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Ultrasonics in medicine and biology

This content has been downloaded from IOPscience. Please scroll down to see the full text. 1977 Phys. Med. Biol. 22 629 (http://iopscience.iop.org/0031-9155/22/4/001) View the table of contents for this issue, or go to the journal homepage for more

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PHYS. MED. BIOL.,

1977, VOL. 22,

NO.

0 1977

4, 629-669.

Review Article

Ultrasonics in Medicine and Biology P. N. T. WELLS,

PH.D.

Department of Medical Physics, Bristol General Hospital, Bristol BSI6 S Y , UK This review was completed in January 1977

Contents 1. Introduction physics 2. Fundamental

.

.

.

detection3.and Generation . 3.1. Magnetostrictive transducers transducers Piezoelectric 3.2. . 3.3. Measurement of ultrasonic power 4. Interactions with biological materials absorption 4.1, Attenuation and tissues 4.1.1. softAbsorption in . elocity 4.2. 4.3. Attenuation and velocity lung and in bone 4.4. Scattering . impedance Characteristic 4.5. . Pulse-echo 5. methods diagnostic . considerations 5.1. Physical . Pulse-echo 5.2. diagnostic instrumentation 6.2.1. Displays . scanners Real-time5.2.2. . Analysis 5.3. of pulse-echo data . Clinical 5.4. applications . . Angiology 5.4.1. Cardiology 5.4.2. . 5.4.3. Endocrinology . Gastroenterology 5.4.4. . 5.4.5. Neurology. gynaecology. 5.4.6. and Obstetrics Oncology 5.4.7. . 5.4.8. Ophthalmology . 5.4.9. Urology . methods diagnostic 6. Doppler . . 6.1. methods Transmission

.

~~

630 631 , 633 . 633 . 633 . 633 . 633 . 633 . 634 . 636 . 636 . 638 . 638 . 639 . 639 . 641 , 642 . 643 . 644 . 646 . 646 , 646 , 647 . 647 . 647 . 647 . 648 . 648 . 648 . 649 . 649 ,

. . . . .

.

~~

Copies of this Review are available from: Offprint Department, Institute of Physics, Techno House, Redcliffe Way,Bristol BSI 6NS, U.K. Price 21.00, including postage (surface mail). 21

y

630

P. N . T. Wells

6.2. Reflection methods . 6.2.1. Continuous wave systems . . 6.2.2. detectors Directional . 6.2.3. Range-measuring Doppler systems . 6.2.4. Two-dimensional Doppler imaging Clinical 6.3. applications . 6.3.1. Angiology . . 6.3.2. Cardiology 6.3.3. Obstetrics . methods diagnostic 7. Other . 7.1. Transmission methods . 7.2. Holography . 7.3. 8. Biological effects . 8.1. Thermal effects . 8.2. Non-thermal effects . 8.2.1. Cavitational effects . 8.3. Frequency dependence of biological effects , 8.4. Cumulative effects . . 8.5. Other biological effects 8.5.1. onEffects blood flow . , 8.5.2. Effects malignant tumours on . 8.5.3. Effect regeneration tissueon 8.5.4. Evidence of genetic effects . 8.6. Possibility of hazard in diagnostics . 9. Therapeutic and surgical methods . 9.1. Physiotherapy . 9.2. Neurosurgery . . surgery 9.3. Vestibular laryngeal papillomatosis 9.4. Treatment of juvenile 9.5. Surgical procedures using direct mechanical effects . 10. Conclusions

. .

. .

. .

. . . .

. .

.

.

.

. . . . . . . . . . . . . . . . . . .

649 649 650 650 651 652 652 653 654 654 654 655 656 657 657 658 658 659 659 659 659 659 660 660 660 660 660 661 663 663 663 665

1. Introduction

Almost fifteen years have elapsed since the subject of ultrasonics was last reviewed in Physics in Nedicine and Biology (Newel1 1963). During this time, there have been many scientific developments as a result of which ultrasonic techniques, and especially diagnostictechniques, now have an established place in biomedicine. The importance of ultrasonic diagnostics is due t o its fundamental differences from investigativemethodssuchasconventionalradiography,scintigraphy and X-ray and radioisotopecomputerizedtomography.Ultrasonicmethods can be used to visualize soft-tissue structures andt o study physiological movements, using exposure conditions which seem to be harmless. The symptoms

Ultrasonics in Medicine and Biology

631

of some diseases, and normal conditions such as pregnancy, are best investigated, a t leastinitially, by ultrasound. Many clinical specialists depend on ultrasonic diagnostic information in managing their patients, and the patients themselves are beginning to demand this kind of investigation. Progressin the scientific study of ultrasonicphysiotherapyhas been disappointingly slow. Much research hasbeen done, however, into the biological effects of ultrasound. The application of ultrasound in surgery has perhaps not met early expectations, but there are some procedures-notably in the treatment of vestibular disturbances-which are routinely used. Recently, research into the effects of tissues on ultrasound has increased substantially. This work.has merit as a purely academic pursuit, and moreover it may lead eventually t o the development of methods of identifying tissues from their ultrasonic characteristics. Newell's (1963) review contains a summary of the basic physics of ultrasound. In the present article, references are given to recent and more comprehensive papers and books ; brief discussions of the ultrasonic field and some methods of generation and detection are included. The larger part of the review is concerned with the medical and biological aspects of the subject. 2.

Fundamental physics The wave equation for longitudinal mechanical waves is a2u

1 a2u "-

"

a22

c2 a t 2

where U is the particledisplacementamplitude, z is the positioninspace along the direction of propagation, t is the time and c is the propagation velocity. c=WP)' (2) where K is the bulk modulus of the medium and p is its density. The solution of eqn (1)leads to therelationships summarizedby Newel1 (1963), and discussed in detail by Wells (1977a). Amongst these relationships, the more important for plane waves are

z = pc

( 3

where Z is the characteristic impedance of the medium;

I

= pcvo2/2

(4)

where I is the intensity of the wave and v. is the peak particle velocity. A t a plane boundary between two media with velocities c, and c2 respectively and

ei = er (sin O,/sin e,) = cl/c2

where e,, 0, and et are respectively the angles of incidence, reflection and refraction. At normal incidence,

4/4= r ( z 2 - z 1 ) / ( Z 2 + ~ 1 ) l 2

(7)

632

P.N . T.Wells

where Ii and I, are respectively the intensities of the incident and reflected waves, and 2, and 2, are the characteristic impedances of the two media. The apparent frequency of a constant frequency source is dependent on the motion of both the source and the receiver: this is the Doppler effect. Iff is the source frequency, and fr is the received frequency, the Doppler shift frequency fu = (f,-f ) is given by

c-v,

fD=

(=-l)f

where v, and vB are respectively the vector velocities of the receiver and the source. When an ultrasonic wave is absorbed (or reflected), a force is exerted in the vectordirection of the propagation of the wave.This effect, loosely called radiation pressure, is not fully understood (e.g. Rooney 1973). In the case of complete absorption.

F

=

W/C

(9)

where F is the radiation force and W is the ultrasonic power. One of the most commonly used ultrasonictransducersin biomedical applications is in the form of a disc of ferroelectric material (see section 3). With continuous wave excitation producing ultrasound of wavelength h,

&/Io= sin2 { ( ~ / h[(a2 ) + x2)t - x ] }

(10)

where Io is the intensity at the surface of the transducer, I, is the intensity a t a distance z from the transducer along the central axis and a is the radius of the transducer. In the far field, the directivity function is given by

D,= 2J,(kasin e) ka sin e

where k = 2rf, where the frequency f = c/h, and J, is the first order Bessel function; e is the angle relating D, to the central axis of the beam. If the transducer produces a transient ultrasonic disturbance, as distinct from a steady state continuous wave, the directivity of the ultrasonic beam is modified because, at any particular point in the field, the contributions from the different elementary parts of the surface of the transducer may be unequal. Therearetwoapproaches tothe solution of this problem. Firstly,the computation of the field maybe based ontime considerations andthe knowledge of the form of theexcitation of the transducer(Beaver 1974, Robinson, Lees and Bess 1974). Secondly, the analysis may take account of the time-averaged distribution of energy within the ultrasonic beam during the propagation of the pulse. This involves the concept of the frequency spectrum of the radiated energy, and the problem may be solved by weighted superimposition of the fields of single-frequency radiation at variousfrequencies within the band (Papadakis and Fowler 1971).

