience, acceptability
and safety. Ultrasound scores highly in all of these: spatial resolution of a millimetre can be
obtained in abdominal scanning, tissue contrast is good (and can be enhanced by using contrast
agents), it is a real-time method, convenient to use, very acceptable to patients, and apparently
quite safe. It can also be a relatively inexpensive technology. Its main disadvantages are that
its images are spoilt by the presence of bone or gas, and the operator needs a high level of skill,
both in image acquisition and interpretation.
Without diagnostic imaging tools, the doctor is blind. Imaging is one of the cornerstones
of diagnosis in modern medical practice: the others are clinical examination and the various
branches of pathology. Moreover, the applications of imaging are not limited to diagnosis.
Increasingly, open access surgery is being replaced by minimally-invasive procedures, and
image-guided intervention is becoming more common. For many of these purposes, ultrasonic
imaging is the best method.
2. Physical foundations
2.1. Wave propagation
Ultrasonics is concerned with the propagation in various media of mechanical waves with
frequencies above the range of human hearing which, for convenience, means waves with
frequencies of more than 20 kHz.
Consider an infinitely-small-amplitude plane pressure wave propagating in a perfectly
elastic isotropic medium. The wave equation is
1 ∂ 2u
∂ 2u
= 2 2,
2
∂z
c ∂t
(1)
where u is the particle displacement amplitude, z is the position in space along the direction of
propagation, t is the time and c is the propagation speed. The speed is related to the elasticity
K and the density ρ of the medium in which the wave is travelling, according to the equation
c = (K/ρ)1/2 .
(2)
At a plane boundary between two media with speeds c1 and c2 respectively,
θi = θt
(sin θi / sin θt ) = c1 /c2 ,
(3)
(4)
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Figure 1. A simple representation of the ultrasonic beam produced by a disc transducer in a
homogeneous medium.
where θi , θr and θt are respectively the angles of incidence, reflexion and refraction. At normal
incidence,
Ir /Ii = [(Z2 − Z1 )/(Z2 + Z1 )]2 ,
(5)
where Ii and Ir are respectively the intensities of the incident and reflected waves and Z1 and
Z2 are the characteristic impedances of the two media. The characteristic impedance of a
medium is given by
Z = ρc.
(6)
The situation to which equation (5) applies is called specular reflexion and it implies that
the reflecting boundary is both smooth and extensive in relation to the wavelength λ. By
defining a quantity ψ related to the size of an obstacle, two situations can be distinguished,
each with a corresponding value of scattering cross section S:
S=1
when ψ λ
4 6
when ψ λ
S=k ψ
(7)
(8)
where k = 2πf and the frequency f = c/λ. Thus, specular reflexion is described by equation
(7) and Rayleigh scattering (Wells 1977), by equation (8). With obstacles of intermediate size
(or with rough surfaces), directional scattering occurs.
2.2. Beam formation
The aperture of the ultrasonic transducer used in medical imaging is usually in the form of a
circle or a rectangle. As illustrated in figure 1, the ultrasonic beam can be considered to consist
of a near field and a far field. With continuous wave excitation of a disc transducer,
Iz /I0 = sin2 {(π/λ)[(a 2 + z2 )1/2 − z2 ]},
(9)
where I0 is the intensity at the surface of the transducer, Iz is the intensity at a distance z from
the transducer along the central axis of the beam and a is the radius of the transducer. In the
far field, beyond the last axial maximum (at z = a 2 /λ, provided that a 2 λ2 ), the directivity
function is
2J1 (ka sin θ)
,
(10)
Ds =
ka sin θ
where θ is the angle between Ds and the central axis of the beam and J1 is the first-order Bessel
function. In the near field, the beam is roughly cylindrical with a series of axial maxima and
Ultrasonic imaging of the human body
675
minima of decreasing complexity moving away from the transducer. Also, in the near field,
the beam can be focused by a lens or other means.
If the transducer is excited to produce a transient disturbance, the ultrasonic pulse has its
energy spread over a spectrum of frequency, corresponding to its bandwidth. This means that
single values cannot be assigned to λ or k in equations (9) and (10). Physically, the beam
diffraction pattern is smeared to an extent which changes during the passage of the pulse.
2.3. The Doppler effect
When an ultrasonic wave is scattered by a target that has a component of velocity along
the direction of beam propagation, the frequency of the scattered ultrasound is shifted by the
Doppler effect. If θ is the angle between the direction of target motion and that of the ultrasonic
beam,
v = −fD c/(2f cos θ),
(11)
where v is the speed of the target and fD is the difference between the frequencies of the
ultrasound transmitted from the transducer and backscattered along the ultrasonic beam,
provided that v c. The negative sign means that the frequency is shifted downwards if
the target is moving away from the transducer.
2.4. Radiation force
The energy carried by an ultrasonic wave results in radiation force when the wave direction
changes (e.g., as the result of reflexion) or when ultrasound is absorbed from the wave. The
radiation force, F , resulting from complete absorption is given by
F = P /c,
(12)
where P is the ultrasonic power.
2.5. Basic ultrasonic properties of biological materials
The basic properties of biological materials that are of importance in ultrasonic imaging are
attenuation, speed and reflectivity. Typical values for various tissues are given in table 1.
Attenuation is due to the combined effects of absorption and, because tissue is inhomogeneous
over a range of scales, scattering and reflexion. In soft tissues, absorption is mainly due to
spectra of relaxation processes (Wells 1975), which accounts for the nearly linear frequency
dependence. In contrast, there is a square law dependence on frequency in simple liquids, in
which absorption is due to viscosity.
For practical purposes, speed dispersion can be neglected in soft tissues; typically, the
speed increases by about 0.01% MHz−1 (Carstensen and Schwan 1959).
When substituted in equation (5), the values of characteristic impedance given in table 1
give an indication of the strongest reflexions likely to occur when an ultrasonic beam travels
through the body. There is almost complete reflexion at the boundary between soft tissue
and air and also when a beam travelling through soft tissue encounters bone. In contrast, the
reflexion at a boundary between different kinds of soft tissues is quite small, leaving most of
the energy to travel across the boundary and into deeper tissues of the body.
Reflexion at the boundaries between different kinds of tissues is a rather idealized situation.
It is the backscattering of ultrasound as it travels through tissues that is generally much more
relevant to the imaging process.
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P N T Wells
Table 1. Properties of some materials relevant to ultrasonic imaging. Frequency dependences of α
apply at least within the range 1–10 MHz (except for bone, for which the range is 1–2 MHz). Data
collected by Duck (1990) and Wells (1977). Gaps in the table indicate that published data are not
readily available.
Material
Air
Blood
Brain
Fat
Liver
Muscle
Skull bone
Soft tissue
(mean values)
Water
Propagation
speed, c
(m s−1 )
Characteristic
impedance, Z
(106 kg m−2 s−1 )
330
1570
1540
1450
1550
1590
4000
1540
0.0004
1.61
1.58
1.38
1.65
1.70
7.80
1.63
1480
1.48
Attenuation
coefficient, α
at 1 MHz
(dB cm−1 )
1.2
0.2
0.9
0.6
0.9
1.5–3.5
13
0.6
0.002
Frequency
dependence
of α
f2
f 1,3
f
f
f
f
f2
f
f2
Nonlinear
parameter,
B/A
6.1
6.6
10
6.8
7.4
5.2
2.6. The physical dimensions of ultrasonic imaging
Ultrasonic imaging depends on the interactions between the structures of the (human) body
and ultrasonic radiation. As a starting point, the ultrasonic wavelength can be considered to
determine the spatial resolution. Consider the process of imaging the contents of the abdominal
cavity. Structures of interest for imaging are likely to be located at distances of up to about
150 mm beyond the skin surface and spatial resolution of the order of a millimetre is needed. To
obtain this resolution, the wavelength of the ultrasound needs to be not greater than 1 mm, which
is the wavelength at a frequency of 1.5 MHz. The problem is that the attenuation increases
with the frequency, so that the distance over which useful levels of energy can be propagated
becomes less as the frequency is increased. Typically, 3 MHz might be the maximum frequency
for 150 mm penetration. The corresponding wavelength is 0.5 mm. Assuming a mean value
of attenuation equal to 0.6 dB cm−1 MHz−1 (see table 1), the total attenuation, for the round
trip, is about 54 dB.
From equation (9), the aperture required for a disc transducer to produce a beam with a
near field depth of 150 mm at 0.5 mm wavelength is 17 mm in diameter; operation in the near
field is necessary for beam focusing to be effective.
At a frequency of 3 MHz, substitution in equation (11) shows that a Doppler shift frequency
of 4 kHz occurs when ultrasound is reflected by a target moving at a velocity of 1 m s−1 with
respect to the direction of the ultrasonic beam. The fact that the Doppler frequency usually
lies in the audible range in ordinary physiological situations is serendipitous.
2.7. Tissue inhomogeneity
Although it is possible to assign average values to speed, attenuation and scattering in soft
tissue, the different kinds of soft tissues each have their own individual properties and these
properties are inhomogeneously distributed within them. For example, table 1 gives a range of
values for attenuation in muscle, in which attenuation across the muscle fibres is around twice
that along the fibres.
An ultrasonic beam is distorted as it travels through tissue as a result of tissue
inhomogeneity. This is discussed further in sections 2.9 and 4.4.
Ultrasonic imaging of the human body
677
2.8. Nonlinear propagation
Although it is often convenient to assume that the pressure of a wave is proportional to the
particle displacement amplitude, this is really only justifiable at small amplitudes. Except at
infinitely-small amplitudes, the nonlinear relationship between pressure and density becomes a
significant factor. Physically, nonlinear propagation transfers wave energy to higher harmonics
(because propagation speed increases with density) so that an initially sinusoidal wave is
converted to a sawtooth. The resultant shock wave is accompanied by excess attenuation
(because attenuation increases with frequency), so that the rate of deposition of wave energy
has a spatial peak at some distance from the source. As the wave amplitude falls, so the
waveform reverts towards a sinusoidal shape.
The nonlinearity of a medium can be described in terms of its nonlinearity parameter B/A.
The quantities A and B are the coefficients of the first- and second-order terms of the Taylor
series expansion of the equation relating pressure to density in the medium. Typical values of
B/A are given in table 1.
The distribution of energy deposition when a beam propagates in a homogeneous nonlinear
medium is determined by the combined effects of nonlinearity, absorption and diffraction.
Aanonsen et al (1984) solved the nonlinear wave equation and, using a modification of their
code, Baker (1997) demonstrated that the location of peak intensity loss from a beam generated
by a circular source shifts away from the source and occupies a smaller diameter as the intensity
is increased. Physically, this is because of the build-up of higher harmonic frequencies in the
beam at the intensities at which the effect is significant.
2.9. Beam and pulse propagation in real tissue
Figure 2(a) represents a pulsed ultrasonic beam propagating in an idealized lossless
homogeneous material. The shape of the pulse is unaltered during propagation; its amplitude
in the cylindrical near field is constant and only reduces in the far field as the result of beam
divergence.
Real tissue possesses frequency-dependent attenuation and is nonlinear and
inhomogeneous. Figure 2(b) represents the propagation of a pulsed ultrasonic beam in tissue.
The pulse that leaves the transducer is the same as that represented in figure 2(a). After
travelling some distance in real tissue, however, a finite amplitude pulse is not only attenuated
but has also been partially converted to a shock wave as the result of nonlinearity. Deeper
into the tissue, the pulse has been further attenuated, preferentially at its higher frequency
components, and has reverted to a less shocked shape with the remaining energy distributed
across lower frequencies. The beam has also been deviated and distorted by inhomogeneity.
3. Generation and detection of ultrasound
3.1. Transducer materials
The ideal transducer for ultrasonic imaging would be perfectly matched to the (human) body,
have high efficiency as a transmitter and high sensitivity as a receiver, a wide dynamic range
and a wide frequency response for pulse operation.
Transducers based on the piezoelectric effect are almost always used in ultrasonic imaging.
Piezoelectricity occurs in many natural materials, quartz perhaps being the best known, but
these natural materials have characteristics which are not very suitable for medical ultrasonics.
The artificial ferroelectric ceramics (Jaffe et al 1955) approach much more closely to the
ideal. When the crystalline structure of a material has no centre of symmetry, it is said to
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P N T Wells
Figure 2. Simple representations of the ultrasonic beam produced by pulsed excitation of a disc
transducer. (a) The beam in an idealized lossless homogeneous medium. The beam shape is that
illustrated in figure 1. The pulse shape does not change during propagation, buts its amplitude
is reduced in the far field as the result of beam divergence. (b) The beam in an inhomogeneous
nonlinear attenuating medium, such as real tissue. The beam is deviated as the result of the
inhomogeneity: under some circumstances, there may be several cross sectional maxima. The
pulse develops shocks as the result of the nonlinearity: this leads to rapid attenuation of the higher
frequency components and the centre frequency of the attenuated pulse is shifted downwards.
be noncentrosymmetric. Ferroelectric ceramics have noncentrosymmetric unit cells which,
when below the Curie temperature, are randomly orientated, but which can be permanently
preferentially aligned by the brief application of a large polarizing potential.
The behaviour of a piezoelectric transducer can be described, in terms of its efficiency as
a transmitter and sensitivity as a receiver, by its piezoelectric transmitting d and receiving g
coefficients, as follows:
d = gεT
kp2 s E = dg,
(13)
(14)
where ε T is the free dielectric constant of the material, s E is its elastic compliance at constant
field and kp is its planar coupling coefficient.
In biomedical applications, the lead zirconate titanate (PZT) ceramic ferroelectric
materials have for many years been the most popular transducers and this continues to be
the case although composites of PZT and plastic polymers are beginning to be used. The
piezoelectric polymer material, polyvinylidene difluoride (PVDF), has some advantages,
particularly as a high frequency receiver.
Table 2 gives values for important electromechanical properties of some materials which
are used as transducers in biomedical applications. Most commonly, transducers are operated
either at resonant frequency or over a band of frequencies containing the resonant frequency.
For fundamental frequency operation of a lead zirconate titanate transducer in the thickness
mode, λ/2 ≈ 2 mm at 1 MHz (and proportionately less at higher frequencies). In summary,
the data in table 2 show that PZT4 is a good transducer material for ultrasonic power generation
Ultrasonic imaging of the human body
679
Table 2. Electromechanical properties of some piezoelectric materials. Data from Bowen et al
(1996) and Hadjicostis et al (1984).