Ultrasonics in Medicine and Biology

633

Ultrasonic beams can be focused, for example, by means of lenses (Fry and Dunn 1962), or by curved transducers (O’Neil 1949), or by concave mirrors (Olofsson 1963). 3. Generationanddetection 3.1. Magnetostrictive transducers At very low ultrasonic frequencies, ferromagnetic metals such as nickel are used as magnetostrictive transducers. Even with laminated construction, they become unacceptably inefficient as the frequencyis increased aboveabout 30 kHz. Ferrite materials are satisfactory as magnetostrictive transducers a t frequencies of up to about 60 kHz (van der Burgt 1958). 3.2. Piezoelectric transducers Piezoelectric materialsare commonly used astransducersin biomedical applications. Ferroelectrics, such as lead zirconate titanate (Jaffe, Roth and Marzullo 1955), are very satisfactory a t frequencies of up to around 10 MHz : they can be manufactured in a wide variety of shapes and sizes and they have a lower electrical impedance than quartz. Short pulses of ultrasound can be generatedand received by probeswithappropriate backing and matching arrangements (Kossoff 1966, Kasai, Okuyama and Kikuchi 1973). 3.3. Measurement of ultrasonic power Calorimetry and radiation force measurement (e.g. Wells, Bullen, Follett, Freundlich and James 1963, Zieniuk and Chivers 1976) are both widely used for power measurement in the biomedical applications of ultrasound. Thermocouple probes (Fry and Fry 1954a, b) are used for intensity determinations. Optical techniques (e.g. Mezrich, Etzold and Vilkomerson 1975) are used to detect ultrasound, but generally not to measure power. 4. Interactions with biologicalmaterials 4.1. Attenuation and absorption

The amplitudeof a planewave propagating in the z-direction can be expressed in the form:

A, = A,exp (-pax) (12) where A, is the peak value at z = 0 of a wave variable such as particle velocity, and A , is the peakvalue a t z of the samevariable; pa is the amplitude attenuation coefficient, measured,forexample,inneperspercentimetre. Attenuation is more commonly expressed in terms of a coefficient a , measured in decibels per centimetre ; then CY

= 20 (log e) pa N 8 . 6 9 ~ ~ .

(13)

Attenuation refers to the total propagation loss; absorption refers only to that component of attenuation by which ultrasonic wave energy is converted intoheat.Thus,attenuation mechanisms include reflection, scatteringand absorption.

P.N . T.Wells

634

4.1,l . Absorption in soft tissues. Most of the published data for absorption in biological materials are given in fig. 1. Some of these data are actually values of attenuation, but often the literature is not clear enough for this distinction 50

2ol 10

0.1

05

1 f

5 10 IMsriZl

50

103

500 IOCC

Fig. 1. Absorption in normal biological materials, expressed in terms of data.Data sources: [l] Calderon,Vilkomerson,Mezrich, Etzold, Kingsley and Haskin 1976; [2] Carstensen and Schwan 1959; [3] Chivers and Hill 1975; [4] Colombati and Petralia 1950; [5] Dankwerts 1974; [B] Dunn 1962; [7] Dunn 1974; [8] Dunn, Edmonds and Fry 1969; [g] Dunn and Fry 1961; [lo] Edmonds, Beuld, Dyro and Hussey 1970; [l11 Esche 1952;

.if:

collected

[l21 Goldman andHueter 1956, 1957 ; [l31 Gramberg 1956; [l41 Hueter 1948; [l51 Hueter 1952; [l61 Hueter 1968; [l71 Kessler 1973; [ 181 Mayer and Vogel 1965 ; [l91 Mountford and Wells 1972a; [20] Pauly and Schwan 1971; [21] Pinkerton 1949; [22] Pohlman 1939; [23] Schneider, Muller-Landau and Mayer 1969; [24] White and Curry 1975.

to be applied. Moreover, the data are so sparse that it would not be realistic to limit attenuation to tissues from man. Similar mammalian tissues are here grouped together, regardless of species. It is seldom possible to take proper account of temperatures a t which the measurements were made : in some tissues the temperature coefficients of attenuation are positive,inothers they are

Ultrasonics in Biology Medicine and

635

negative. The 'freshness' of the tissues is another factor which may be uncertain. Attenuation does seem to be independent of intensity, however, a t least at low values of intensity, provided that heating is taken into account. It is apparent from fig. 1 that, for biological tissues in the frequency range 0.1-50 MHz, c g f h

(14)

where g and h depend on the characteristics of the particular tissue and the conditions of measurement(such as temperature) and have fairlyconstant values over limited ranges of frequency. The value of h is generally only a little greater than uuity. It is impossible to separate the datafor different soft tissues, except to note that the attenuation muscle in seems to be rather greater than that in other soft tissues. In comparison with most non-biological materials, the attenuation in soft tissues has two remarkable features. Firstly, the attenuation coefficients are surprisingly high. Secondly, the linear dependence of attenuation coefficient on frequency cannot be explained in terms of classical absorption (which is due to viscosity and in which the dependence is quadratic). There isnow adequate evidence that this behaviour is due to relaxation processes. The relationship between both the absorption per wavelength (pax)and the velocity to the relative frequency (fif,), for a medium with a single relaxation frequency f,, is shown in fig. 2. The magnitude of (pah)maxdepends on the

Fig. 2. Frequency dependence of absorption per wavelength and velocity in a medium with a single relaxation process.

proportion of the inputwave energy which is involved in the relaxationprocess. Although a single relaxation process has an associated monotonic frequency, a t higher frequencies the classical absorption mechanism due to viscosity becomes increasingly important. It can also be seen in fig. 2 that the velocity increases across the relaxation region from a low-frequency value c. to a high-frequencyvalue c,. This phenomenon of dispersion is apparent as an increase in the stiffness of the medium a t frequencies above fr.

636

P. N . T.Wells

For a medium with a single relaxation frequency, it may be shown (see, for example, Markham, Beyer and Lindsay 1951) that :

and

where wp = 271fP. Hence, approximately,

c-c. a pa.

(17)

It was pointed out by Schwan (1962) that a variation of as little as 20% in the value of ah over a frequency decadecould be accounted for by theexistence of two relaxation processes, one relaxation frequency occurring nearthe lower end of the frequencyrangewith theother a t amuch higher frequency. Although the relaxation processes which are involved are not fully understood (Wells 1975), a more plausible explanation is that there is a spectrum of more than two relaxation processes (Schneider, Muller-Landau and Mayer 1969). 4.2. Velocity Most of the published data for velocity in biological materials are given in fig. 3. It is difficult to draw conclusions from these data, for the same reasons that make it hard to interpretabsorption data. Soft tissues have similar values of velocity (about 1500 ms-l), close to that of water. Velocity dispersion has not been measured except in haemoglobin solution (Carstensen and Schwan 1959). The velocity is about 1530 ms-1 in a solution of 0.165 gml-1 of haemoglobin at 15 "C ; the velocity increases by about 1.60 m s-1 linearlywith frequency over the range 200 kHz-l0 MHz. Substitution in eqn (17) gives Cl0

- Cl 21 100ll

(18)

where cl and cl0 are the velocities (in m S-1) a t 1 and 10 MHz respectively, and a1 is the absorption coefficient (in dB cm-l) at 1 MHz. Assuming that the same consta,nt of proportionality is appropriate, in soft tissues the velocity dispersion would be about 1 m s-1MHz-1. Thisis negligible in comparison with the experimental ranges and errors of measurement for 'solid' tissues, as illustrated in fig. 3. 4.3. Attenuation and velocity in lung and bone As shown in figs 2 and 3, in comparison with other tissues lung has a high attenuation coefficient and SL. low velocity (Dunn and Fry 1961, Dunn 1974). The rate of increase in velocity with frequency is much greater than could be accounted for by relaxation ; it could be due to anincreasing proportion of the wave energy being channelled through soft tissue paths, rather than across gas-filled cavities.

-

637

Ultrasonics in Medicine and Biology I C o l l a g e n[ l 6 1

a

Bone [6,9,11, 16,18,19.20,23

1

E y e lens[2,1L1

m

M u s c l e [ l , L , ? o , ~ l ,12,15 , l 7

I

1

I Foetalhead[2L

I Spleen [10,11.12,17 1

[L,10, 17. 2 5 1

Nerve tissue

I Kidney [10,11.12.17 1 I

Blood L31

I Liver [IO.11, 12,17 l 115 1

I B r e a s t ,l a c t a t i n g

I

Limb, soft tissue[l71

I Mllk [ l 5 1 37OC Normalsaline[221 I Haemoglobinsolution 15 1

20°Cll

I Eye,vitreous [2,1L1

I Breast,premenopausal

L151

I Eye, a q u e o u s [ 2 , 1 & 1

l

I Cerejrosplnalfluid[21 2 0 ’ ~II 3 7 ’ ~ Water 1131

I Breast,postmenopausal Fat

I

L.1’0.15

L151

1

I1MHz 1 5 M H z Lung17,8l

I

1000 500

I

I

I

l

I

I

1500

2000

2500

3000

3500

Velocity ( m

I LOO0

i L500

S”]

Fig. 3. Velocity in normal biological materials : collected data, showing published ranges for each material. Data sources: [l] [2] [3] [4] [5] [6] [7] [S] [g]

[lo] [l11 [121

[l31

Bakke and Gyfre 1974; Begui 1954; Bradley and Saoherio 1972; Buschmann, Voss and Kemmerling 1970 ; Carstensen and Sohwan 1959; Craven, Constantini, Greenfield Stern and 1973; Dunn 1974; Dunn and Fry 1961; Floriani, Debevoise and Hyatt 1967; Frucht 1953; Goldman and Hueter 1956, 1957; Goldman and Richards 1954; del Grosso and Mader 1972;

[ 141 Jansson and Sundmmk 1961 ; [ 151 Kossoff, Fry and Jellins

[l61 [l71 [1 S] [l91 [201 [21]

[22] [23] [24] [25]

1973; Lees 1971; Ludwig 1950; Martin and MoElhaney 197 1 ; Rich, Klinik, Smith and Graham 1966 ; Theicman and Pfander 1949; van Venrooij 1971; Vigoureux 1952; Wells 1966; Willocks, Donald, Duggan and Day 1964: Wladimiroff, Craft and Talbert 1975.

I n addition to having a high attenuation coefficient, bone also has a high velocity. The attenuation coefficient of bone is roughly proportional to the square of the frequency, up to about 2 MHz; above this frequency there is a

638

P. N . T.Wells

lower power dependence onfrequency (Hueter 1952). This suggests that scatteringmay be animportant loss mechanism at low frequencies. The behaviour of compact ivory bone seems to be free from these complexities, presumably because of its simpler structure (White and Curry 1975). 4.4. Scattering

The total disturbanceof the incident wave interacting withan inhomogeneous medium may be considered to be due t o interference between the incident wave and a particular set of Huygen’s wave sources distributed through the medium. The interfering radiation pattern of the obstacle is called the scattered wave field. Scattering of waves by an obstacle is a function of the scattering crossof the totalpower scattered section S of the obstacle. This is defined as the ratio bythe obstacle tothe incident wave intensity.Threesituationsmay be distinguished (obstacle size) > (wavelength) : S = 1 ;

(19)

(obstacle size) < (wavelength) : S cc k4a6;

(20)

(obstacle size) N (wavelength)

(21 1

and where a is the radius of the scatterer and k = 2 ~ f . I n medical ultrasonics these situations correspond respectively to the extensive smooth surfaces of large organs, to small targets such as blood cells and to structures such as blood vessels and muscle fibrils. The scattering from permanent structures, in which the relative positions of the elements remain constant, is a deterministic quantity. This oondition may be approximated in living solid tissues, and the structural ordering of such tissues may be studied by ultrasonic waves using methods analogous to X-ray crystallography (Hill 1974). Nicholas and Hill (1975) have investigated the possibility that theultrasonic diffraction patterns of different tissues might be sufficiently different to allow them to be distinguished. The preliminary results are promising, but extension from in vitro experimentation to in vivo clinical practice would be made difficult by the uncertainty in the position of the resolution cell. 4.5. Characteristic impedance Table 1 gives typical values of density p for several biological materials,

and the corresponding values of characteristic impedance Z = pc calculated from the data in fig. 3. These are bulkvalues:their relevance to clinical applications of ultrasoundislimited to estimations of order-of-magnitude calculations of organ boundary effects. Scattering from discontinuities within tissue structures depends on localized variations in characteristic impedance, and thecollagen content of connective tissue is likely to be an important factor (Fields and Dunn 1973).

Ultrasonics in Medicine and Biology

639

Table 1. Densities and characteristicimpedances of some biological materials

Material Blood Bone Brain Fat Kidney Liver Lung Muscle Spleen Water

Density (g m1-l) 1.06 1.38-1.81 1.03 0.92 1.04 1.06 0.40

1.07 1.06 1.00

Characteristic impedance (kg m--2S-l) x 108 1.62 3.75-7.38 1.55-1'66 1-35 1.62 1-64-1.68 0.26 1-65-1.74 1.65-1.67 1-52

5. Pulse-echo diagnostic methods 5.1. Physical considerations

The ultrasonic pulse-echo method depends on the estimations of the ranges of echo-producing targets lying along the ultrasonicbeam, from measurements of the times of flight of ultrasonic pulses transmitted to and reflected from the targets. Within the limitations imposed by noise and by the maximum transmitted power, the maximum useful dynamic range of the echoes received in conventional medical diagnosticpulse-echosystems is about 100 dB.This dynamic range is sharedbetween the variations inecho amplitude a t particular ranges and the attenuation of echoes which increases with distance. Compensation for attenuation can be partially provided by the application of swept gain. I n practice, at any particularrange, an echo amplitudevariation of about 30 d B is the maximum which may usefully be employed, since the azimuthal resolution is unlikely to be acceptable with a larger dynamic range. Therefore, around 70 dB is available to provide swept gain compensation for attenuation. The resolution of any imaging system may be defined in several different ways. The usual definition is that the resolution is equal to the reciprocal of the minimum distance (in range or in azimuth) between two point targets, a t which separate registrationscan just be distinguishedon the display, An alternative definition, equivalent in concept but usually more convenient in ultrasonic diagnostic practice, is that the resolution is equal to the reciprocal of the distance which appears on the display to be occupied by a point target in the field. Measurements based on this definition avoid problems which arise due to interference between waves scatteredbytwo closely spacedpoint (or line) targets. The resolution cell is the volume of material within which the interaction providing the data takes place. Except in simple and idealized situations, the dimensions of the resolution cell depend on the range of the target. As shown

640

P.N . T.Wells

in fig. 4, the length of the resolution cell depends on the duration of the ultrasonic pulse, and its width, on the width of the ultrasonic beam. The effective values of the pulse duration and the beam width are determined by the dynamic range lying between the maximum echo amplitude and the threshold of the diagnostic system.

Echo

de

~T;'I!

L

4

Fig. 4. Factors determining the dimensions of the resolution cell in a n ultrasonic pulseecho system. The axial length of the resolution cell corresponds to the duration of the echo pulse envelope which exceeds the threshold level of the system: this is illustrated in the bottom right-hand side of the diagram. The diameter of the resolution cell is equal to the effective beamwidth, which is determined by the system threshold as illustrated in the top left-handside of the diagram andwhich depends on thespace-position along the central axis and the time-position within the pulse. The diagram shows how the maximum diameter of the resolution cell is determined.