Material
Electromechanical property
Coupling coefficient, kp
Charge constant, d33
(×10−12 m V−1 )
Voltage constant, g33
(×10−3 V m N−1 )
Dielectric constant, ε T
(×10−11 F m−1 )
Elastic compliance, S E
(×10−12 m2 N−1 )
Mechanical Q
Density, ρ
(kg m−3 )
Wave velocity, c
(m s−1 )
Characteristic impedance, Z
(×106 kg m−2 s−1 )
Curie temperature, Tc
(◦ C)
PZT4
0.57
315
PZT5
0.60
374
PVDF
0.20
35
PZT/polymer 1-3
0.63
650
24.6
24.8
152
66
1150
1500
10
460
24
26
133
108
600
7600
75
7700
10
1400
20
1800
4000
3760
3000
3330
30.4
29.0
4.2
6.0
320
365
130
350
but, because of its high Q, it is not as good as PZT5 in pulse operation. PVDF is apparently
a good receiving transducer, but its low dielectric constant and its low d33 value mean that
its sensitivity is greatly reduced by, for example, connecting cable capacitance, and it is
an inefficient transmitter. The great advantage of composite transducer materials, such as
PZT/polymer 1-3, is that they have a relatively low characteristic impedance (which is also an
advantage of PVDF), whilst retaining a value of coupling coefficient which actually may even
be somewhat higher than that of PZT.
3.2. Probe construction: elementary considerations
In medical ultrasound, a device that contains a transducer is often referred to as a ‘probe’,
since it is used to generate and detect the ultrasound that probes the internal structures of
the body. As explained in section 2.2, a simple kind of transducer is in the shape of a disc of
piezoelectric material: the diameter, in terms of the wavelength, determines the geometry of the
ultrasonic beam. Unless its behaviour is modified by matching and backing (see section 3.3),
such a transducer has its fundamental resonance at the frequency at which its thickness is
half a wavelength; for PZT, this is about 2 mm at 1 MHz and proportionately less at higher
frequencies. The sharpness of the resonance is described by the Q of the transducer material,
so that, for short pulse operation, a low Q is desirable.
3.3. Matching, backing and loading: pulse operation
The instantaneous particle pressure (p) in a medium supporting the propagation of a wave is
given by
p = ρcv0 sin ωt,
(15)
where v0 is the peak particle velocity, ω = 2πf and t is the time. The derivation of equation (5)
depends on the continuities of both particle velocity and particle pressure across the interface
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P N T Wells
between two media. It follows from equation (5) that
It /Ii = 4Z2 Z1 /(Z2 + Z1 )2 ,
(16)
where It is the intensity of the wave transmitted into the second medium. It is usually the case,
in medical ultrasonics, that the transducer has a different (higher) characteristic impedance than
that of the medium (water or soft tissue) in which the ultrasound is propagated. Substitution
in equation (16) shows that, with a PZT transducer, there is about 20% transmission from
the transducer into the load. If a lossless layer of thickness l, characteristic impedance Z3
and wavelength λ is included between the transducer and the load, however, the transmission
across the layer is given by
It /Ii = 4Z2 Z1 [(Z2 + Z1 )2 cos2 kl + (Z1 + Z2 Z1 /Z3 )2 sin2 kl],
(17)
where k = 2π/λ. Then, if l = nλ/4, where n is an odd integer,It /Ii = 1 (i.e., there is 100%
transmission). Thus, the impedance of the transducer can be matched to that of the load by
means of a λ/4 layer of a material of intermediate impedance (typically 7 × 106 kg m−2 s−1
to match PZT to water). At higher frequencies, a quarter-wavelength layer may be too thin
to be reliable, in which case a 3λ/4 layer may be used. A particularly neat solution to the
problem of fabricating a quarter-wavelength matching layer is to integrate the layer with the
front surface of the transducer. Seyed-Bolorforosh (1996) has described how this can be done
by cutting tiny orthogonal grooves on the front of the transducer so that little pillars of PZT of
the desired height stand proud. After electroding the surface of the transducer at the base of the
grooves, epoxy resin is poured over the structure to fill them. The characteristic impedance of
the integrated composite material thus formed can be designed to have the appropriate value,
by proper choice of the dimensions of the pillars.
For continuous wave operation, a transducer with a high Q is desirable. This means that
its resonant frequency is well-defined and the thickness of the quarter wavelength matching
layer can be calculated precisely. Because the device operates at resonance, however, it is
usually satisfactory for a half-wavelength layer to be used between the transducer and the load.
According to equation (17), the characteristic impedance of such a layer does not affect the
transmission; high efficiency is obtained provided that the attenuation in the layer (and in the
transducer) is low. A quarter-wavelength matching layer does have an important advantage
with pulse operation. Because the frequency spectrum of a pulse covers a range that increases as
the pulse length becomes shorter, however, the thickness corresponding to a quarter wavelength
becomes increasingly ill-defined.
Besides matching the characteristic impedance of the transducer to that of the load, there
are two other factors that have an important influence on the performance of the probe. The
first is the backing, or rear surface loading, of the transducer. A simple approach, for short
pulse operation, is to back the transducer with a medium similar in characteristic impedance to
that of the transducer, and which absorbs any ultrasound that leaves the transducer from its rear
surface (Nguyen et al 1996). This has the effect of damping the resonance of the transducer
so that it is equally sensitive over a wide range of frequency. The penalty is that the sensitivity
is very significantly reduced. Good performance is obtained with impedance matching of
the front surface of the transducer with a quarter-wavelength plate and with an air-backed
rear surface, or, better still, with a rear half-wavelength plate tuned to a frequency which,
in conjunction with the resonant frequency of the front plate, extends the overall frequency
response of the device.
The other factor that influences the performance of the probe is the electrical circuit to
which it is connected. At resonant frequency, the transducer behaves like an inductor L and
a capacitor C connected in series (with a resistor R, also in series, to account for loss in the
device), with a second capacitor Cs connected in parallel. The series elements correspond to
Ultrasonic imaging of the human body
681
the mechanical behaviour of the device, the resonant frequency ω0 = 2πf0 being determined
by the equation ω0 = 1/(LC)1/2 . Also, Q = ω0 L = 1/ω0 CR. The effect of the capacitor
Cs is to reduce the sensitivity of the device. By connecting an inductor Ls in series with the
transducer, the total impedance of the reactive components loading the transducer becomes
equal to zero at resonant frequency, i.e., when ω0 = 1/(Ls Cs )1/2 .
It can now be seen that the probe designer can select the characteristics of the matching
and backing layers and the electrical loading of the transducer to optimize the performance
of the device. Generally, this means that the sensitivity is constant and has its maximum
practicable value over the desired frequency bandwidth. Such a device is also likely to have
the maximum practicable dynamic range: this means that the electrical ripple that immediately
follows the transmission or reception of a brief pulse of ultrasound has the minimum amplitude
and duration.
3.4. Beam studies
The pressure distribution of the ultrasonic beam produced by a transducer with any size
and shape of aperture, and with any temporal distribution, can be theoretically calculated.
For example, equations (9) and (10) describe the beam produced by a disc transducer with
continuous wave excitation. In practice, however, insufficient data are likely to be available to
allow a reliable estimate of beam shape to be calculated and it is necessary to measure or to
observe the spatial distribution of, for example, the pressure field. In addition, the integrated
beam power and the intensity distribution may be of interest. The most important of the
available measurement and observation methods are described in the following sub-sections.
3.4.1. Hydrophones. Hydrophones are detectors based on transducers that respond directly
to the ultrasonic field. The output of a hyrophone is an electrical signal that follows the
instantaneous value of the ultrasonic pressure, ideally over a small area.
A small area hydrophone needs to be small in relation to the wavelength of the ultrasound
that it is required to measure. This makes the device nondirectional. There are two main types
of construction. A needle or fibre hydrophone has a sensitive tip and can be used to probe
an ultrasonic field. A membrane hydrophone consists of a relatively large sheet of a thin film
of plastic piezoelectric material that has a negligible effect on the field; a tiny element of the
membrane is electroded to form a sensitive area.
A typical piezoelectric needle hydrophone consists of a polyvinylidene difluoride disc,
500 µm in diameter and 15 µm in thickness, attached to an insulating layer at the tip of a
600 µm diameter stainless steel tube (Lewin 1981). The needle lumen contains a backing
material that has a higher characteristic impedance than water, so that resonance occurs at the
frequency at which the transducer is λ/4 in thickness.
There are two main types of fibre-optic hydrophones. In one type, the ultrasonic pressure
wave modulates the refractive index of the fluid in front of the tip of the fibre, leading to a
change in reflectivity that can be measured with an optical detection system at the other end
of the fibre (Staudenraus and Eisenmenger 1993). A problem with this type of hydrophone is
its lack of sensitivity. The second type of fibre-optic hydrophone has a metal-coated tip and
operates as a baseband or heterodyne interferometer. In a variant of this approach, a robust
device, capable of withstanding shock waves, consists of a cut fibre with one or more hard
dielectric optical layers formed by sputtering. For example, a single-mode fibre with a core
diameter of 3.5 µm with a λ/4 layer at the tip has been constructed, with a bandwidth of
70 MHz (Koch 1996).
A typical membrane hydrophone consists of an annular frame with a diameter of 100 mm,
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P N T Wells
over which a 9 µm-thick sheet of polyvinylidene difluoride is stretched (Preston et al 1983).
This thickness corresponds to λ/2 at 170 MHz. At the centre of the disc of film, a sensitive
element is formed by vacuum-deposited gold electrodes with diameters in the range 0.2–
1.0 mm, with nonoverlapping connexion tracks. The membrane essentially has a negligible
effect on an ultrasonic field in the 1–15 MHz frequency range and the sensitivity is flat within
1 or 2 dB.
3.4.2. Target plotting. A method that is quite widely used for plotting the effective distribution
of an ultrasonic beam, provided that it can be pulsed to allow echoes to be detected, is to scan
a small target point-by-point within the field. For echo amplitude measurement, a null method
can be used with a calibrated attenuator appropriately positioned in the electrical signal path.
The choice of target has to recognize the compromise between the desirability of small size
(to minimize directionality) and the need for the echo to be detectable. Also, the reflectivity
of the target should vary with the ultrasonic frequency in a well-defined fashion. Lypacewicz
and Hill (1974) concluded that a small stainless steel ball bearing is an excellent choice: it is
self-aligning and nondirectional, and can be supported by a wire soldered to its rear surface.
3.4.3. Schlieren observation. The variation in the density of a medium supporting an
ultrasonic wave produces corresponding changes in the optical refractive index if the medium
is transparent. The schlieren method depends on this phenomenon. A parallel beam of light
is arranged to pass through a transparent medium (usually water) in which the ultrasonic
beam is travelling. The light is then focused onto an obstruction, so that none reaches the
observer when there is no ultrasonic field. When ultrasound changes the refractive index of
the medium, however, some of the light that passes through the disturbed area no longer falls
on the obstruction; that which is deviated constitutes a bright image of the ultrasonic beam.
Thus, the beam becomes visible against the dark background. A closed-circuit TV system can
be conveniently used for display.
Pulsing the light synchronously with the ultrasonic pulse generator results in stroboscopic
visualization of the field. If viewed via a TV camera, an indication of the beam profile can
be obtained by displaying the amplitude-time waveform of an appropriately selected TV scan
line (Follett 1986).
3.4.4. Magnetic resonance imaging. Nuclear magnetic resonance signals can be influenced
by motion within the material being studied by the application of a magnetic field gradient
superimposed on the static, uniform magnetic field that is used for spin polarization (Hahn
1960). Particle displacements of a few micrometres accompany the propagating waves
typically used for ultrasonic imaging. Ultrasonic waves in agar gel have been visualized
by a clinical MRI scanner with synchronized phase-locked gradient waveforms (Walker et al
1998). The method has the potential to enable the propagation of ultrasonic beams used for
imaging to be studied noninvasively in living tissues.
3.4.5. Power and intensity measurement. There are three main methods for measuring
the power transported in an ultrasonic beam. These are by calorimetry, radiation force
measurement, and calibrated hydrophone.
In calorimetry, the quantity of heat produced in unit time when an ultrasonic beam is
completely absorbed is equated to the ultrasonic beam power. Essentially, there are two
types of calorimeter. In a flow calorimeter, a liquid is continuously pumped through the
vessel in which the ultrasound is absorbed. The power is calculated from the difference in
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the temperatures of the outflow and inflow liquid, and the rate of flow of the liquid and its
specific heat. A convenient improvement is to place an electrical heating element between
the temperature measurement sites. The reduction in electrical power needed to maintain
a constant temperature difference when the ultrasound is switched on is then equal to the
ultrasonic power. This is independent of the flow rate and the specific heat, and the electrical
power can be controlled by a feedback loop. This type of calorimeter is relatively sensitive
and has a fast response (Torr and Watmough 1977).
In the second type of calorimeter, flowing liquid is not used to maintain a steady state.
Instead, the rate of rise of temperature of the device is measured whilst the ultrasound is being
absorbed. The water equivalent of the device needs to be known, for the ultrasonic power to
be calculated, or it must be calibrated against another power measurement method. Although
the initial rate of rise of temperature is only slightly affected by cooling to the environment, a
cooling correction needs to be applied if the heating leads to a significant temperature gradient.
An example of a calorimeter of this type, consisting of a hollow sphere filled with carbon
tetrachloride (the characteristic impedance of which is close to that of water, and which is
relatively absorbent) and containing an array of thermocouples has been described by Wells
et al (1963).
The importance of ensuring that the calorimeter responds only to heat generated by the
absorption of ultrasound, and not to that resulting from the inefficiency of the transducer, does
not seem to have been widely realized. Effective thermal isolation is necessary.
The relationship between the radiation force produced when an ultrasonic beam travelling
at a known velocity is absorbed, and the ultrasonic power transported by the beam, is given
by equation (12). For an ultrasonic wave with a power of 1 W travelling in water, a force of
670 µN is produced when the wave is completely absorbed. This is equivalent to the force
of gravity acting on a mass of 69 mg. The force is doubled if the beam is totally reflected at
normal incidence. Thus, the ultrasonic powers used in imaging can be measured by waterimmersed instruments based on, for example, modified analytical balances (Rooney 1973) or
electronically-controlled servo sensors (Farmery and Whittingham 1978).