Biological softtissues haveattenuation coefficients which increase with ultrasonicfrequency.The energy in an ultrasonic pulse is spread over a spectrum of frequency. Consequently the higherfrequency components of the pulseenergy are preferentially attenuatedduring propagationthrough tissue, This has two main effects. Firstly, t'he pulse becomes stretched as it

Ultrasonics in Biology Medicine and

641

propagates. Secondly, in a wide-band system, the rate of swept gain which needs to be applied to compensate for attenuation decreases with increasing range. I n apracticalsystem,asweptgainrate of 1-3 dB cm-lMHz-l is generally satisfactory, where the frequency is the zero-crossing frequency at the centre of the transmitted pulse. These several considerations need to be taken into account in compromising between resolution and penetration (i.e. ma,ximum range). With a plane disc transducer, a satisfactory compromise is generally achieved at low megahertz frequencies with a transducer diameter of 20 wavelengths and a penetration of 200 wavelengths; with a dynamic range of 10 dB, the range resolution is about 1.5 wavelengths, and the azimuthal resolution, about 15 wavelengths a t a range of 120 wavelengths. At higher frequencies, it is better slightly t o increase the aperture of the transducer. The widthof the ultrasonic beam in the near zone can be reduced-and hence, the azimuthal resolution can be improved-over a limited range (the depth-offocus) by the application of focusing. With a fixed focal length system, the ultimate limitationto theimprovement in resolution at thefocus with increasing aperture is set by the wavelength of the ultra,sound, but the depth-of-focus becomes less as the aperture is made larger. One of the fundamental limitations to the pulse-echo method in medical diagnosis is due to multiple reflection artifacts. Multiple reflections arise when the echo from a structure travelsalong an extended route by repeatedreflection from one or more interfaces. These multiple reflections result in registrations on the display at positions where there are nocorresponding reflecting structures within the patient. Multiple reflection artifacts duet o gas-containing structures and to bone, and to the layers of the abdominal wall, are perhaps the most common. Although they are troublesome, they can usually be recognized quite easily from their regular spacings. I n addition to causing multiple reflection artifacts, both gassy and bony structures have higher attenuations and very different velocities to those of soft tissues, and their presence seriously restricts ultrasonic investigations. 5.2. Pulse-echo diagnostic instrumentation

The signal chain of a typical pulse-echo ultrasonic system is shown in fig. 5 . The various components of the system are described in detail elsewhere (Wells 1977a). I n essence, the clock determines the pulserepetitionrate of the system : a typical value is 1000 s-l. The transmitter generates a short-duration unipolar pulse with a typical rise-time of 50 ns; this pulse is fed through the attenuator which controls the overall system sensitivity to the transducer in the ultrasonicprobe.Theradiofrequency amplifier hasalinearamplitude characteristic and a frequency bandpass centred on and equal in width to the nominal ultrasonic frequency. The gain of the RF amplifier is controlled by the sweptgaingenerator, thus compensatingfor the increasing attenuation of later-arriving echoes; typically the maximum gain is around 50 dB. The fullwave demodulator has a maximum dynamic range of around 40 dB, but this

P.N . T . Wells

642 Repeililon :c!e

c.ocx

S b e p t galn r a t e

W

l.

Swept 5 a 1 n "gererator

Edge enhancewent

1'

Trcrsrm:!er

CIS31Jy'

Fig. 5. Block diagram of the signal chain of a typical pulse-echo ultrasonic system.

canbereduced by the suppression control if it is desired to reject smallamplitude signals. The video frequency amplifier mayhavea compression (e.g. logarithmic)amplitudecharacteristic, tomatchthe signal dynamic range at the output of the demodulator to that of the display; it may also process the signal, for example by leading edge enhancement of echo wavetrains. The maximum gain of the video amplifier is typically around 40 dB. 5.2.1. Displays. The various displays used in ultrasonic pulse-echo diagnosis are illustrated in fig. 6. In the A-scope, the probe is often hand-held. The ultrasonic timebase is on the x-axis of the display. In the ophthalmologicalexample in fig. 6, the transducer is spaced away by means of a water-filled tube from the surface of Transducer

Articulated

Probe

Water-filled

1 :

Probe

Movement pattern

-2:, Focusedbeam

I

E -I

id1 (cl la I Fig. 6. Displays corresponding to variousmethods of scanning. ( a ) A-scan. ( b ) T-P recording ("scan). (c) Two dimensional B-scan. ( d ) C-scan.

Ultrasonics in Medicine and Biology

64 3

the structure under examination: this allows echoes to be detected from the tip of the probe. The timebase is triggered after a corresponding fixed delay from the clock pulse. The output from the video amplifier is displayed on the y-axis. There are several methods of time-position (T-P) recording. I n fig. 6, the "scope illustrated has a hand-held probe. The ultrasonic timebase is on the y-axis, with the output of the video amplifier used to control the brightness of the display by z-modulation. A timebase with a relatively slow sweep time is applied to the x-axis; around 3 S is suitable for cardiological studies. An alternative and a very convenient method of timeposition recording uses a strip of photographic paper continuously driven past a cathode-ray tube with a fibre-optic faceplate. Conventional two-dimensional B-scanning, as illustrated in fig. 6, depends on the movement of the ultrasonicprobe by hand across the patient in a controlled scanplane.Thescanningarrangement shown is one of several designs. The x- and y-axes aredrivenbyseparate ultrasonictimebases, triggered by the clock, with velocities determined by a resolver in the scanner which measures the direction of the ultrasonic beam across the patient. The x- and y-coordinates of the start of the timebases on the display are controlled by x- and y-resolvers in the scanner. The output from the video amplifier is arranged to z-modulate the display. A two-dimensional image is formed by storing the displayed echo information whilst the probe is moved across the patient. This storage may be accomplished photographically, or on a directview electronic storage tube of the transmission-control or bistable type, or on a scan converter. The scan converteris widely used in thisapplication, because i t has a wide dynamicrange (good grey-scale), it can be operatedin an equilibrium writing mode (which minimizes the non-uniformity in the brightness of the image resulting fromunskilled scanning), andit has a high resolution (which permits post-scanning 'zooming'). An important class of scanner uses water-bath coupling between the transducer and the patient. In an advanced instrument(Kossoff, Carpenter, Robinson,Radovanovich andGarrett 1976), eightlarge-apertureannulararrays, eachwithswept focusing, are used in sequence automatically to produce pictures a t a relatively high speed. I n C-scanning, the probe is moved in a regular pattern (e.g. a raster) in a plane parallel to that being investigated in the patient, as shown in fig. 6. The depth of the scan plane is determined by a fixed time gate, and, in order to obtain high resolution and to minimize the effect of specular reflections, a wide aperturetransducerisoften used, range-gated atthe focus. The C-scope is not yet much used in routine clinical practice. 5.2.2. Real-time scanners. The maximum pulse repetition rate which can be used in an ultrasonic system is limited by the timewhich is required to obtain echoes from structures situated at the maximum range of penetration. For example, the echo delay associated with a range of 150 mm (a typicalmaximum range in cardiology) isabout 200 p . I n order toavoid artifacts in this situation,

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it would be necessary to allow an interval of about 500 p to elapse between the transmission of pulses. This corresponds to apulserepetitionrate of 2000 S-l. The pulse repetition rate is equal to the product of the image frame rate and the number of lines per frame ; if an image containing 100 lines is tolerable, it can be updated at a frame rate of 20 s-l. Frame rates in excess of 20 S-1 are flicker-free and permit ‘real-time’ visualization. There are several types of real-time scanners, some of which are illustrated in fig. 7 . High-speed mechanical scanners are simple in principle, but they lack versatility and some are difficult to couple to the patient. Multi-element Transducer element Circular track Parabol:c

/

mirror

I

Transduc

tef ron t

\

‘ membrane ( a1

(bl

IC)

(dl

Fig. 7 . Real-time scanningsystems. ( a ) Single-element transducer oscillat,ing about an isocentre in front of its front surface, generating a sector scan. ( b ) Single-element transducer rotating at the focus of a parabolic mirror, generating a linear scan. (c) Multi-element lineararray,withelementsoperatedingroupssteppedby (cl) Multi-element intervals of one elementspacing, generating a linearscan. phased array, with controlled time delays in the transmitting andreceiving signal paths associated with each element, generating a sector scan.

systems, although requiringmore complicated electronics and generally having poorer resolution, side-lobe performance and near-field capability,are more versatile and easier to couple. In ultrasonic diagnosis, real-time imaging hasthree mainadvantagesin relation to conventionaltwo-dimensionalscanning. Firstly,it allows rapid physiological events to be studied, Secondly, physiological movements (such as those dueto thecardiac pulse or t o foetal breathing)do not distort theimage. Thirdly, it allows the operator more quickly to interpret the anatomical relationships within the volume of the patient which is being investigated. 5.3. Analysis of pulse-echo data The ultrasonic diagnostic methods which have the widest applications in clinical medicine are those which produce images which are interpreted by trained observers whose judgements are based on experience. This has reached an advanced stage of development (Kossoff, Garrett, Carpenter, Jellins and Dadd 1976). Numerical information can be obtained by analysis of pulse-echo data (Wells, Atkinson, Halliwell, Morris, Mountford and Woodcock 1976), which can supplement or even replace ultrasonic and other diagnostic images and tests.

Ultrasonics in Medicine and Biology 645 The time delay between the transmission of an ultrasonic pulse and the reception of its echo can be used to estimate distance if the velocity is known. Clinical applicationsincludeocularbiometry andbrain midline structure localization. The rate of change of structure position provides an estimate of structure velocity,forexampleincardiacvalvestudies.Distancemeasurements can be extended to three dimensions for the estimation of the volume of structures such as the liver. The possibility that the amplitudes of echoes arising fromwithintissue volumes might provide diagnostic information was first tested by Wild and Reid (1952). They showed that the echoes frommalignant lesions inthe breast are greater ;n amplitude than those from normal breast tissue, and vice versa with benign lesions. Subsequently Schentke and Renger (1964) reported that both the mean echo amplitude and the echo-spacing frequency increase in several diseases of the liver including cirrhosis. Wells, McCarthy, Ross and Read (1969) studied the echo amplitudes from normal and cirrhoticliver: they calculated an index

A

h,/T

=

(22)

i

where h, is the amplitude of the ith internal echo and T is the duration of the echo wavetrainsubjected to analysis.Theapplication of this method t'o envelope-detected A-scans gave an accuracy of about 80% in separating the two groups on the basis of the higher values of A in cirrhotics. Mountford and Wells (1972a, b) refined the method by making sufficient measurements to give statistical significance to estimatesforindividuals.Theresultsfor echo amplitude are shown in fig. 8 ; the separation between the cirrhotic and the normal groups is equivalent to 6 dB. The decrement of echo amplitude from within the liver was found to be 1.1 dB cm-1MHz-1. This is equivalent to the attenuation coefficient if the distribution of scatterers within the liver may be Norxcl

*Or

Relative ecno amphiude ( d B 1

Fig. 8. Histogram showing the hepatic echo amplitude distributions of individuals with normal livers and with cirrhosis. The echo amplitude range is divided into intervals of 0.882 dB (data of Mountford and Wells 1972b).

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assumed to be homogeneous, and the value is in good agreement with more conventional measurements. There is now much interest in the possibility of identifying tissues from their ultrasonic characteristics. To this end research into the analysis ofpulseecho data is being carried out by many investigators (Linzer 1976). 5.4. Clinical applications It is impossible, in a short review such as this, to give more than a brief mention even of the most important applications of pulse+xho ultrasonic methods in clinical practice and to present a few typical results. For more information, see Wells (1977a,b). 5.4.1. Angiology. Many of the abdominal vessels, including the aorta, inferior venacava,portal vein and hepatic, mesenteric and renal vessels, can be visualized (Leopold 1975). Aneurysms and thrombi can be excellently demonstrated. High-resolution images of superficial vessels, such as those in the neck, have recently been produced (Marich, Green, Evans and Harrison 1976). 5.4.2. Cardiology. The possibility that clinicallyuseful information about cardiac structure and function might be obtained from ultrasonic pulse-echo measurements was first proposed by Edler and Hertz (1954). Nowadays timeposition echocardiography is an essential adjunctin cardiological investigations. I n addition to measurements of cardiac valve function (Gramiak and Waag 1975)"some typical recordings are shown in fig. 9"the method is useful in the diagnosis of left atrial myxoma (Edler 1965), congenital heart

Fig. 9. Mitral valve echocardiograms. (a)Normalmitral valve. ( b ) Mitral stenosis. The dot markers (most clearly visible in the part of the recording corresponding to the normal left ventricular cavity) are spaced at intervals of 1 S and B mm.

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disease (Chesler, Joffe, Beck and Schrire 1971), pericardial effusion (Feigenbaum, Zaky and Waldhausen 1967) and hypertrophic cardiomyopathy (Feizi and Emanuel 1975). Two-dimensional imaging of the heart, either by realtime scanning (Bom, Hugenholtz, Kloster, Roelandt, Popp, Pridie and Sahn 1974, von Ramm and Thurstone 1976) or by Ecc-gating of aconventional scanner (Waagand Gramiak 1976), is being increasingly used and is particularly useful in elucidating unusual anatomy. 5.4.3. Endocrinology. Thethyroidglandisreadily accessible to ultrasonic scanning (Jellins, Kossoff, Wiseman, Reeve and Hales 1975). Not only is it possible to detect, and to distinguish between, cysts and tumours, but also echography is complementary to scintigraphy, which revealsfunctioning thyroid tissue. The normal suprarenal gland is difficult to visualize, except on the left side in children (Lyons, Murphy and Arneil 1972). I n patients with suspected suprarenal tumours, ultrasonic scanning is a useful preliminary to venography (Davidson, Morley, Hurley and Holford 1975). 5.4.4. Gastroenterology. Howry and Bliss (1952) were the first to publish twodimensional ultrasonicscans of the liver. Since the advent of good-quality grey-scale instruments, echography hasbecome beyond question more valuable than scintigraphyinthe management of the cancer patient (Taylor and Carpenter 1975). It is also very useful in distinguishing between medical and surgical causes of jaundice.Pancreasscanning is helpful asa confirmatory investigation in patients with malignant disease (Lutz 1975). Finally, ultrasonic scanningis more reliable and sensitive than clinical examinationin detecting ascites (McCarthy, Wells, Ross and Read 1969). 5.4.5. Neurology. Midline echography has been practisedfor twenty years, since the original publication of Leksell (1956). The method using the A-scope is not objective, and its accuracy depends on the skill and clinical acumen of theoperator(White,Kraus, Clark and Campbell 1969). Instruments using logic circuitsautomatically to identify and to measure the position of the midline structures (e.g. Williams 1973) appear, however, to be clinically useful (1974). Other according to evaluationssuchas that of WhiteandHanna transcranial studies of the brain are virtually useless, due to uncertainty in the position of the resolution cell. 5.4.6. Obstetrics and gynaecology. Donald (1974) thinks that a routine ultrasonic examination should be made of every pregnancy, between weeks 20-24, to establish maturity, to detectmultiple pregnancy and to locate the placenta. A typical scan is shown in fig. 10. Of these procedures, the establishment of maturity isof mostinterest to physicists-and, for different reasons, to it is based on the measurement of thebiparietal obstetricians-because diameter of the foetus. Some scanners have special displays with electronic calipers to facilitate the making of this measurement(Hall,Fleming and Abdulla 1970). As scanners are improved, so it is becoming possible with more

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certainty to detect foet.al abnormalities such as spina bifida (Campbell, PryseDavies, Coltart, Seller and Singer 1975). Another development of potentially great importance is the monitoring of foetal breathing (Dawes 1974), since this

Fig. 10. Typical two-dimensional ultrasonic scan, showing a foetus at 26 weeks’ gestation. This longitudinal scan through the maternal abdomen was made with a ‘grey scale’ system.