There are several approaches to the measurement of ultrasonic intensity. For example, a
calibrated hydrophone may be used. The signal detected by a hydrophone is proportional to the
particle displacement amplitude. The intensity of the ultrasound is proportional to the square
of the displacement amplitude. Provided that the total beam power is known (for example, by
radiation force measurement), scanning the hydrophone point-by-point across the beam and
integrating the squares of the resultant voltage measurements provides the data necessary to
calibrate the hydrophone. Other methods of intensity measurement include the radiation force
on a small suspended spherical target (Hasegawa and Yoshioka 1969), the temperature rise
of a small thermocouple junction embedded in an absorbent disc (Fry and Fry 1954), and the
displacement of a thin reflective pellicle observed by laser interferometry (Vilkomerson et al
1977). Generally, methods that use piezoelectric or optical detectors are likely to be more
sensitive, and to have better time resolution, than those using thermal detectors. Therefore,
they are more likely to be suitable for measuring the relatively low time-averaged intensities
that are usually used in ultrasonic imaging.
4. Image formation
The first attempts to use ultrasound for medical diagnosis were based on the expectation that
it would be possible to demonstrate tissue masses within the body and, particularly, within
the brain, because of differences in attenuation. Dussik et al (1947) constructed a scanner
in which a beam of ultrasound was directed through the patient’s head and detected by a
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receiver placed in line with the transmitter. The images which were formed by scanning
the beam in a raster pattern seemed to represent the intracerebral structures, including the
ventricles. Using a similar scanner operating at a frequency of 2.5 MHz and an intensity of
about 1 W cm−2 , Hueter and Bolt (1951) concluded that ‘a preliminary evaluation indicates
that the echo-reflection method is considerably less promising (than the transmission method)
for general ventriculography, mainly because of the small amount of reflection at the interface
between the tissue and the ventricular fluid’. The subsequent demonstration (Ballantine et al
1954) that an empty skull gave rise to similar pictures, because of the coincidental transmission
properties of the bone, halted work on this approach and arguably held back progress in
ultrasonic imaging research for several years.
Pulse–echo ultrasound was shown to have practical value for the detection of flaws in
metals during World War II; the publications of Firestone (1946) in the USA and Desch et al
(1946) in the UK appeared as soon as the constraints of military secrecy were relaxed. Using
this same technique, research into medical applications soon began in Denver, where Howry
and Bliss (1952) constructed a water-immersion two-dimensional scanner, and in Minneapolis,
where Wild and Reid (1952) started to develop high-frequency real-time two-dimensional
imaging. From this early work, researchers world-wide and in increasing numbers began to
explore the potential of the new technique, although perhaps initially more slowly in the USA
than elsewhere, because of the set-back with transmission imaging.
4.1. Principles of pulse–echo ultrasound
The pulse–echo method depends on the measurement of the time that elapses between the
transmission of a pulse of ultrasound and the reception of its echo from a reflecting or scattering
target (from which the distance to the source of the echo can be calculated, if the propagation
speed is known), and the measurement of the amplitude of the echo (which is related to the
ultrasonic properties of the target). The spatial resolution is determined, in elevation and in
azimuth, by the cross sectional dimensions of the ultrasonic beam, and, in depth, by the duration
of the ultrasonic pulse. The maximum depth of penetration is that at which the amplitudes
of echoes are just detectable; this depends on the attenuation of ultrasound in the tissue,
which is itself dependent on the ultrasonic frequency. Ultimately, the resolution is limited by
diffraction. This means that spatial resolution improves as the wavelength is decreased (i.e.,
as the frequency is increased), but this has to be set against the consequential reduction in
penetration.
4.1.1. Line scanning. Ultrasonic pulse–echo information is generally acquired along an
ultrasonic beam and, whilst this is in progress, the beam has effectively to be in a fixed
spatial position. Thus, the pulse–echo wavetrain is a single line of information in which
time corresponds to depth and amplitude, to the reflectivity, or backscattering strength, of the
tissue structures along the beam. Displayed on, for example, a cathode ray oscilloscope, the
wavetrain is called an ‘A-scan’, using terminology originating in radar.
The essential components of an ultrasonic A-scope are illustrated in figure 3. In a typical
arrangement designed to penetrate 150 mm into the body, the ultrasonic centre frequency could
be 3 MHz and the transducer could have a diameter of 17 mm (see section 2.6). Range ambiguity
would arise with pulse repetition frequencies in excess of 5000 Hz (because ultrasound travels
300 mm in soft tissue in 200 µs). The pulse repetition frequency is determined by the rate
generator and, in practice, a frequency of 200 Hz would typically be used as this is high enough
to avoid display flicker and it seems prudent to minimize the exposure of tissue to ultrasound.
Echoes from deeper structures are increasingly attenuated by the intervening tissue; the swept
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Figure 3. Basic elements of the A-scope. The output from the receiver is connected to the vertical
(y) deflexion plates of the cathode ray tube, and that from the timebase generator, to the horizontal
(x) plates.
gain generator increases the amplification of the receiver, following the transmission of each
ultrasonic pulse, to compensate for this.
4.1.2. Two-dimensional scanning. The production of an image of a plane section through
the body can be accomplished by scanning the ultrasonic beam across the plane whilst pulse–
echo wavetrains are acquired. A two-dimensional image is formed by relating the positions
of registrations on the display to the positions of the corresponding echo-producing structures
within the patient, as illustrated in figure 4. This requires a means to determine the ultrasonic
beam position in the scan plane so that the two timebases of the display, one horizontal and
one vertical, produce a resultant image line with the appropriate position and orientation. The
brightness along this line is controlled by the amplitude of the corresponding echo signal.
Various methods can be used to control the direction of the ultrasonic beam and to scan
it through the patient. The first type of two-dimensional scanner, that came into widespread
clinical use in the mid-1960s, used a system of sliding or rotating linkages to constrain the
ultrasonic probe within a fixed plane, the orientation of which could be selected according
to the anatomical section to be imaged (Donald et al 1958, Holmes et al 1965, Wells 1966).
The configuration that turned out to be most popular was that devised by Wells (1966) and
used two articulated arms with a system of wires, pulleys and potentiometers to measure the
position and direction of the ultrasonic beam within the scan plane. The probe was moved by
hand across the patient’s skin, with a water-based gel to exclude air; the process of acquiring
a single image typically occupied 5–30 s. The scanners were generally designed to produce
tissue maps with emphasis on the display of organ boundaries. The best scans were considered
to be those in which the anatomy was depicted by thin white lines on a black background, for
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Figure 4. The principles of two-dimensional B-scanning. The timebase line on the display is
the resultant of horizontal and vertical timebases controlled by the scanner to correspond to the
orientation and position of the ultrasonic beam in the patient. The amplitudes of the received echoes
control the brightness of the display.
which purpose a black-and-white bistable display is ideal. The most accomplished operators
developed the skill to oscillate the probe through an angle whilst moving it around the body
(a process known as ‘compound scanning’), thus improving the chances of achieving normal
incidence with organ boundaries and so obtaining the strong echoes necessary to form good
images with the insensitive equipment that was then available.
In the mid-1970s, however, mainly because of the work of Kossoff (1972), it was realized
that the echo amplitude conveys useful diagnostic information and that grey-scale displays are
generally greatly superior to those limited to presenting black-and-white images. Nowadays,
grey-scale imaging is universally used. The grey-scale capability can be described in terms
of the dynamic range. This can be expressed as the separation (in dB) between the minimum
and maximum echo amplitudes over which changes in echo amplitude produce perceptible
changes in image brightness.
The introduction of grey-scale scanning was accelerated by the arrival of the scan
conversion memory tube. In this device, the image formed by the intensity-modulated timebase
corresponding to the ultrasonic beam position in the patient is stored as a charge pattern on an
insulated target within a cathode ray tube. The charge pattern is then read out by raster-scanning
the target in a TV-compatible format. This made it possible to replace the old-fashioned
photographic recording methods, and the bistable storage tube, with a convenient method of
grey-scale display. Previously sceptical doctors began to appreciate the potential of ultrasound
to provide clinically useful images that, with a little experience, could be understood by all.
4.1.3. Real-time scanning. The next major advance in clinical ultrasonic imaging came with
the development of real-time scanning. With a few exceptions, the scanners that were in clinical
use up to the mid-1980s required at least a few seconds to acquire an image. Apart from the
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687
limitation that motion, such as that of the heart, could not be directly observed, the process
of ultrasonic diagnosis consisted of making a scan, examining the image, deciding whether
another scan was needed, and so on until there was reasonable confidence in the result. The
process was revolutionized by the development of scanners that acquired grey-scale images
with frame rates of 15–20 s−1 and higher, producing flicker-free displays in real time, so that the
image immediately followed changes in scan plane orientation and physiological movement
could be observed.
The maximum pulse repetition rate for unambiguous pulse–echo signal acquisition
depends on the required depth of penetration into the patient. For example, it is 5000 Hz for
a maximum penetration of 150 mm (see section 4.1.1). This means that 5000 lines of image
information can be acquired per second. Provided that the beam can be scanned through the
tissue plane with sufficient speed, the product of the number of lines per frame and the image
frame rate is equal to the pulse repetition rate. Thus, in this example, real-time images each
consisting of 250 lines could be acquired at a frame rate of 20 s−1 , and so on.
The first practicable real-time scanners employed mechanical means to sweep the
ultrasonic beam through the tissue plane. In one system, two transducers were mounted
opposite each other on the rim of a rotating wheel, to produce radial beams; the wheel was at
the focus of a parabolic mirror in a water bath, so that a scan with a rectangular format could
be made through a flexible membrane forming one wall of the water bath, facing the mirror
and in contact with the patient’s skin (Pätzold et al 1970). Although it was cumbersome to
use and sometimes hard to obtain the desired scan plane, this scanner had the advantage of
not needing a scan converter (because the ultrasonic beam positions were all in parallel, so
the vertical timebase of the display needed only to be translated horizontally to provide image
registration). Being commercially available and with virtually no competition at the outset, it
led the field for several years.
Real-time scanning with a hand-held probe became possible with the introduction of
systems employing either an oscillating transducer scanning through an oil layer behind a
thin membrane in contact with the skin (McDicken et al 1974) or in direct contact with
the skin (Eggleton et al 1975, Schuette et al 1978), or a continuously-rotating wheel with
radially-mounted transducers, first in contact with the skin (Holm et al 1975) and, later, when
commercially available, within a liquid-filled casing. All these types of scanners steered the
ultrasonic beam through a sector and, although it was feasible, in principle, to generate the
appropriate timebase control directly, it turned out to be easier and more convenient to convert
the scan to a TV format. This was because digital scan converters with sufficient dynamic range
and resolution, and based on solid-state random access memories, had just become available at
economic prices. Nowadays these devices are universally used for scan conversion and image
storage, except in the very least expensive scanners.
4.2. Transducer array scanning
The methods of two-dimensional scanning that have so far been described all employ singleelement transducers. What this means is that the transducer, usually in the form of a disc, has
an aperture that is large enough to produce a directional beam with a near field long enough
to allow focusing to be used to optimize the resolution in azimuth and elevation. The beam
steering is carried out mechanically.
It was probably Buschman (1965) who first used an array of transducers to produce
an ultrasonic image. His probe had ten small transducers mounted on an arc-shaped support
designed to fit over the eye. The transducers were activated sequentially to produce a scan with
ten discrete lines of image information. Then, well ahead of his time, Somer (1968) described
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P N T Wells
the first phased array real-time two-dimensional sector scanner. This was a remarkable
achievement. A 10 mm × 11 mm, 1.3 MHz, 21 element transducer array was constructed,
together with the electronic circuitry required to steer the transmitted beam through a sector by
introducing the appropriate time delays in the impulses applied to each element in the array.
In this first instrument, there was no provision to steer the received beam. The idea was to
use the central transducer element in the array as a nondirectional receiver. Somer (1968)
was working with neurologists who wanted to produce images of the brain through the intact
skull. It was probably mainly because of the unfavourable ultrasonic properties of the skull
(reminiscent of the problems explained in the early work described at the beginning of this
section) that the results were disappointing.
In what can now be seen to be a natural extension of Buschman’s (1965) array of
transducers, Bom et al (1971, 1973) constructed linear arrays with 20 transducer elements.
Overall, the probe face was 80 mm long and 10 mm wide. Operating at a nominal frequency
of 3 or 4.5 MHz according to the construction of the probe, each element had a diameter of
3 mm. The length of the near field of the beam of each element was only 4.5 mm at 3 MHz, and
the half-angle of divergence was 12◦ , so the imaging characteristics were far from satisfactory.
They were the best that were practicable, however, because the compromise was between
resolution and image line density: an image with only 20 lines was considered to be as sparse
as could be tolerated. The advantages of the system were the simplicity of the display (each
element acquired a pulse–echo wavetrain in sequence, and the horizontal timebase was shifted
vertically according to the element that was activated) and its rapid frame rate (150 s−1 ). The
instrument was vigorously promoted commercially and the potential of real-time imaging to
revolutionize cardiological diagnosis became apparent through its use. Only a small number
of enthusiasts acquired instruments, however, because the system was inadequate for routine
clinical use.
This type of array of single-element transducers is really quite different from what is
nowadays understood by a transducer array. A modern transducer array allows the size of the
aperture to be selected and the ultrasonic beam to be focused and steered. Figure 5 shows the
principles of beam focusing and steering with an array. In this diagram, only four elements
are represented, to simplify the description: a real aperture would typically have at least 16
elements. The important point is that each individual element is narrow enough effectively
to be nondirectional in the scan plane. As viewed in the diagram, each transducer element
emits a cylindrical wavelet in response to electrical excitation. When all the elements in the
aperture are excited simultaneously (figure 5(a)), the wavelets combine to form a wavefront
parallel to the aperture, so that the beam travels directly away from the array. With a linear
timing excitation gradient across the array (figure 5(b)), however, the beam is steered in a
direction the angle of which to the normal can be changed by changing the direction and slope
of the gradient. In figure 5(c), the situation that arises when the distribution of the timing at
excitation across the array is cylindrical, is that the corresponding ultrasonic wavefront is also
cylindrical: this means that the beam is brought to a focus at the centre of the cylinder. Finally,
both the direction and focal length of the beam can be controlled by simultaneously changing
both the gradient and cylindrical radius of the excitation (figure 5(d)).
This explanation of beam control is in the context of the formation of a transmitted beam.
The same control of a received beam can be provided by delay lines in the individual signal
paths associated with each element in the array, prior to summing the signals that have passed
through the delay lines.
Figure 6 shows a linear array consisting of a large number of tiny elements. A typical
modern array might have 128 elements, each 0.5 mm wide and 7.5 mm long, extending over
a distance of 75 mm. In this example, an aperture with a width of 10 mm can be formed by
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689
Figure 5. Principles of beam forming and steering with a transducer consisting of an array of narrow
elements, illustrated with four such elements. The same principles apply both on transmission and
on reception. (a) Simultaneous excitation produces a beam normal to the array. (b) Linear timegraded excitation steers the beam away from the normal. (c) Cylindrical time-graded excitation
focuses the beam. (d) Superimposed linear and cylindrical time-graded excitation steers and focuses
the beam.
utilising 17 elements in a group. By stepping along the array (Whittingham 1976), one element
at a time, there are 111 discrete beam positions along the array. Even more beam positions
can be formed, if the aperture size is altered alternately by one or two elements during the
stepping process. Beam focusing in elevation is provided by the cylindrical lens. (Note that
suitable lens materials (i.e., materials that have characteristic impedances similar to that of
water or tissue, but different propagation speeds) usually have higher propagation speeds than
water or tissue. Consequently, a focusing transducer has a concave section.) This means that
focusing in elevation has to be at a fixed depth, both on transmission and reception. The same
applies to focusing in azimuth of the transmitted beam: the focusing conditions cannot be
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P N T Wells
Figure 6. Linear transducer array with electronically-controlled focusing in azimuth and lens
focusing in elevation. In this example, there are 12 elements in the active aperture (shown by
stippling); there are 51 elements in the array, giving 42 separate lines in the image.
changed after the beam has left the aperture. To give an idea of the magnitude of the delay
required to focus a beam, a simple trigonometrical calculation shows that the outermost limit
of a 10 mm wide aperture has to be excited 170 ns before the centre of the aperture in order to
focus the beam at a depth of 50 mm. On reception, however, the focal length in azimuth can be
swept continuously to coincide with the instantaneous position of the echo-producing targets,
by dynamically adjusting the delays in the received signal paths from each of the transducer
elements within the aperture. The time delays may be introduced either by analogue or by
digital circuits. Although analogue delay lines have largely been superseded by digital circuits,
it is of interest to note that three different analogue approaches have been adopted. In one
approach, the signal is transmitted along a tapped coaxial cable, and a switch extracts the
signal from the tap that happens to provide the desired delay. Alternatively, the delay can be
provided by a tapped series of inductor–capacitor elements. Finally, the capacitative element
in a single inductor–capacitor circuit can be under voltage control, for example, by changing
the depletion layer thickness in a semiconductor diode.
Analogue delay lines have largely been replaced by digital techniques, now that fast
sampling and processing speeds with sufficient dynamic range are readily available. At a
nominal frequency of 5 MHz, for example, digital sampling at a frequency of 50 MHz is
adequate to process a typical ultrasonic pulse. After analogue preprocessing of the received
echo signals, a dynamic range of 50 dB is likely to be sufficient. Digitization to 8 bits
corresponds to 256 levels or 48 dB. The signal having been satisfactorily digitized at 50 MHz,
a time shift of 20 ns can be introduced by shifting the waveform by one sample period. This
is generally adequate for beam focusing and steering.
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Figure 7.
Examples of typical probes employing
transducer arrays. Left-to-right: a phased array probe
for sector scanning; an endovaginal probe; a large
curvilinear array for general purpose abdominal scanning;
and a smaller curvilinear array for scanning relatively
superficial structures. The sizes can be gauged from the
15 mm diameter of the endovaginal probe.
The ultrasonic beam from an array can be steered through an angle by introducing a linear
time gradient along the aperture. For example, to steer the beam from a 10 mm aperture
through an angle of 45◦ , the time difference across the aperture is 4.7 µs. If the beam is
both to be focused and steered, a cylindrical timing profile needs to be superimposed on the
linear gradient. Although a linear array can be operated in this way, usually to produce a
scan with the format of a parallelogram, beam steering is more commonly used over the entire
aperture of what is commonly called a ‘phased array’. This produces a sector scan format.
A phased array typically has 64 elements in a 15 mm aperture, each element being 200 µm
wide and 10 mm long. The manufacture of such an array, with a separate electrical connexion
to each element and, usually, multiplexing electronic circuits within the probe casing, and the
specialized multicoaxial connecting cable, requires a high level of precision engineering.
The characteristic impedance of a ceramic transducer is substantially different from that of
water or soft tissues (see table 2). This reduces the sensitivity of the transducer (i.e., increases
its insertion loss). A quarter-wave matching layer is used (see section 3.3) to improve the
sensitivity of the system and, together with a matching layer on the rear surface of the transducer,
the characteristics can be tuned to provide optimal sensitivity over a wide frequency band.
Examples of typical probes employing transducer arrays are shown in figure 7.
4.3. Signal processing and display for grey-scale pulse–echo imaging
The essential components of the signal processing chain for array scanning are shown in
figure 8. The ultrasonic pulse is usually generated by applying a brief (typically 10 ns)
monopolar signal, of around 100 V in amplitude, to the transducer (or to each transducer
element in the array). The nominal frequency of the ultrasound is determined by the resonance
of the transducer and its impedance matching layers. In figure 8, the ultrasonic beam is
steered and focused, both on transmission (by the multiple time-controlled transmitters) and
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Figure 8. The essential components of the signal processing chain for scanning with a transducer
array.
on reception (by the multiple controlled delay lines), as determined by the beam former. The
voltage produced by the transducer, in response to echoes from tissues within the patient, is
typically in the range 500 mV to 50 µV; i.e., it covers a dynamic range of 80 dB. The nonlinear
preamplifiers compress this dynamic range to 50 dB, so that the amplitudes of the signals fed
to the multiple controlled delay lines (whether analogue or digital) cover the voltage range
1 V to about 3 mV. The swept gain amplifier is able to compensate for substantially more than
a range of 50 dB of tissue attenuation (because of the action of the nonlinear preamplifiers),
so the voltage range fed to the demodulator is around 5 V to 150 mV. Although the swept
gain time function is usually selected by the operator, an ingenious adaptive technique may be
used that is based on the assumption that local attenuation is correlated with local backscatter
(Hughes and Duck 1997). The pulse-shaping video amplifier has a frequency response that
optimizes the appearance of the image and produces an output voltage suited to the type of
display (e.g., cathode ray tube or liquid crystal).
4.4. Resolution
The spatial resolution depends on the profiles of the ultrasonic beam and pulse (see section 3.4)
and the characteristics of the signal processing and display system (see section 4.3). Typically,
the spatial resolution in elevation is likely to be about three times worse than in azimuth. A
transducer array can be used either as a simple aperture, in which all the elements are active
throughout the signal acquisition process, or the size of the aperture can be adjusted to optimize
the spatial resolution throughout the depth of penetration. In designing an imaging system,
an obvious criterion is to maintain a constant f -number (focal length/diameter), independent
of the axial position of the target. This can be achieved by expanding the effective size of the
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aperture (i.e., by increasing the number of elements active in the array) with time following
the transmission of the pulse, so that the receiving beam width at the focus remains constant
(Harris et al 1991). In addition to this, there are several other advantages with this approach.
Because of tissue inhomogeneity (see section 2.7), there is a maximum aperture size beyond
which there is no further improvement in spatial resolution; indeed, the resolution may actually
become worse (Mosfeghi and Waag 1988). Although techniques have been tried to compensate
for the effects of tissue inhomogeneity with both one-dimensional (Smith et al 1986) and twodimensional (Ries and Smith 1995) arrays, none has proved to be easily implementable.
The temporal resolution of an ultrasonic imaging system is limited by the rate at which
image frames of adequate quality can be acquired. The discussion presented in section 4.1.3
applies to the situation in which only one ultrasonic beam is active at any particular time. The
image frame rate can be increased, however, by parallel processing, although usually at the
expense of some image degradation. For example, Shattuck et al (1984) used a broadened
transmitted beam within which four narrow received beams could be simultaneously formed.
In this case, the cost was the reduction in both sensitivity and spatial resolution. In principle,
it would be possible, if the array were sufficiently long, simultaneously to employ two entirely
separate pulse–echo beams; the cost would be an increase in noise due to crosstalk, which
would be evident as a reduction in image contrast resolution.
The contrast resolution of an ultrasonic imaging system is a measure of its ability to register
perceptibly different display brightnesses from targets of minimally differing reflectivities. In
a perfect imaging system, the brightness transfer characteristic could be selected to allow even
the minimal change in target reflectivity to produce a perceptible brightness change. The
display of a real imaging system contains ‘noise’, however, including that generated in its
electronic circuits. This noise reduces the contrast resolution. Another important source of
noise is due to the finite size of the ultrasonic beam and pulse and, particularly, to the side
lobes of the beam which are due to the finite size of the aperture and the grating lobes which
accompany a beam formed by an array of transducer elements (von Ramm and Smith 1983).
For a given number of elements in the aperture, the grating lobes tend to increase in amplitude
as the beam is steered away from the perpendicular direction. The effect of the finite beam
width and the ancillary lobes is that echo-producing targets lying away from the central axis
of the beam may give rise to signals that reduce the contrast resolution of the image.
4.5. Speckle
In pulse–echo ultrasonic imaging, it is the backscattered waves that provide the diagnostic
information. Although in reality the situation is complicated, blood is a tissue that can be
considered to consist of isotropic Rayleigh scatterers (see section 2.1). The backscattering
increases with the fourth power of the frequency. Of course, attenuation in intervening tissue
also increases with frequency, so the frequency that gives the maximum echo amplitude from
blood is determined by the combination of these two effects (Reid and Baker 1971).
Two-dimensional ultrasonic images of blood and the fine structure of soft tissues are
actually speckle patterns (Wells and Halliwell 1981, Wagner et al 1983). The ultrasonic pulse
occupies a volume of tissue that contains some number of individual scatterers of varying
strength and position; the amplitude of the corresponding electrical echo signal from the
receiving transducer is the result of interference between the scattered waves, each of which has
its own particular phase angle (O’Donnell 1983, Finette 1987). In many tissues, the scattering
is primarily due to collagen. Although the correlation function for tissue has not yet been
determined, the Gaussian model of scattering has so far provided a consistent description of
tissue structure (Insana et al 1990). Initially, as the frequency is increased, the backscattered
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echo amplitude increases towards a maximum value but, at higher frequencies, increasing
attenuation in the intervening tissue dominates and the echo amplitude falls.
Because it does not have a one-to-one correspondence with scatterers in the tissue, speckle
is sometimes thought merely to be an annoying image artifact, in the same category as noise.
More importantly, speckle reduces target detectability; strictly speaking, it does not affect
contrast resolution although it does reduce the useful spatial resolution. Speckle can be
reduced by summing uncorrelated images of the scan plane. The requisite uncorrelated images
can be obtained either by scanning from several different directions, by scanning at several
different ultrasonic frequencies (Gehlbach and Sommer 1987), or by collecting scans from
a fixed position while small physiological movements result in differing speckle patterns in
images with essentially the same anatomical information (Wells and Halliwell 1981). In some
situations, none of these methods may be practicable. If this is the case, speckle may be
reduced by adaptively filtering the image in the signal processing circuits (Bamber and Daft
1986, Chen et al 1996).
Although it might be supposed that speckle suppression would lead to an improvement in
image perception, this may not necessarily be so. The speckle in ultrasonic images is not fully
developed and its texture is influenced by larger-than-Rayleigh scatterers. For this reason,
the image textures from different tissues may have differing appearances, which can assist in
clinical image interpretation.
4.6. Examples of real-time grey-scale scanning
Figure 9 shows a large curvilinear transducer array being used to scan a pregnant woman and
figure 10 is typical of the kind of image that is produced.
Ultrasonic imaging has an important place in almost every area of clinical investigation.
For more information, reference should be made to the numerous textbooks that are available:
a good starting point is that edited by McGahan and Goldberg (1997).
4.7. Blood flow and tissue motion imaging
4.7.1. Ultrasonic scattering by blood. At the typical ultrasonic frequency of 3 MHz, the
wavelength in blood is about 500 µm. An individual red blood cell is a biconcave disc, with
a diameter of about 8 µm and a thickness of about 2 µm. Scattering of ultrasound by blood
can be modelled in several ways. For example, Brody and Meindl (1974) treated blood as a
suspension of point scatterers. In fact, however, blood cells are quite closely packed and so
the individual cells do not behave like uncorrelated scatterers but actually interact strongly.
This problem was avoided by Angelsen (1980), who modelled blood as a continuous medium
with fluctuations in density and compressibility. Human blood cells have a tendency to form
clumps, or ‘rouleaux’, which can survive even under normal flow conditions (Machi et al
1983). Nevertheless, scattering tends to decrease with increasing shear rate (Yuan and Shung
1989). With this as a model, Mo and Cobbold (1986) concluded that the backscattering of
blood can be considered to be a Gaussian random process.
If blood really did consist of a suspension of uncorrelated point scatterers and the
ultrasonic detection process was incoherent, it is arguable that blood flow could not be
detected by ultrasound. The ultrasonic power backscattered by the blood would remain
constant and there would be no discrete targets whose motion would either give rise to
a Doppler shift frequency or whose displacement could be observed over time. In fact,
however, scattering by blood is an example of the process that gives rise to speckle. Blood
behaves as an array of ensembles that give rise to fluctuations in backscattered power that fade
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Figure 9. An ultrasonic scanner being used for an obstetrical investigation. The probe is the large
curvilinear array shown in figure 7.
sufficiently slowly to allow enough time for their motion to be observed (Atkinson and Berry
1974).
4.7.2. The continuous wave Doppler method. Consider a beam of ultrasonic waves of
constant frequency and amplitude travelling through the body and encountering a vessel
containing flowing blood. The ultrasound detected as the result of backscattering from
stationary tissues has the same frequency as that of the transmitted ultrasound; that
backscattered by the flowing blood has its frequency shifted by the Doppler effect (see
section 2.3). If the same transducer was used both to transmit the ultrasound (for which
a 10 V signal would be likely to be required) and to receive the backscattered ultrasound
from the flowing blood (producing a signal of about 10 µV), the receiver would have to
accommodate a dynamic range of about 120 dB. This would be very difficult to achieve.