may give an index of foetal status during the second half of pregnancy. I n gynaecology, ultrasound can be helpful in distinguishing betweenmasses (Donald 1974). 5.4.7. Oncology. Ultrasonic visualization of malignant lesions may be of great value in planning radiotherapy (Brascho 1974) and in assessing the effects of radiotherapy and chemotherapy (Kobayashi, Takatani, Hattori and Kimura 1974). Ultrasonic investigations of the breast have so far been rather disappointing. Although quite pretty pictures have been produced (e.g. Jellins, Kossoff, Reeve and Barraclough 1975) the effort needed to obtainand to interpret them is so great that the method could not be used for screening even a highly selected group of the population, which is necessary for an impact to be made on the breaet cancer problem. 5.4.8. Ophthalmology. Intraocular tumours can be detected with the A-scope, and Ossoinig (1974) has described criteria by which different types of tumour may be identified according to their echo-producing characteristics. The A-scope is also valuable in the measurement of the axial lengths of the components of the eye (Giglio, Ludlam and Wittenberg 1968). Almost all twodimensional ophthalmic scanners use water-bath coupling. The images are clinically useful in the diagnosis of many kindsof ocular abnormalities (Sutherland and Forrester 1974). 5.4.9. Urology. Since the publication of a paper by Holmes (1966), there has been an increasing interest in ultrasonic pulse-echo scanning of the kidney. Both kidneys can normally be seen in either longitudinal or transverse posterior scans and the right kidney can be seen from the anterior through the liver. The main clinical application of the method is when intravenous pyelography

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has failed, and it is also useful in distinguishing between cystic and solid lesions (Mountford, ROSS,Burwood and Knapp 1971). Scanning the bladder allows an estimate to be made of urinary volume (Alftan and Mattson 1969), and it is also useful for staging bladder tumours (Barnett and Morley 1971). Enlargement of the prostate can be assessed either by radial scans made with a rectal probe(King, Wilkiemeyer, Boyce and McKinney 1973) or by conventional scans made through the urine-filled bladder (Whittingham and Bishop 1973). 6. Dopplerdiagnosticmethods 6.1. Transmission methods The difference At between the upstream and downstream ultrasonic pulse transit times measured between two transducers at a fixed distance d apart depends on the flow velocity. If 6' is the angle between the direction of flow and the centralaxis of the ultrasonic beam,

v 2d = At c"(

cos 6')

(23)

where v is the flow velocity and c$ v. This relationship has been used as the basis of ultrasonic flowmeters, both for blood (Franklin,Baker, Ellis and Rushmer 1959) and gas (Blumenfeld, Wilson and Turney 1974). Whilst being elegant in principle, in practice A, is very small in relationto the transit timet (for example, if d = 20 mm in blood, 6' = 15" and v = 100 mms-l, A, is about 0.01% of t ) . Themeasurement of suchasmall difference poses formidable problems in instrument design and, at least for blood flow, the incentive to solve these problems has largely disappeared with the development of Doppler reflection methods (see section 6.2). 6.2. Rejlection methods

The Doppler shiftf D in the frequency of reflected ultrasound may be used as a measure of the vector velocity of the movement of the reflecting surface or ensemble. If y is the angle between the direction of flow and the effective axis of the ultrasonic beam,

v

= -fDc/(2fcosy)

(24)

where v is the flow velocity and C ~ V For . example, if the measurement is being madein blood, fD = 1 kHz when v = 100 mms-1, f = 8 MHz and y = 15".

The reflection (backscattered) ultrasonic Doppler method is used in many instruments, simple and complex, for measuring the velocities of structures and flow within the living body, 6.2.1. Continuous wave systems. The block diagram of acontinuous Wave Doppler system is shown fig. in 11. It was first described in English by Satomura (1957). I n essence, the transmitter is a continuouswave oscillator which Causes the transmitting transducer to emit ultrasonic waves into the patient. Echoes returning from within the patient are detected by the receiving transducer. These echo signals are equal in frequency to that of the transmitter when the

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reflections are from stationary structures (and so are cross-talk and leakage signals), but they are shifted in frequency by the Doppler effect if the corresponding reflecting structures(or ensembles) havecomponents of velocity along the ultrasonic beam. All the signals are amplified together by the radio

E l e c t r i c a l l e a k a g e andultrasonlccrosstalk

+transducer

Skin

Blood vessel walls

\ \ \ ' \

Fig. 11. Block diagram of typical continuous wave ultrasonic Doppler system; a blood flow velocity detector is shown here.

frequency amplifier (gain typically30 dB), so that the input to the demodulator is made up of signals both equal in frequency to that of the transmitter and Doppler-shifted by reflection from moving targets.The output from the demodulator is fed through alow-pass filter which extracts thesignals with the difference frequencies. These signals are amplified, and either an operator listens to them or they are analysed, for example, by a zero-crossing frequency 1975), soundspectrograph (Woodcock 1970, Coghlan, Taylor meter(Lunt and King 1974), maximumfrequency follower (Flax, Webster and Updike 1971, Sainz, Roberts and Pinardi 1976) or average velocity computer (Arts and Roevros 1972). A typical frequency spectrum is shown in fig. 12. 6.2.2. Directional detectors. Simple Doppler systems merely measure the magnitude of the frequency difference between the transmitted and received ultrasonic signals and not the sign of the difference. As indicated in eqn (24), this sign carries the information about the direction of the movement of the target, either towards or away from the probe. The directional information is of value in some diagnostic situations. If the Doppler signal always consists, a t any instant in time, of movement only in direction, a phase quadrature detector and logic circuitry can be used to indicatewhetherthe movement is inthe forwardorreversedirection (McLeod 1967). If signals corresponding to both forward and reverse movements exist simultaneously, however, quadrature detection is unsatisfactory. Single sideband andheterodyne detectionmethods arethenappropriate (Coghlan and Taylor 1976). 6.2.3. Range-measuring Doppler systems. Unmodulated continuous wave Doppler systems are unable to measure the distance between the probe and the

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651

moving structure. It is possible to code the transmitted signal so that the pulse-echo transit time may be determined, thus allowing the range of the target to be estimated. The simplest method of coding is to pulse the transmitter (Wells 1969, Peronneau, Hinglais, Pellet and Leger 1970). The upper frequency limit (which is related to the maximum value of target velocity x

c

0 0

-B

r

8 21 C

3 U

e

L

a

c

5i

-a L (U

a

(3

Fig. 12. Frequency spectrum of Doppler signal detected transcutaneously by an ultrasonic probe positioned over a normal common femoral artery. Positive Doppler shift signals correspond to flow away from the heart, and vice versa. The pulse wave travels along the limb at 8 k i t e velocity. Arterial disease may be classified according to the transit time of the pulse wave along the limb, and the modification of the maximum frequency envelope ofthe frequency spectrum.

which can be measured) which can be detected without ambiguity depends on the sampling rate. The maximum sampling rate is itself limited by therequired penetration. I n principle, thislimitation can be avoided bythe use of a random code on a continuously transmitted carrier (Waag, Mykleburst, Rhoads and Gramiak 1972), but neither this nor related techniques seem to have been developed to the stage of clinical application. A fundamental property of all methods designed to measure both velocity and position is that the velocity and range resolutions are inextricably linked (Atkinson 1976). I n physical terms, if the position of a target is precisely measured, the velocity cannot simultaneously be measured because to do this requires motion to occur. 6.2.4. Two-dimensional Doppler imaging. I n vascular investigations it is sometimes helpful to have a two-dimensional map showing the positions of blood vessels. An example of this type of image is the contrast angiogram familiar in radiology. The Doppler shift in the frequency of backscattered ultrasoundis a special characteristic offlowingblood, andthishas been

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652

exploited by Reid and Spencer (1972) in constructing an instrument for noninvasive mapping of vessels in which blood is flowing. The basic principles, subsequently developed by Spencer, Reid,Davisand Paulson (1974), are illustrated in fig. 13. The probe is arranged so that the ultrasonic beam is at

Doppler probe

Positlon sensing resolvers

p!

x andy lmeb se eneraqors

Bloodvessel

Focused ultrasonic beam

Dlsplay

Fig. 13. Two-dimensional scanner for mapping blood flow distributions from ultrasonic Doppler signals detectedtranscutaneously(basedon Spencer, Reid, Davis and Paulson 1974).

least slightly inclined to the direction of flow in the vessels which it is desired to visualize. When the beam passes through moving blood, the Doppler detector generates an output which is filtered (to eliminate artifacts due t o low velocitymovements) and, provided that the output exceeds apreset threshold level in amplitude, the electron beam of the display is switched on. A two-dimensional map showing those regions in which flow is detected is constructed on the display by scanning the probe over the area of the patient’s skin beyond which the blood vessel lies. When arteries and veins are close together, a directionally sensitive circuit can be arranged to display where flow is in the opposite direction t o that in the vessel (usually the artery) under investigation. A further refinement of the method is t o incorporate a pulsed Doppler. This measures the range of the flowing blood and allows vessel lumen cross-sections to be visualized (Mozersky, Hokanson, Baker, Sumner and Strandness 1971, Fish 1972). The combination of pulse-echo imaging of blood vessels with pulsed Doppler measurement of blood velocity has opened up the possibility of studying both vascular structure a,nd function in real time (Marich et al. 1976). 6.3. Clinical applications 6.3.1. Angiology. In theearliest application of the Doppler blood flow detector

in the evaluation of arterial disease, the course of the artery was traced along the limb whilst the investigator listened for sudden changes in the characteristics of the signal (Strandness, Schultz, Sumner and Rushmer 1967). Quantitative methods of grading arterial disease have since been developed.