Therefore, with continuous wave ultrasonic Doppler systems, it is usual for separate transducers
(usually mounted side-by-side in the same probe) to be used for transmitting and receiving.
Typically, the receiver then has only to accommodate a dynamic range of about 60 dB. A signal
corresponding to the Doppler shifted echoes is obtained by multiplying the transmitted and
received signals, which can conveniently be done with a diode detector following the ultrasonic
frequency amplifier connected to the receiving transducer.
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Figure 10. Scan of a fetus in late pregnancy, made using a large curvilinear array of the type shown
in figures 7 and 9. In this longitudinal scan, the mother’s head is to the left. The uterine cavity
occupies most of the image: the fetal head is on the right, and the fetal body and one of the fetal
legs and feet can be seen.
4.7.3. The pulsed Doppler method. The continuous wave Doppler method (see section 4.7.2)
provides no explicit information about the distance between the ultrasonic transducer and the
moving target. This information can be provided, however, by combining the pulse–echo
principle (see section 4.1.1) with the Doppler method of motion detection (Wells 1969). The
principles are illustrated in figure 11. The master oscillator runs continuously at the frequency
of the ultrasound to be transmitted. The clock pulses are obtained by dividing down from
the master oscillator frequency (to maintain phase coherence). These clock pulses trigger
the transmit sample-length monostable at the pulse repetition frequency of the system. This
monostable generates a pulse that opens the transmit pulse gate for the period that ultrasound
needs to be emitted by the transducer to produce the desired sample length. With an ultrasonic
speed of 1500 m s−1 , this corresponds to 0.67 µs mm−1 . Echoes received by the transducer are
amplified and fed to the phase quadrature detector (Nippa et al 1975). This circuit separates the
signals according to whether they have frequencies higher (i.e., with leading phase) or lower
(i.e., with lagging phase) than that of the transmitted ultrasound. Higher frequency signals
correspond to flow towards the transducer, and vice versa for lower frequency signals. The
forward and reverse flow signals from the phase quadrature detector are fed to a heterodyne
processor, where they are mixed with a signal at a pilot frequency. The pilot frequency is
chosen to be greater than the highest reverse flow frequency signal. In the output from the
heterodyne processor, the pilot frequency corresponds to zero flow velocity; reverse flow
signals have frequencies lower than the pilot frequency, whereas forward flow signals have
higher frequencies. This arrangement makes it simple subsequently to analyse the signals in
terms of their frequency spectrum. Following the transmission of each ultrasonic pulse, the
output from the heterodyne processor consists of a wavetrain in which later time corresponds
to greater depth. The sample depth monostable, which is also triggered by the clock pulse,
introduces a time delay chosen by the operator to correspond to the echo delay time (about
Ultrasonic imaging of the human body
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Figure 11. Block diagram of a pulsed Doppler system providing depth and length adjustment of
the sample volume and phase quadrature detection of blood flow direction.
1.33 µs mm−1 , because the pulse has to travel to the target and the echo has to return) from
the beginning of the sample volume from which it is desired to collect the Doppler signals.
The receive sample length monostable then opens the receive pulse gate for a period equal
to the transmitted pulse duration, so that the output from the receive pulse gate is actually
the corresponding sample of what is, in effect, the Doppler signal. The amplitude of this
signal is held in the sample-and-hold circuit until it is updated by the sample derived from the
next transmitted ultrasonic pulse. The output from the sample-and-hold circuit is smoothed by
filtering, subjected to audio amplification and fed to the frequency spectrum analyser (typically
based on the fast Fourier transform method) and displayed (e.g., on a strip chart).
4.7.4. Doppler flow and motion imaging. The purpose of methods of flow and motion imaging
is to produce a two-dimensional (or three-dimensional) image of an anatomical structure, with
the presence of blood flow (or tissue motion) indicated in its correct spatial position by some
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form of image coding (usually colour). The code can carry information about characteristics
of flow or motion (e.g., its speed, direction, velocity or quantity). The clinical value of the
method may be greatly enhanced if the process can proceed in real time or, at least, if it can be
fast enough to follow physiological motion or changes in the position of the scanning probe.
The simplest way to produce an image of blood flowing in a vessel system is to make
use of the fact that Doppler-shifted blood flow signals are detected only when the ultrasonic
beam passes through flowing blood. A continuous wave Doppler probe (see section 4.7.2) is
mounted on a scanning system that measures the position of the probe in what is essentially
a two-dimensional plane, with the ultrasonic beam at an angle to the plane. The probe is
scanned slowly by hand, in contact with the patient’s skin under which lie the blood vessels to
be imaged. Only where Doppler signals are detected does the instrument cause registrations
to appear on the display (Reid and Spencer 1972). The same principle can be extended to
three-dimensional imaging by using a pulsed Doppler system with a multiplicity of receiving
channels gated over a range of depths, with the probe mounted on a scanner providing threedimensional measurement (Fish 1981).
The first two-dimensional blood flow imager (Reid and Spencer 1972) produced simple
black-and-white maps, with blood flow shown as being either present or not present. It was
Curry and White (1978) who first produced images coded in colour to show the velocity of
blood flow. Their equipment used a logic circuit to exclude reverse-flow signals and three filters
to separate forward-flow signals into normal, moderately increased and markedly increased
frequency ranges. By this means, the image was coded in red, yellow and blue. This allowed
the extent of localized narrowing of arteries, with associated increase in blood flow velocity,
to be assessed. In Fish’s (1981) three-dimensional scanner, the direction of blood flow was
indicated by colour coding, with blue for flow towards the transducer and red, for flow away.
This logical choice of colour corresponds to the colours of the light from stars, as originally
hypothesized by Doppler (1843).
Although manually-scanned ultrasonic Doppler blood flow imaging opened up a new area
of clinical investigation, it never came into widespread use. To use the technique successfully,
a great deal of skill was needed. Also, as the technique was being developed and assessed,
ultrasonic duplex scanning was introduced by Barber et al (1974). A duplex scanner is one
that enables two-dimensional ultrasonic pulse–echo imaging to be used to guide the placement
of an ultrasonic beam for Doppler signal acquisition, thus to identify the anatomical location
of the origin of the Doppler signal. Duplex scanners can be based on either mechanical
(see section 4.1.3) or transducer array (see section 4.2) real-time scanning systems. For the
acquisition of Doppler signals, the ultrasonic beam has effectively to be stationary and to dwell
in the appropriate position sufficiently frequently and for a sufficiently long time during the
acquisition period to allow adequate sampling of the target motion. The moving parts of a
mechanical scanner cannot be stopped and started rapidly and so only a stored image is usually
available during Doppler signal acquisition. This can be a problem if there is movement of
the patient or the probe. Scanning of the beam of an array scanner can be stopped and started
instantaneously, so simultaneous operation is possible with this type of system.
The arrival of the duplex scanner had a considerable impact on the practice of ultrasonic
imaging. It opened up important new areas of clinical investigation, particularly in the heart and
the vascular system. It had none of the inconvenience of the manually-scanned flow imaging
techniques and these largely fell into disuse. An example of a duplex real-time grey-scale scan
and the simultaneously-acquired Doppler frequency spectrum is shown in figure 12.
The basic principles of combining pulse–echo and Doppler two-dimensional images are
illustrated in figure 13. Essentially, the idea is an extension of duplex scanning but with
multiple Doppler receiving channels. Using a modified duplex scanner, Brandestini and
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Figure 12. Duplex scan of a carotid artery, one of the main arteries supplying blood to the head,
which is on the left of the grey-scale image. The scan was made at the region in the neck where the
artery bifurcates into the external carotid, which supplies the face, and the internal carotid, which
supplies the brain. The oblique line represents the direction of the ultrasonic beam selected for the
acquisition of the Doppler signals from the sample volume, indicated by the broad bar. The marker
passing through the sample volume was adjusted by the operator to align with the artery, to enable
the machine to estimate blood flow velocity from Doppler shift frequency: see equation (11). The
Doppler frequency spectrum is displayed on the right. The vertical scale indicates the blood flow
velocity (m s−1 ); the large markers on the horizontal scale indicate time in 1 s intervals. Note that
blood flow does not reverse direction in the carotid artery during diastole.
Forster (1978) obtained a real-time two-dimensional pulse–echo image through which they
swept the ultrasonic beam of a 128-point pulsed Doppler system to superimpose blood flow
signals, colour coded in red and blue, on the grey scale anatomical image. The system was
developed until it was eventually possible to operate at four frames per second (Eyers et al
1981). What ultimately limited the speed was the process of Doppler frequency estimation
that was employed.
It was the publication in English of a full paper (Kasai et al 1985) that drew widespread
attention to the results of work that evidently had been in progress in Japan for several years. It
showed that autocorrelation detection provided a rapid means of frequency estimation and made
it possible for colour flow imaging to be performed in real time. The basis of the autocorrelation
detector is that the echo wavetrains from stationary targets do not change with time, whereas
sequential echo wavetrains from moving targets have corresponding changes in relative phase.
As shown in figure 14, the autocorrelation detector produces an output signal that depends on
the relative phases of consecutive pairs of received echo wavetrains. Thus, the echo wavetrains
themselves are their own references for phase comparison. The autocorrelation detector
functions by multiplying two echo wavetrains, one currently being received by the transducer
and the other, having been received from the immediately preceding pulse transmission and
delayed for a time exactly equal to the interval between pulse transmissions. The output from
the autocorrelator has constant amplitude except where consecutive wavetrains have phase
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Figure 13. Block diagram of a real-time two-dimensional colour flow imaging system. In this
example, a phased array transducer is used to produce a sector scan. (Alternatively, a linear array
transducer can be used to produce a scan in parallelogram format.) The image formatter and
scan converter accepts the colour and grey-scale signals together with scan position data from the
ultrasonic beam steering circuit. The operation of the system is synchronized by a clock pulse
generator (for simplicity, this is not shown).
Figure 14. Block diagram of an autocorrelation detector and associated gating, scan-converting
and colour-processing circuits. The multiplier performs the process of autocorrelation.
differences. In figure 14, the separate processes of velocity and velocity variance calculation
are indicated; the value of the velocity variance can be considered to be a measure of the width
of the Doppler frequency spectrum, which increases with the degree of flow disturbance.
The colour processor in figure 14 assigns luminance, hue and saturation to the display,
following one of the schemes described in section 4.7.6.
4.7.5. Flow and motion imaging by time-domain processing. Doppler flow and motion
imaging, as described in section 4.7.3, is a narrow frequency band method in which processing
is performed in the frequency domain. Embree and O’Brien (1985) and Bonnefous and Pesque
(1986) independently described how flow and motion information can also be obtained from
broadband ultrasonic echoes by means of time-domain processing. The method depends on
temporal tracking of the spatial position of individual coherent blood ensembles or tissue
constituents. In principle, it can be applied directly to the amplified ultrasonic signals or
to the video signals as they appear on the display, although the implementations of the two
approaches are quite different.
Time-domain processing of the ultrasonic signals enables blood flow (or moving tissue)
velocity to be determined by a one-dimensional correlation between the echo wavetrains
acquired with consecutively transmitted ultrasonic pulses. The time t required for an ultrasonic
pulse to complete the round trip between the transducer and a moving scatterer situated a
Ultrasonic imaging of the human body
701
distance z from the transducer is given by
t = 2z/c.
(18)
If the interval between pulses (which is equal to the reciprocal of the pulse repetition
frequency) is equal to T and v is the speed of the scatterer, then
v = −τ c/2T ,
(19)
where τ is the change in the time of arrival of the echo from the scatterer between consecutive
ultrasonic pulses. The time-domain formulation is clearly similar to the frequency-domain
(Doppler) formulation given in equation (11) and would need to include the cosine function in
the same way if the scatterer was crossing the ultrasonic beam at an angle to its axis. The cross
correlation (the function that is performed by the multiplier in the time-domain realization of
figure 14) can be achieved by shifting the relative time positions of consecutive wavetrains;
when the product has its maximum value, the two wavetrains are shifted by (T + τ ) and, since
T is known, τ can be estimated. The other principal difference between frequency- and timedomain realizations of the circuit in figure 14 is that, for time-domain processing, the digital
sampling frequency needs to be increased by a factor of about ten to provide the necessary
broadband capability.
Doppler flow imaging (see section 4.7.4) came into widespread clinical use well before the
feasibility of time-domain processing became apparent. Time-domain processing is somewhat
faster, however, it has better spatial resolution and it is not subject to an artifact corresponding
to the Nyquist limit (see section 4.7.3). Time-domain processing is usually less sensitive
than the Doppler method and the relative risks of the theoretical biological hazards of highintensity short-pulse and low-intensity delivery of similar quantities of energy require further
investigation (Wells 1995; and see section 5).
It is the time-domain processing of the video signals that is perhaps the most obvious
method of determining the velocity of blood flow or tissue motion. If a target can be visualized
by real-time two-dimensional imaging (in which the image is actually formed by the video
signals), its velocity can be estimated directly from the measurement of the distance that it
travels over a known interval of time. For example, Trahey et al (1987) tracked blood flow in
sequential two-dimensional image frames by means of a localized two-dimensional speckle
correlation search.
Another approach, applicable to blood flow studies but not to solid tissue motion, makes
use of the decorrelation of the speckle pattern, which occurs at a rate that depends on the blood
flow velocity (Gardiner and Fox 1989). The potential of this method remains to be explored.
4.7.6. Colour coding schemes. Colour is characterized by luminosity (brightness or shade),
hue (which is determined by the wavelength of the light, but which is, of course, a purely
subjective phenomenon for the observer), and saturation (which is the degree to which a colour
departs from white and approaches a pure spectral colour). The wavelengths of the visible
spectrum extend from about 390 nm (violet) to 740 nm (red). Under favourable conditions,
an observer with normal colour vision can discriminate between hues in the middle part of the
spectrum with wavelength differences of around 1.2 nm; about 130 steps of hue difference can
be perceived across the entire visible spectrum. As the saturation of a hue is decreased to form
a tint, about 20 different levels can be identified.
Two colour coding schemes are in common clinical use for blood flow studies. In ‘colour
velocity imaging’, the image is colour-coded according to the velocity of blood flow. Increasing
flow velocity towards the probe is coded red-orange-yellow-white; increasing flow velocity
in the opposite direction is displayed in dark blue-blue-light blue-white. The colour scales
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Figure 15. Heart scan made with the probe positioned on the anterior thorax at the lower edge of
the rib-cage. The left ventricle is the large dark area near the apex of the scan. During diastole,
blood from the left atrium, which is the chamber below the left ventricle in this scan, enters the left
ventricle through the mitral valve. In the normal, the mitral valve closes in systole and blood is
pumped through the aortic valve into the aorta. This scan was made during systole, and it can be
seen that the mitral valve is not fully closed: this allows blood to flow back from the left ventricle
into the left atrium, as is evident from the green colour which represents the Doppler-shifted signal
from the reverse blood flow (see colour bar on the right). The severity of the condition, which
is known as ‘mitral regurgitation’, can be judged from the area of the colour-coded region. The
distance markers are separated by intervals of 10 mm.
can be interpreted without the need to refer to a coding key. Perversely, however, the colour
directions are the opposite to those that Doppler (1843) hypothesized for the light from stars,
and adopted by Fish (1981) in his colour flow scanner (see section 4.6.3). In addition to
colour coding according to velocity, colour can be used to indicate the variance of the velocity
estimate, which is related to the degree of flow disturbance. This can be done, for example,
by restricting the limits of the velocity image to the reds and blues and injecting an increasing
amount of yellow as the flow disturbance increases. The yellow mixes with red to form orange
and with blue to form green.
Examples of the clinical applications of colour velocity imaging are shown in
figures 15–17.
In ‘colour power imaging’, which is the other principal method of colour flow image
display, red or blue luminosity (depending on the preference of the observer) is used to
indicate the power, or amplitude, of the blood flow signal. By appropriately selecting the
gain of the system, stationary echoes are displayed with low luminosity, whilst blood flow
appears proportionately brighter as the signal power increases (Rubin et al 1994). The method
does not distinguish between forward and reverse flow. Its principal advantage is that it is
more sensitive than colour velocity imaging. This is because, as the system gain is increased
to increase the sensitivity in colour velocity imaging, the entire spectrum of colour begins to
obliterate the display, as the broadband noise level is approached. In colour power imaging,
noise merely increases the background image luminosity; blood flow signals, even though
only weak, cause a proportionate further increase in the localized luminosity. An example of
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703
Figure 16. Colour-velocity image of the umbilical cord of a fetus, floating in the amniotic fluid
within the uterine cavity. The cord contains two arteries and one vein, which can be distinguished
in the scan by the directional colour-coding (indicated in the colour bar on the right). The distance
markers are separated by intervals of 5 mm.
Figure 17. The brain is supplied with blood by the internal branches of the two carotid arteries,
one on each side of the neck. In this colour velocity scan, the patient’s head is on the left. The
carotid artery is partially obstructed by atheromatous plaque, which results in flow disturbance:
this is evident from the mosaic of colours, in contrast with the uniform red in the undiseased part of
the artery. The blue colour represents blood flowing from the head, in the jugular vein. The colour
bar shows the coding scheme and the distance markers are separated by intervals of 5 mm.
a colour power image is shown in figure 18.
Scanners used for colour-coded imaging of blood flow have a control (typically labelled
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Figure 18. Transverse scan of kidney, made with the probe on the patient’s back, a few centimetres
to the left of the spine. The blood flow is colour-coded according to the power of the Doppler-shifted
signal. Although the coding scheme is not affected by the direction of the blood flow (the blue in
the colour bar is not assigned), the renal artery and renal vein can be seen lying side-by-side in the
lower right corner of the colour box. The small vessels in the outer cortex of the kidney can also
be seen, which would probably not be possible if the colour-coding was according to the velocity
of the blood flow. The distance scale is in centimetres.
‘colour write priority’) that allows the operator to inhibit the display of colour over those areas
of the image that are characterized by the presence of higher amplitude echoes than those that
are obtained from blood. This prevents the appearance of colour due to the motion of solid
tissues. Under some circumstances, however, information about the motion of solid tissues
can be of clinical value. To provide a colour-coded image of solid tissue motion, the colour
write priority is adjusted to inhibit the display of colour from image areas with low amplitude
echoes: only high amplitude echoes are displayed, colour-coded according to their velocity or
power. The process is sometimes called ‘tissue Doppler imaging’ (McDicken et al 1992). As
with colour flow imaging, colour tissue motion imaging may be coded according to velocity
or power (Hoskins and McDicken 1997).
Colour Doppler imaging, whether of flow or tissue motion, is sensitive to the angle between
the direction of the ultrasonic beam and the target movement, as indicated in equation (11). It
is possible, however, to determine the direction of target movement by measuring its velocity
from several different positions on the scan plane. Thus, it is possible to produce a twodimensional image in which the colour represents the speed of target movement (Hoskins
1997).
The distinction between blood flowing in a large vessel and in solid tissue with the
minimum of capillary blood flow to maintain viability ignores the intermediate situation in
which there is significant perfusion and a multiplicity of small vessels within solid tissue.
For example, malignant tissue is characterized by numerous small vessels with direct shunts
between arteries and veins: the associated ultrasonic Doppler blood flow signals are well
Ultrasonic imaging of the human body
705
correlated with the presence of cancer (Wells et al 1977). By carefully adjusting the upper and
lower thresholds of Doppler signal amplitude between which colour is displayed in the image,
the extent of tissue vascularity can be estimated by counting the proportion of coloured pixels
(Cosgrove et al 1993).
4.8. Three-dimensional image acquisition and display
Baum and Greenwood (1961) prepared a series of photographically-reversed (i.e., the higher
the echo amplitude, the darker the registration) contiguous two-dimensional ultrasonic scans
of the eye, on transparent plates. By stacking these plates with appropriate spacing, they
constructed a three-dimensional image of the eye. Such an image enables even a relatively
unskilled observer to appreciate ultrasonic information in a three-dimensional form. With
this method, however, the deeper echoes are visible only poorly because they are obscured by
superficial echoes.
Although the pioneering work was followed by other attempts at ultrasonic threedimensional image acquisition and display, with a number of small improvements such as
holography (Redman et al 1969) and stereoscopy (McDicken et al 1972), real progress was
not possible until digital image storage and manipulation became practicable (Halliwell et al
1989).
Imaging involves the three stages of acquisition, processing and display. For a typical
penetration of 150 mm, the pulse repetition rate might be 5000 s−1 . As explained in
section 4.1.3, this corresponds to an image frame rate of 20 s−1 with 250 lines per frame.
The important point is that two-dimensional ultrasonic scanning can be in real time, with the
depths of penetration and image line densities that are needed to be clinically useful. It is true
that the image frame rate has to be reduced in duplex and, to a greater extent, in colour flow
scanning (see section 4.7) but, even so, one of the important advantages of two-dimensional
ultrasonic imaging is that it is essentially a real-time process.
Three-dimensional ultrasonic imaging requires the acquisition of a set of lines of image
information (echo wavetrains) not just from a two-dimensional tissue plane, but from a threedimensional tissue volume. Imagine that a volume of tissue 100 mm2 × 150 mm deep is
to be scanned in three dimensions, by acquiring a set of contiguous two-dimensional scans.
Allowing for the fact that the spatial resolution in elevation is likely to be around three times
worse than that in azimuth (see section 4.4), 33 contiguous scan planes each of 100 lines
would be appropriate. This means that 3300 lines would need to be acquired and, at a pulse
repetition rate of 5000 s−1 , 1.5 image data sets could be acquired per second. Although this
is not fast enough to qualify as real time, it certainly is adequate to be clinically useful for
the investigation of relatively static anatomical structures. Also, by synchronizing the image
acquisition with the electrocardiogram, satisfactory three-dimensional scans of the heart in the
various phases of the cardiac cycle can be acquired in a few seconds. Contemporary techniques
in x-ray computed tomography, magnetic resonance imaging and nuclear medicine cannot even
remotely approach the speed of three-dimensional ultrasonic imaging.
Figure 19 shows three ways in which three-dimensional image data can be acquired with a
phased array sector scanner. All these methods can be implemented by a motorized movement
of the transducer within a hand-held device. Another approach is by free-hand scanning with
a transducer designed for two-dimensional imaging, with the position and orientation of the
probe being measured in three-dimensional space. Although the operator needs to be skilled
in order to acquire a regularly-spaced image data set, the probe has hardly to be any more
bulky than for traditional scanning. It is prudent to ensure that the spatial resolution obtainable
in two-dimensional scanning is maintained in the three-dimensional image and, in abdominal
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Figure 19. Three-dimensional image data acquisition with a phased array scanner. (a) Rotation of
the transducer about an axis normal to the centre of the array scans a conical volume of tissue. (b)
Rotation of the transducer about the longitudinal axis of the array scans a pyramid-shaped volume
of tissue. (c) Translation of the array along an axis normal to the longitudinal axis of the array
scans a wedge-shaped volume of tissue.
imaging, this means that the spatial measuring system typically needs to have positional and
angular accuracies of better than around 0.5 mm and 0.5◦ , respectively. Articulated arms with
angle encoders, acoustic methods using spark gaps and microphones, and light-emitting diodes
with stereoscopic TV monitoring, have all been tried. Currently, however, the most popular
method seems to be that based on electromagnetic field transmitters and receivers (Detmer
et al 1994, Barry et al 1997). In this method, the receiving sensor is typically a 10 mm cube,
containing three orthogonal coils. It is attached to the ultrasonic probe. The transmitting coils
are mounted in a fixed spatial relationship with the couch on which the patient lies. These
coils are pulsed in sequence to emit three orthogonal magnetic fields. The process is repeated
fast enough to track even rapid changes in the position and orientation of the ultrasonic probe.
Typically, a skilled operator can acquire a good grey-scale image data set in around 5 s, or a
colour flow image data set in around 20 s.
In principle, it is possible to acquire a three-dimensional image data set without any
mechanical scanning, by the use of a two-dimensional transducer array. This is because such
an array can steer the ultrasonic beam both in azimuth and in elevation. Progress has been
made in the construction and operation of such arrays (Smith et al 1991, von Ramm et al 1991)
but formidable problems remain, mainly due to the large number of small transducer elements
that are required.
An image data set, whether acquired by automatic or by manual scanning, can be
considered to be a three-dimensional block of small volume elements (called ‘voxels’), the
brightness or colour of which carry the same information as the picture elements (‘pixels’) in
an ordinary two-dimensional scan. An important difference from two-dimensional imaging,
however, is that the complete three-dimensional data set cannot be collected in true real time
with contemporary technology.
A three-dimensional ultrasonic image data set can be processed and displayed on a
computerized workstation of the kind currently used in radiology for three-dimensional xray computed tomography and magnetic resonance imaging (Fishman et al 1991). An
example of such a display is shown in figure 20. Unlike three-dimensional CT and MRI
(in which, for example, bone can easily be distinguished from soft tissue), however, ultrasonic
Ultrasonic imaging of the human body
707
Figure 20. Display of three-dimensional scan of ovary obtained with an endovaginal probe
incorporating a curvilinear transducer array, mechanically rotated through a sector in the plane
orthogonal to that of the array scan. The three-dimensional image can be displayed in sections with
positions and orientations selected by the operator, either in real time or from the stored data.
images generally cannot satisfactorily be automatically ‘segmented’ into separate anatomical or
structural elements. (An exception is the lumens of blood vessels, where these are characterized
by the presence of colour-coded flowing blood.) This is because ultrasonic images are
essentially ‘noisy’, consisting largely of speckle due to the coherent nature of ultrasound (see
section 4.5). Consequently, surface fitting does not work well with grey-scale images (although
progress is being made in developing active contour models (Chalana et al 1996), integrated
edge maps (Aarnik et al 1998) and minimum cross-entropy thresholding (Zimmer et al 1996))
and segmentation methods are the subject of much contemporary research. Three-dimensional
ultrasonic angiograms, with power Doppler as the segmentation tool, are already being used
clinically (Richie et al 1996); they indicate what may become possible with grey-scale images
in the future.
Following successful segmentation, the ‘surface rendering’ display method can be used.
Lines-of-sight, or rays, are constructed from any chosen observation point and a perspective
view of the intersections of the segmented surface by these rays is displayed (Carson et al
1992). Stereoscopic viewing or rotation of the image gives the illusion of three dimensions.
Even when satisfactory segmentation is not possible, ‘volume rendering’ may be used to
display three-dimensional ultrasonic images (Nelson and Pretorius 1998). As with surface
rendering display, rays are cast from a point. The brightness (or colour) of each ray, however,
is determined by the local image brightness (or colour) as the ray passes through the image
volume. The image volume can then be viewed from any chosen direction and from any
chosen distance. In this way, stereoscopic pairs of images can be created, or the viewpoint
of the two-dimensional display of the observed image volume can be rotated to create depth
cues. Unless deep structures are entirely surrounded by a brighter shell, it is likely that they
all will become visible at least during part of the cycle of rotation.
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P N T Wells
Two-dimensional imaging of the velocity of tissue motion or blood flow using time-domain
processing is discussed in section 4.7.5. The time-domain processing method based on echo
tracking has been extended to three dimensions by Bohs et al (1995), who have presented
images of blood flow characteristics in laminar conditions and in jets.
From three-dimensional imaging, it is a simple step, in principle, to four-dimensional
(space and time) imaging. Preliminary studies of four-dimensional imaging of the fetal heart
have been carried out (Deng et al 1996). The data were acquired over about 20 s, using a gating
signal derived from a fixed-position pulse–echo ultrasonic beam for stroboscopic visualization.
4.9. Specialized imaging methods
4.9.1. Endoluminal scanning. ‘Endoluminal scanning’ is the term used to describe any
method of imaging in which scanning is accomplished by means of a probe positioned within
an anatomical cavity. Such a cavity may have a natural opening (e.g., the oesophagus, the
rectum and the vagina) or access may require tissue puncture (e.g., the arteries, the veins, the
abdominal cavity and the chambers of the heart).
The principal advantage of endoluminal scanning is that it brings the transducer into closer
proximity to the structures to be scanned than would otherwise be possible. This means that
higher ultrasonic frequencies can be used and this, together with the smaller thickness of
intervening tissue to distort the ultrasonic beam, results in images with better resolution and
clarity. The main disadvantage is that endoluminal scanning is a more invasive procedure.