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For example, Lewis, Papathanaiou, Yao and Eastcott (1972) used the Doppler motion detector to detect arterial wall pulsations and thus to measure blood pressure along the leg by means of a sphygmomanometer; they showed that the blood pressure at the ankle is not much altered by exercise in the normal, but in disease it falls typically by 50-100%. Woodcock (1970) was the first to explain the significance of changes in the time-velocity waveform, based on simultaneous observations of the ultrasonic Doppler arterial signals at separate sites on the limb. He determined the pulse transit time between the two measurement sites, the pulsatility index (defined as the sumof the maximum oscillatory energy of the Fourier harmonics of the Doppler frequency spectrum, divided by the mean energy over the cardiac cycle) at eachsite,andthe dampingfactor (defined astheratio of the pulsatility indices at theupstream anddownstream sites)between the two sites. to grade Fitzgerald, Gosling and Woodcock (1971) used thesequantities isshort or long, arterial disease into four classes. Thus,thetransittime according to whether the length of the collateral is short or long; the damping factor is small or large, according to whether the diameter of the collateral is large or small.Recentstudieshaveintroduced the concept of the transfer functions of peripheral arteries in the leg. Calculation of the impulse response from the maximum frequency envelopes of the Doppler signals detected at two sites on the leg (Morris, Woodcock and Wells 1975) is possible with diseased patients ; and the occurrence of time-separated peaks in the impulse response suggests the existence of separate vascular pathways (Woodcock, Morris and Wells 1975). Calculation of the impulse response isunstablewithnormal transferfunctions, butotheranalyticalmethods, involving s- or z-transfer functions, are giving promising results (Morris, Owens, Payne, Wells and Woodcock 1976). Methods of surveying the cerebral circulationby means of simple continuouswave Doppler instruments have been described (e.g. Mol and Rijcken 1974, Muller and Gonzalez 1974, Planiol and Pourcelot 1974). Despite difficulties associated with zero-crossing counting and directional detection, it is possible toobtain clues about suchdisordersascarotidstenosis,arterio-venous aneurysm,subclavianstealsyndrome and innominate artery stenosis. Twodimensional Doppler mapping is useful in studying the carotid arteries (Thomas, Spencer, Jones, Edmark and Stavney 1974), as stenosis and the existence and extent of calcification and plaque can be demonstrated, and for the detection of deep vein thrombosis (Day, Fish and Kakkar 1976). Dopplerstudies of venous flow arelimited bythe absence of natural pulsation. Deep venous thrombosis,particularlyin the upper leg, can be detected from the changing characteristics of the flow signals at the superficial femoral vein, in response to squeezing of the calf or foot (Evans 1971). 6.3.2. Cardiology. In the normal, the jugular venous flow pattern as recorded by a transcutaneous Doppler detector has a characteristic pattern. Specific changes occur in this pattern in patients with tricuspid regurgitation or leftto-right shunt (Kalmanson, Veyrat, Chiche and Witchitz 1974).

654

P.N . T . Wells

The possibility of making transcutaneous measurements of blood velocity in the thoracic aorta using the Doppler method was first proposed by Light (1969). He has used a 2 MHz continuous wave system, and, in his more recent work, he has adopted the suprasternal notchposition for the probe (Light and Cross 1972). The orientation of the aortic archis such that theultrasonic beam can be directed to intersect the axial direction of the blood flow: the highest Doppler shift frequency which is detected then corresponds to the maximum to display the blood velocity.The use of areal-timesoundspectrograph directionally-detected Doppler shift signals allows the operator to obtain the optimum orientation, and to recognize flow signals from branch arteries which, since they serve the head and neck, are in the opposite direction in the ultrasonic beam to theflow in the aortic arch. The clinical usefulness of the method isstill being assessed (Light 1974, Sequira,Light, Cross and Raftery 1976). The part of the aorta which is monitored is close to the heart, so information on left heart function is obtained. Preliminary results indicate that, in any particularindividual, the instantaneouscardiacoutput is proportional to It is the corresponding measured velocity: this is helpful inintensivecare, also possible to derive indices of earlysystolicacceleration,peakvelocity, and the durations of the acceleration and deceleration phases of the systolic period. 6.3.3. Obstetrics. Themovements of the foetalheart can be detectedby means of a simple 2 MHz Doppler system, often a t 9-10 weeks' gestation and almostalwaysafter 12 weeks. Themethodisvaluablein the diagnosis of foetal death in the last 6 months of pregnancy. During labour, the normal foetal heart rate is in the range 120-160 min-l, there is no change in rate during uterine contractions and there is a beat-to-beat variability of 5 min-l

or less. Unfavourable patterns include bradycardiawithoutbeat-to-beat variations or with decelerations duringcontractions, andtachycardiawith decelerations during contractions (e.g. Liu, Thomas and Blackwell 1975). The possible importance of the foetal breathing pattern as anindex of foetal status is mentioned in section 5.4.6. It has recently been demonstrated that continuous wave Doppler shift signals can be detectedtranscutaneously, which apparently arise from the redistribution of fluid within the foetal lung (Boyce, Dawes, Gough and Poore 1976). Theobservation may allow many problems which occur with pulse-echo methods of measuring foetal breathing to be avoided. 7. Other diagnosticmethods 7 . 1 . Transmission methods

The att'empts t o image the brain by two-dimensional display of the intensity of ultrasound transmitted through the skull (Dussik, Dussik and Wyt 1947) failed because of the predominant effect of the transmission characteristics of the skull (Guttner, Fielder and Patzold 1952). Early attempts to use image converters to visualize the shadows of biological structures (e.g. Jacobs 1965)

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were almost as disappointing, for reasons which are not clear in view of the comparative success of an imaging system using acombination of lenses, scanning prisms and a linear array (Green, Schaefer, Jones and Suarez 1974, Zatz 1975). Even when pulsing is used to avoidstanding wave artifacts, however, the images remain degraded by diffraction, refraction and deviation of the ultrasonicbeam.This problem might be reduced bytimedelay spectrometry (Heyser and le Croisette 1974), in which the transmitted ultrasound is coded with a linear frequency sweep, thus allowing the signal which travels along the direct path to be measured. The construction of two-dimensional X-ray images by computed tomography has become a powerful diagnosticmethod(e.g. Brooks and Di Chiro 1976). I n principle, analogous images should be obtainable using ultrasound. Greenleaf, Johnson, Lee, Herman and Wood (1974) tested this possibility, scanning both a phantom and an excised canine heart, using 5 MHz ultrasonic pulses of 2 MHz bandwidth. The images were interesting, but not sufficiently clear to encourage the development of a clinical instrument. Velocity profiles, as distinctfromattenuation profiles, however, have given quite good two1975) ; dimensional reconstructions (Greenleaf, Johnson, Samaoya and Duck these images do not have an X-ray analogue. 7.2. Holography