Probes for endorectal and endovaginal scanning are usually based on small (10–
20 mm×5 mm) high frequency (5–10 MHz) linear or curved linear transducer arrays (see
section 4.2) mounted at the end of a suitably-shaped rod, the other end of which is fitted with a
handle for the operator. Because these probes are designed to scan within cavities with natural
openings, they do not need to be sterilized before use: it is sufficient for them to be cleaned
with antiseptic solution. An example of an endovaginal scan of an early pregnancy is shown
in figure 21.
For endo-oesophageal scanning, a small transducer is mounted at the tip of a long (up
to 1 m) flexible tube, with a maximum diameter of about 10 mm. The endoscope may
carry an optical imaging channel and there may also be provision for surgical instruments
to perform interventions under direct vision. Radial scanning around the axis of the endoscope
is accomplished by mechanical rotation of the transducer, which typically has a diameter of
3–5 mm, has lens focusing of the ultrasonic beam and operates at 5–10 MHz. The patient,
who is usually sedated, has to swallow the tip of the endoscope; a little local anaesthetic at the
back of the throat minimizes the discomfort of the procedure. Since the endoscope does not
penetrate the skin, it needs only to be cleaned with antiseptic solution.
For intra-abdominal scanning, a procedure which is becoming increasingly common as an
adjunct to minimally invasive laparoscopic surgery, rigid or flexible probes can be used. For
example, a typical ultrasonic laparoscope has a 5–10 MHz linear transducer array, 30 mm by
3 mm, mounted at the tip of a 50 cm rigid rod. The angle of the tip can be controlled in two
planes by the operator, using knobs on the handle of the laparoscope, to select the scan plane.
Since the laparoscope is introduced through a port through the abdominal wall, it has to be
sterilized before use. This is conveniently done by exposing it to an atmosphere of ethylene
oxide gas for 3–7 h, followed by 2 h aeration for detoxification.
For intravascular scanning, there are two distinct approaches. Both employ a flexible
catheter about 1.2 m long and with a diameter of about 3 mm, with a 20–30 MHz transducer
mounted at the tip to produce a radial scan: in one approach, the beam of a single-element
transducer is mechanically rotated (Wells 1966) and, in the other, the beam of a cylindrical
Ultrasonic imaging of the human body
709
Figure 21. Scan of early pregnancy made using an endovaginal probe such as that shown in figure 7.
The uterine cavity is the dark area in the middle of the picture. The fetus lies close to the probe.
The length of the fetus (crown-to-rump) extends between the two markers, which were positioned
by the operator. The scanner incorporates a look-up table that displays the corresponding time
since conception (in this case, 6.5 weeks). The distance scale is in centimetres.
array of (typically 64) transducer elements is electronically rotated (Bom et al 1972, O’Donnell
et al 1997). Mechanical scanning is usually accomplished by rotating the transducer through
a flexible drive-shaft by means of a motor mounted at the other end of the catheter, as shown
in figure 22 (Wells 1966, ten Hoff et al 1995), although micromotors small enough to be
incorporated into the catheter have recently been described (Erbel et al 1997). The catheter
is introduced into an appropriate blood vessel through a skin puncture and advanced until the
tip is positioned to scan the region of interest. For example, a femoral artery puncture in the
leg affords access to the coronary arteries in the heart. Examples of such scans are shown in
figure 23.
Endoluminal scanning is usually used for two-dimensional imaging. The operator can
obtain a three-dimensional impression of the scanned anatomy by changing the scan plane.
The scan plane can also be adjusted automatically in known increments, e.g., by pulling-back
an intravascular catheter, to acquire an image data set for three-dimensional display (Li et al
1995).
4.9.2. Synthetic aperture imaging. Synthetic aperture imaging depends on the synthesis of
the equivalent of a large aperture transducer by means of a small transducer which is scanned
to fill the aperture whilst acquiring pulse–echo information from the tissues to be imaged. It is
a two-stage process. In the first stage, the tissues are scanned from a line or an area, by a line
(fan beam) or point (conical beam) transmitter, and the scattered ultrasound is itself scanned
by a line or a point receiver. In a monostatic synthetic aperture, the same transducer is used
as the transmitter and the receiver. The technique has been tried for ultrasonic body scanning
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P N T Wells
Figure 22. Intravascular catheter probe for scanning blood vessels from within their lumens. Left
panel: detail of the scanning tip, showing the transducer, in side view, set a slight angle to the
mechanical drive shaft axis, to reduce the amplitude of the echo from the walls of the catheter;
the catheter is filled with liquid to provide ultrasonic coupling. Right panel: the complete device
showing the length of the catheter, the scanning tip and the connector (with slip rings, for electrical
connexion) for the motor drive to rotate the transducer.
(Burckhardt et al 1974, Ylitalo and Ermert 1994). Because a diverging beam has to be used,
the sensitivity decreases rapidly with depth. The sensitivity increases as the divergence is
reduced, but this also results in reduced spatial resolution (Ylitalo 1996). Nevertheless, the
imaging performance is theoretically better than that of a traditional B-scanner, although the
time required to synthesize the aperture may mean that physiological motion degrades the
image. Moreover, tissue inhomogeneity reduces the coherence of the ultrasound, on which the
method depends.
Synthetic aperture processing does have a useful role in intravascular ultrasonic scanning
with a small cylindrical transducer array (see section 4.9.1) and it may also reduce the image
acquisition time in three-dimensional ultrasonic microscanning (see section 4.9.5).
4.9.3. Computed tomography. The invention of x-ray computed tomography (Hounsfield
1973) had a profound effect on the practice of clinical radiology. The method depends on
the acquisition of an angular set of x-ray attenuation profiles from a two-dimensional tissue
plane, from which the image is produced by the process of back-projection reconstruction
(Ramachandran and Lakshminarayanan 1971). It works well partly because x-rays travel in
straight lines. In principle, ultrasonic computed tomography depends on the acquisition of a
complete set of projections of the characteristic to be imaged (i.e., velocity or attenuation).
Figure 24 shows how ultrasonic computed tomography can be used for imaging the
breast. Ultrasonic computed tomography depends on the validity of the assumption that
line-of-sight propagation is maintained across the tissues being scanned. Unfortunately,
however, an ultrasonic beam is deviated by refraction and distorted by inhomogeneities in
tissue. Consequently, and beginning with the work of Greenleaf et al (1974), who scanned an
excised dog heart at 5 MHz, the results have until recently been consistently disappointing. CT-
Ultrasonic imaging of the human body
711
Figure 23. Intravascular ultrasonic scans of coronary arteries. The scans were made with a highfrequency cylindrical array transducer with a diameter of about 3 mm, introduced via a puncture
into the femoral artery in the leg. The circle in the blood vessel lumen represents the transducer
array. Top panel: normal artery. Bottom panel: artery affected by eccentric atheromatous plaque
(indicated by arrows).
derived data on speed-of-ultrasound, however, has been used by Jago and Whittingham (1991)
to correct misregistration in superimposed B-scans obtained from multiple angles around an
excised sheep kidney, with encouraging results.
4.9.4. Elasticity imaging. Doctors all learn the art of paplation as an essential element of
clinical examination. Palpation depends on the differences in the hardness of different tissues
and thus can reveal the presence of abnormalities if they are close to the surface. An obvious
example is the discovery of a cancer of the breast, detectable because malignant tumour tissue
is harder than normal glandular and fatty tissues.
The goal of ultrasonic elasticity imaging is to map tissue properties such as Young’s
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P N T Wells
Figure 24. Ultrasonic computed tomography. In this example, time-of-flight tomography of
a breast immersed in water is being performed by translate-rotate scanning. (a) Mechanical
scanning system. (b) Profiles of time-of-flight measurements obtained by translation at three
angular positions: in a real system, such profiles are acquired at, typically, 128 angular positions
around the entire tomographic plane.
modulus (or stiffness), Poisson’s ratio and viscosity in an anatomically meaningful presentation
to provide useful clinical information (Gao et al 1996). The stress–strain relationship for most
tissues is nonlinear and stress tends to relax over time under constant strain (i.e., cyclic loading
and unloading is characterized by hysteresis).
Several methods of ultrasonic elasticity imaging have been demonstrated. In vibration
amplitude sonoelastography, a low-frequency (20–1000 Hz) vibration is externally applied to
excite internal tissue motion of which an image is produced by Doppler detection (Lerner
et al 1988). In vibration phase gradient sonoelastography, both the amplitude and the phase
of externally-excited low-frequency internal tissue motion are measured. By assuming that
viscosity at low frequencies is negligible and that shear waves predominate, Levinson et al
(1995) obtained phase gradient images of thigh muscle under various conditions of active
muscle contraction: both the speed of vibration propagation and the value of Young’s modulus
increase with increasing contraction. The method of compression strain elastography, in
which the tissue is externally compressed and pre- and post-compression ultrasonic A-scan
line pairs are crosscorrelated to produce a set of strain profiles, has been demonstrated by
Ophir et al (1991) and Kallel et al (1998). In an ingenious variant of the method, de Korte
et al (1998) used an intravascular imaging system to obtain strain elastograms of diseased
arteries in vitro, with the change in strain being produced by change in pressure. With further
refinement, this method may be practicable in vivo, by making use of the arterial pressure
pulse. O’Donnell et al (1994) have devised a method of compression strain elastography with
relatively large displacements (around ten wavelengths), measuring the overall displacement
by summing the small displacements resulting from incremental step loading from the Ascan crosscorrelations. Finally, Bohs and Trahey (1991) have used a two-dimensional speckle
tracking method employing a sum-of-absolute-difference criterion with a search kernel to
measure flow and tissue motion and this has been adapted by Walker et al (1993) for vibration
amplitude sonoelastography.
Ultrasonic elasticity imaging has not yet been used, except rather crudely, in routine clinical
practice. It is a very promising method, however, because it should have spatial resolution at
Ultrasonic imaging of the human body
713
Figure 25. Geometry of the ultrasonic beam produced by a
spherical bowl transducer.
least comparable with that of real-time grey-scale imaging combined with potentially better
tissue discrimination.
4.9.5. Microscanning. Microscanning is defined as the two- and three-dimensional display
of biomedical soft tissue structures with spatial resolution in the range 10–100 µm. This
resolution range lies between the smallest structures that can be seen under direct vision
and the largest structures that are traditionally of interest to histopathologists using the light
microscope. Although resolutions exceeding that of the light microscope can be obtained with
the scanning acoustic microscope (Lemons and Quate 1974) operating at frequencies above
3000 MHz (i.e., with wavelength of less than 500 nm), the need to section the tissue, the long
scanning times and the relatively high cost may be the reasons why this has not been introduced
into clinical practice. The scanning laser acoustic microscope (Kessler 1974), which is a realtime instrument in which the image is acquired by laser-scanning a liquid surface levitated by
ultrasound transmitted through the thin specimen, operates at frequencies of around 100 MHz
(i.e., with wavelengths of around 15 µm). The need for sectioning and the high cost have
prevented the method from becoming popular. Microscanning, with its ability to explore
tissues in three dimensions without the need for serial sectioning, has previously been largely
neglected but has the potential to advance biomedical science and clinical practice.
The geometry of a focused ultrasonic beam produced by a spherical bowl transducer is
shown in figure 25. The dimensions of the −3 dB profile of the ellipsoidal focal volume can
be calculated from the following equations (Wells 1977):
df ≈ λ(F /2a)
lf ≈ 15(1 − 0.01θ)df ,
(20)
(21)
where df is the cross-sectional diameter and lf is the axial length of the focal volume, a is the
radius of the spherical bowl and F is the focal length of the transducer, and θ is the half-angle
of convergence of the beam, provided that θ < 50◦ . The focal beam diameter can be used as
the starting-point for the design of an optimized microscanner. This is because it determines
the resolutions in elevation and in azimuth; the range resolution is not determined by the length
of the focal volume, but by the length of the ultrasonic pulse and, generally, this can be made
less than the beam diameter.
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P N T Wells
A typical scanning acoustic microscope operating at a frequency of 100 MHz, using a
transducer with diameter of 3.2 mm and a focal length of 6.4 mm, thus has df ≈ 30 µm and
lf ≈ 380 µm. This means that its optimal resolution cell is around 30 µm in diameter and, by
gating the echo wavetrain into separate receiver channels, 13 appropriately-spaced resolution
cells can be accommodated axially within the focal volume. For imaging a 5 mm tissue
cube from opposite sides in two 2.5 mm-thick slabs, sampling theory requires the scanning
beam spacing in the XY plane is equal to 15 µm; in the Z direction, the process has to be
repeated over about 13 sets of 13 planes, from each side. In order fully to sample the image
information from the 5 mm tissue cube, this means that the beam has to dwell to acquire echoes
on about 2.9 million occasions. Assuming a dwell time of 1 m s−1 (chosen to allow achievable
spatial scanning velocity and the possibility of signal averaging for noise reduction), the total
scanning time is about 48 min. At 100 MHz, the attenuation of ultrasound is about 20 dB cm−1
in water and about 100 dB cm−1 in soft tissue. Therefore, the overall dynamic range of gain
compensation needs to be about 40 dB for a 2.5 mm-thick slab of tissue.
It will be interesting to see whether three-dimensional miroscanning with these constraints
turns out to be clinically useful. If it does, it should be possible substantially to reduce the
image acquisition time by employing synthetic aperture scanning (see section 4.9.2).
4.9.6. Contrast agents. It was Gramiak et al (1969) who first reported that the ultrasonic
echogenicity of blood could be artificially enhanced. They observed that the injection of
indocyanine green dye into the chambers of the heart by means of a catheter (a procedure used to
opacify the chambers and vessels in x-ray imaging) coincidentally caused the blood transiently
to give rise to strong ultrasonic echoes. Subsequently, Kremkau et al (1970) showed that the
effect was not primarily due to the echogenicity of the indocyanine green, but to decompression
cavitation at the tip of the catheter. Two years later, Ziskin et al (1972) demonstrated that ether
was the most effective substance that they tested to enhance echogenicity, presumably because
it boils vigorously at body temperature. In any but tiny quantities, however, ether is toxic or
even lethal. Water saturated with carbon dioxide was found to be satisfactory, since it contains
numerous bubbles but gas emboli are not a complication as the bubbles rapidly dissolve in
blood.
Because gas has a very low characteristic impedance in comparison with that of blood
(see table 1), suspended gas bubbles greatly increase the echogenicity of blood. The scattering
is amplified if the size of the bubble is such that it resonates at the frequency of the incident
ultrasound. The resonant frequency f0 of a gas bubble immersed in a liquid of density ρ and
at pressure P is given by (Minnaert 1933)
f0 ≈ 21 πr(3γ P /ρ)1/2 ,
(22)
where r is the radius of the bubble and γ is the adiabatic ideal gas constant. This means that
a bubble with a diameter of 5 µm (comparable with that of a red blood cell) resonates at a
frequency between 1 and 10 MHz. The actual frequency depends markedly on the nature of
the shell that encapsulates the bubble. Commercially-available microbubble contrast agents
are encapsulated in a variety of substances, such as albumin, lipid and palmitic acid. The shells
stabilize the bubbles and extend their life after injection into blood.
It is a fortunate coincidence that the bubbles that resonate at the ultrasonic frequencies
commonly used for imaging have diameters of a few micrometres. From the point of view of
clinical convenience and the safety of the patient, an injection into an artery is considered to be
very much less desirable than an injection into a vein. Blood enters the right side of the heart
from the veins and then passes through the lungs before returning to the left side of the heart,
which pumps it through the systemic circulation. Consequently, an intravenous injection of
Ultrasonic imaging of the human body
715
ultrasonic contrast agent soon enters the systemic arterial system, provided that the individual
microbubbles are small enough to cross over the capillaries in the lungs. Obviously, red blood
cells do this; provided that the microbubbles are no bigger than red blood cells, they also can
do so. The only condition is that they should survive for long enough to produce the desired
image contrast.
Encapsulation of a microbubble results in an increase in its resonant frequency and a
decrease in the amplitude of scattered ultrasound (de Jong and Hoft 1993). The lifetime of
the microbubble can be extended either by increasing the stiffness of its shell or by using
a gas that dissolves poorly in blood (Frinking and de Jong 1998). Free air bubbles with a
diameter of about 8 µm disappear within 1 s; those with a thin albumin shell persist for up
to 10 min; and those with stiff shells or relatively insoluble gases can last for an hour or
more.
The obvious utility of an ultrasonic contrast agent arises because it increases the
echogenicity of blood and thus, for example, increases the sensitivity of ultrasonic Doppler
blood flow detection. It also results in a significant reduction in image clutter. The
concentration of contrast agent needed to produce useful enhancement of blood echogenicity
is typically 50 mg of concentrated bubble suspension per kg body weight. Used in this
way, microbubbles actually decrease the contrast between blood and the soft tissues that
surround the blood vessel and so the term ‘contrast agent’ is something of a misnomer. It
is possible, however, to use microbubble contrast agents in a way that does increase the
echogenicity of blood beyond that of soft tissues. One way might be to inject more contrast
agent, but this would not be attractive because of the risk associated with introducing gas
into circulating blood, where it might block small vessels and, for example, cause a stroke.
Another problem would be the increase in the ultrasonic attenuation in the blood (Bouakaz
et al 1998). The method that does work, however, exploits the fact that microbubbles
suspended in blood backscatter ultrasound at harmonics of the incident frequency, when
the incident pressure is sufficiently high (Schrope et al 1992). Consequently, if the
receiver is tuned to receive signals at twice the frequency of the transmitted ultrasound,
the echoes backscattered by the contrast agent microbubbles in the blood can actually
have a higher amplitude than those backscattered by the surrounding soft tissues. This
improves the discrimination between blood and tissue echoes and is known as ‘harmonic
imaging’ (Forsberg et al 1996). It has been shown by Zheng and Newhouse (1998) that,
for a given incident ultrasonic frequency, the backscattered signal commences with the
same frequency, but this is followed after an ‘onset delay’ of typically 15 cycles of the
fundamental by the development of the second harmonic. This effect must limit the range
resolution obtainable with harmonic imaging, although this does not yet seem to have been
investigated.
At first sight, it is rather surprising that a stronger backscattered signal is detected at
the second harmonic of the bubble resonant frequency, than at the third harmonic. This
phenomenon has not yet been thoroughly investigated but a possible explanation is that
increasing attenuation in overlying tissues (which increases with the frequency) and the limited
frequency bandwidth of practical transducers are dominant. Another possibility is that the first
subharmonic frequency may be strongly backscattered and Shankar et al (1998) have shown
that it should be more easily detectable than the second harmonic. Improved detectability has
to be weighed against reduced resolution, however, and this is also a topic for research in the
future.
It has been reported (Porter et al 1997) that encapsulated microbubbles can be destroyed
by the ultrasonic exposure conditions used in some diagnostic procedures. The phenomenon
is accompanied by the emission of strong ultrasonic transients that can result in the appearance
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P N T Wells
of high brightness pseudoechoes in pulse–echo grey-scale images and multicoloured transients
in flow and motion images. Thus, it provides a technique further to increase the sensitivity
of ultrasonic imaging systems to the presence of blood containing ultrasonic contrast agent.
Furthermore, the annihilation of contrast agent by exposure to a brief pulse of ultrasound can
provide a relatively precise timing marker for the measurement of blood flow and perfusion
dynamics.
In addition to the use of contrast agents for improved imaging by echo enhancement,
contrast agents can also be used for dynamic studies. Thus, a rapid increase in image brightness
may occur after injection of a bolus of contrast agent, followed by a slower decline (Sehgal
and Arger 1997). Analysis of the brightness-time curve can provide data on blood flow rate
and blood perfusion.
According to equation (22), the resonant frequency of a bubble is proportional
to the square root of its ambient pressure.
Fairbank and Scully (1977) proposed
that a noninvasive method for intracardiac blood pressure measurement could be based
on this relationship. Such a method would be of very great clinical value, since
blood pressure can otherwise be measured only indirectly or invasively. Unfortunately,
however, there are several reasons that make the resonant bubble approach impracticable.
Firstly, bubbles cannot be manufactured with sufficient accuracy to ensure that they
are all the same. Secondly, and of fundamental importance, the bubble resonance
is not sharp enough to provide the pressure discrimination necessary to be clinically
useful.
4.9.7. Tissue harmonic imaging. The reduction in image clutter that can be obtained by
restricting signal detection to the second harmonic frequency backscattered by ultrasonic
microbubble contrast agents is discussed in section 4.9.6. The method is possible because
bubbles behave in a nonlinear fashion when the ultrasonic pressure amplitude is sufficiently
high. It is not only bubbles that demonstrate nonlinearity, however, and table 1 shows that
biological soft tissues also have this property. Consequently, ultrasound backscattered by
tissues has harmonic frequency components that can be detected if the incident ultrasonic
pressure is sufficiently high. As with microbubble contrast agents, it is the second
harmonic frequency that is most relevant. There the direct similarity ends. In tissue
harmonic imaging, the harmonics are not generated during the scattering process, but
during the passage of the ultrasonic pulse through the tissue towards the scatterer (and
not following scattering, because the ultrasonic pressure amplitude is then too small
for nonlinearity to have any material effect). Thus, the process is that described in
section 2.8.
By tuning the receiver to the second harmonic of the transmitted ultrasonic frequency,
two advantages can be gained. The first is that echoes detected from tissues close to the
transducer are relatively weak, because second harmonic generation builds up to a significant
level only as the result of adequate tissue penetration. The second is that contrast resolution
is improved (because the side lobes of the transmitting and receiving beams have different
angular structures) and the spatial resolution is improved (because the distortion of the
transmitted beam by the inhomogeneity of the superficial layers is relatively small and the
receiving beam can be made narrower than the transmitted beam). The practical realization
of tissue harmonic imaging is relatively new, but see, for example, Thomas and Rubin (1998)
for a discussion of the method in the context of echocardiography. Figure 26 illustrates
the improvement in image quality that can result when tissue harmonic imaging is used in
abdominal scanning.
Ultrasonic imaging of the human body
717
Figure 26. Longitudinal scans of the right kidney, which contains a fluid-filled mass, made with
the probe on the anterior abdominal wall immediately below the rib cage. The wedge-shaped
structure near the top of each image (i.e., close to the probe) is the liver. Right panel: traditional
signal processing, with the transmitter and the receiver operating at the same frequency. Left
panel: second harmonic imaging, with the receiving transducer operating at twice the transmitted
frequency. The distance markers are separated by intervals of 10 mm.
5. Safety considerations
Because bioeffects, some of which are harmful, may be caused by ultrasound under certain
exposure conditions, there is a hypothetical possibility that ultrasonic imaging may not be
completely safe (Wells 1986). Moreover, the ultrasonic exposure levels used by commerciallyavailable scanners have been steadily increasing, in order to obtain more information (Duck
and Martin 1991). Consequently, both regulatory authorities and prudent clinicians take the
subject seriously.
The World Federation for Ultrasound in Medicine and Biology (WFUMB 1992) has
published the following statements on thermal effects in clinical applications.
B-mode imaging: known diagnostic ultrasound equipment as used today for simple Bmode imaging operates at acoustic outputs that are not capable of producing harmful
temperature rises. Its use in medicine is therefore not contraindicated on thermal
grounds. This includes endoscopic, transvaginal and transcutaneous applications.
Doppler: it has been demonstrated in experiments with unperfused tissue that some
Doppler diagnostic equipment has the potential to produce biologically significant
temperature rises, specifically at bone-soft tissue interfaces. The effects of elevated
temperatures may be minimized by keeping the time for which the beam passes
through any one point in tissue as short as possible. Where output power can be
controlled, the lowest available power level consistent with obtaining the desired
diagnostic information should be used.
Although the data on humans are sparse, it is clear from animal studies that exposures
resulting in temperatures less than 38.5 ◦ C can be used without reservation on thermal
grounds. This includes obstetric applications.
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P N T Wells
Transducer heating: a substantial source of heating may be the transducer itself.
Tissue heating from this source is localized to the volume in contact with the
transducer.
The possibility that nonthermal effects of ultrasound may be hazardous in some situations
is more contentious (Barnett et al 1994). Cavitation, defined as the formation or activity
of gas- or vapour-filled cavities (bubbles) in a medium exposed to an ultrasonic field, is the
phenomenon of most concern. Other possible nonthermal mechanisms include radiation force,
acoustic torque and acoustic streaming.
Many questions relating to the safety of the ultrasonic exposures used for imaging remain
to be answered. For example, is there a linear relationship between the quantity of image
information and the ultrasonic energy needed to obtain it? The thermal effect of a given
ultrasonic power may be independent of the exposure duty cycle, but what are the nonthermal
effects under different regimes? Pragmatically, users of ultrasound for diagnosis should apply
the ALARA (‘as low as reasonably achievable’) principle to the exposures to which they subject
their patients. The exposure levels should be at the lowest intensities and for the briefest times
necessary to obtain diagnostically-adequate images. To assist in achieving this goal, a predictor
of cavitation known as the mechanical index (MI) has been developed (AIUM/NEMA 1992),
given by the expression
MI = Pr · 3(zsp )fc1/2 ,
(23)
−1
−1
where Pr · 3(zsp ) is the peak rarefactional pressure (MPa) derated by 0.3 dB cm MHz to
the point on the beam axis (zsp ) where the pulse intensity integral is maximum, and fc is the
centre frequency (MHz). Some modern scanners display the MI value on the screen, so the
operator can aim to minimize it.
Although it is right to be concerned about exposure conditions, it is the consequences
of misdiagnosis that are likely to be the greatest hazard of an ultrasonic investigation (Wells
1986). Other risks must not be ignored, but they must be viewed in the proper perspective.
6. Conclusions and future prospects
The physics of ultrasonic imaging is quite well understood. The performance of pulse–
echo grey-scale imaging systems is ultimately limited by the attenuation, nonlinearity and
inhomogeneity of tissues and by the need to minimize the possibility of hazard by minimizing
the ultrasonic exposure. Despite these limitations, however, there is scope for improvement.
The rate of improvement in the capabilities of digital electronic circuits shows no sign of
diminishing and, as costs continue to fall, so ultrasonic signal digitization is moving closer to the
individual transducer element. The potential benefits of contrast agents are only just beginning
to be explored and the reduction in clutter that can be obtained by second harmonic imaging,
both without and with contrast agents, is an excellent example of very recent progress. New
transducer materials will probably result in greater sensitivity and better noise performance
and may make it possible to reduce exposures. Some of these materials may also be used in
affordable two-dimensional transducer arrays for three-dimensional image acquisition.
Although there are many similarities between Doppler (i.e., frequency or phase) and timedomain processing for obtaining information about blood flow and tissue motion, the two
techniques are often considered to be different and sometimes, in competition. Time-domain
processing does have some advantages (e.g., there is no direct equivalent of ambiguity due
to the exceeding of the Nyquist limit), but it is computationally more demanding than the
Doppler technique. As computing becomes more accessible and less expensive, however, this
Ultrasonic imaging of the human body
719
is becoming less of a consideration. In situations in which the greatest sensitivity is required,
colour power imaging has an important low-noise advantage over colour velocity imaging,
whether performed by Doppler or time-domain processing.
Three-dimensional imaging improves the perception of anatomical relationships, whether
by specialists reviewing complicated situations or by untrained observers with routine cases.
There is scope for increasing the rate of three-dimensional image information acquisition by
the use of multiple-beam systems with parallel processing. Image segmentation remains a
problem, however, and this limits the opportunities for useful image display.
Amongst the specialized imaging methods, endoluminal techniques are already in routine
use. Synthetic aperture imaging may become useful, at least until two-dimensional arrays
have been further developed, and in microscanning applications. Computed tomography may
be used to provide information about tissue refraction and attenuation, for the improvement
of traditional pulse–echo images. Elasticity imaging is a very promising technique; it may be
developed into quantitative telepalpation.
Although the primary contemporary role of ultrasonic imaging is in diagnosis, the method
also has important applications in monitoring the progression and regression of disease,
in some areas of screening, and in interventional procedures, both for localization and for
guidance. Ultrasonic imaging is likely to become one of the preferred visualization techniques
in minimally invasive surgery, because of its high speed and ease of use.
The safety record of ultrasonic imaging is impeccable. There is no reason to suppose that
contemporary techniques employ levels of exposure that could cause biological damage, but
it is prudent to be cautious and further research is justified.
Acknowledgments
I am grateful to Dr H S Andrews for providing figures 10 and 21. The following figures are
from product literature: Advanced Technology Laboratories, figures 7, 9, 15–18; Medison,
figure 20; Endosonics, figure 23; and General Electric, figure 26.
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