Ultrasonic holography is a two-stage imaging process. First, an ultrasonic hologram is generated by recording the diffraction pattern of the object in an ultrasonic field, biased by acoherent reference wave. Second,a threedimensional image of the object is created by illuminating the hologram with coherent light. There are many different wide-aperture recording arrangements by which ultrasonic holograms may be generated. These include liquid surface levitation (e.g. Brendon 1974), optical interferometry (e.g. Metherell 1974), the scanned hydrophone(e.g. Aldridge, Clare and Shepherd 1974) and the Pohlman cell it is generally (e.g.Lafferty andStephens 1971). In opticalholography, necessary for thescatteredand reference signals t o be added before the hologram is recorded, because there is no suitable phase-sensitive light detector. The same applies in ultrasonic holography with slowly responding detectors, suchas the surfacelevit'ationmethod. Many ultrasonicdetect'ors, however, respond both to amplitude and phase;if such a detector isused as thereceiver, an electrical signal derived from the transmitting oscillator may be used as the phase reference in the recording process. Much effort has been expended in trying to develop ultrasonic holography for medical imaging. The results have been uniformly disappointing, and it is easy to understand some of the reasons for this. The image has a longitudinal dist'ortion of about 2000 : 1, due to the difference between the wavelengths of the ultrasound used to generate the hologram and the light used for reconstruction. Specular reflections, being of a largeamplitude,maskscattered echoes. There is diffraction in the near fields of the large apertures, of both

656

P.N . T.Wells

the transducers and the object. Phase coherence is destroyed by variations in the velocities in different biological tissues. Ultrasonic waves are deviated and distorted as they travel through biological tissues. Several at least of these problems appear to be insuperable. 7.3. Microscopy

Thepotentialvalue of ultrasonic microscopy has become apparentas a result of the development of two different methods of high-frequency uitrasonic imaging. I n one method, the ultrasonicshadow of the object (a thin section-as in an optical microscope-suspended in a coupling liquid) is cast onto a flat reflecting surface, and this is scanned by an optical probe (Korpel, Kessler and Palermo 1971). The frame rate is 30 s-l. Structural elements of down to around 10 pm in size can be visualized with an ultrasonic frequency of 220 MHz (Kessler 1974). At this frequency it seems that the image contrast is primarily due to variation in characteristic impedance in the object rather than in attenuation. At gigahertz frequencies, thecontrast due toattenuation differences in microscopic structures is much greater than that ata few hundred megahertz, and viscosity (as distinct from relaxation;see section 4.1.1) is the predominant attenuation mechanism. Uniform irradiationand scanning of the whole specimen are impracticable at these frequencies, but Lemons and Quate (1975) have avoided this problem by scanning the specimen with a highly focused ultrasonic beam, as shown in fig. 14. The spherical aberration due to sapphire Mylar film Llquld l w a t e r lc e l l t o perm!: scannlng m o i l o n of speclrren

pi;

/Spec!mermounted

'

31

m y l a r fllrr

CC nm

Fig. 14. Principal components in a scanning acoustic microscope designed to operate at frequencies of around 1 GHz. The lens apertures are f 0.65. The transducers are piezoelectric films. The mylar membrane is mounted on a metal ring attached to a scanning mechanism (after Lemons and Quate 1975).

lenses is below the resolution limit imposed by diffraction, and according to Bennett (1976) the resolution at 1 GHz is around 1 pm. Moreover, the simple configuration of the transmission scanning microscope could be elaborated t o allow scattering in thespecimen to be studied.

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Ultrasonic microscopy promises to open up a new area of histopathology. The tissue properties which are visualized, being different from those which can be seen by light microscopy, may give insight into cellular structure and function; and the technique may in some situations eliminate the need for staining. A resolution of around 0.1 pm may be achieved a t a frequency of 10 GHz; this surpasses that of light microscopy. 8. Biological effects 8.1. Thermal effects Ultrasound is attenuatedasit passes through biological materials.The wave energy which is absorbed is converted to heat, and this results in an increase in the temperature. Many of the so-called ‘biological effects of ultrasound’ are atleast partly due to this heat. Thermal effects are most important at low megahertz frequencies (where significant volumes of tissuecan be heatedrapidly) at mediumintensities (where thetemperature rises are significant, but mechanical effects donotpredominate). Simple examples include the increaseduringirradiation inthe conductionvelocity of the nervous impulse (Shealy and Henneman 1962, Lele 1963) and the effects on active transport (Carney, Lawrence and Ricketts 1972, Chapman 1974). The use of focused ultrasound to induce trackless lesions is discussed in section 9.2. The relationship between intensity and threshold exposure time for lesion production is shown in fig. 15; there is remarkably good agreement

I

90-6

10-3

I

10-2 Time

l

I

I

10”

1

10

(S)

Fig. 15. The relationship between intensity and single-pulse time duration, for threshold lesion production in white matter of mammalian brain (data from laboratories in Illinois, London and Massachusetts, collected byFry, Kossoff, Eggleton and Dunn 1970).

between the resultsobtainedin different laboratories. At intensities below , calculation and experiment confirm that the lesions about 700 W ~ r n - ~both are caused by heat (Pond 1970, Robinson and Lele 1972). Thephenomena which occur a t higher intensities are discussed in section 8.2.1.

P. N . T.Wells

658 8.2. Non-thermal effects

Since the publication of the results of experimentsin which ultrasonic radiation blocked nerves, apparently by non-thermal mechanisms (Fry, Wulff, Tucker and Fry 1950, Fry, Tucker, Fry and Wulff 1951), it has been accepted that heating is not theonly effect of ultrasound on biological materials. 8.2.1.Cavitational effects. Transientcavitation is the phenomenon in which voids suddenly grow from nuclei in the supporting liquid, and then collapse, under the influence of the changing pressure in an ultrasonic field. The whole process occupies less time than thewave period. Nottingk and Neppiras (1950) have developed atheoryto describe the motion of bubbles intransient cavitation. It predicts that cavitationphenomena diminish and finally disappear as the frequency is increased; this has been demonstrated by Esche (1952) and Gaertner (1954). During bubble collapse, a strong pressure pulse is set up in the liquid and high temperature occurs in the bubble. Whether or not transient cavitation occurs in any particular exposure situation depends on several factors. It can occur easily a t low kilohertz frequencies in liquids of low viscosity, but at megahertz frequencies in soft tissues it appears that intensities of around 1 kW cm-2 or more may be required (this value is based on the considerations mentioned in section 8.1). The behaviour of a gas-filled bubble pre-existing in an ultrasound field of intensity below that necessary to cause transient cavitation is known as stable which the surrounding liquid is an cavitation. A resonant system exists in inert mass which is set into vibration, the elasticity being provided by the gas in the bubble (Minnaert 1933); the resonance frequency f, of the system is given by fr =

(1/2xrb)[3y(P0

+ 2u/rb)/Plt (25

)

where r b is the bubble radius, y is the ratio of the specific heats of the gas, P, is the hydrostatic pressure, U is the surface tension and p is the density of the liquid. A resonant air-filled bubble in water at atmospheric pressure has a at lower diameter of about 0.7 pm a t 1 MHz ; it isproportionatelylarger frequencies. One of the effects of stable cavitation is that streaming occurs in the liquid (1959) demonstrated and in the neighbourhood of a resonant bubble. Elder analysed four regimes of streaming, the transitions between which occur with increasing ultrasonic amplitude. A velocity gradient exists in a liquid inwhich the streaming pattern is localized, and in some circumstances this velocity gradient may be high enough to break cell walls or to rupturebiological macromolecules. The process has been quitethoroughlystudied at low kilohertz frequencies with aqueous suspensions (e.g. Neppiras and Hughes 1964, Prit,chard, Hughes and Peacock 1966, Thacker 1973a) and breakagerates correspond to those found to occur in shearing fields with equivalent velocity gradients. The biological effects of stable cavitation at low megahertz frequencies in Clarke, Crowe and Hammick liquids are also quite well understood.Hill,

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(1969) demonstrated that stable cavitation is effective a t these frequencies if the specimen is rotated to neutralize unidirectional forces. Hill (1972) showed that at least about 1000 cycles of oscillation are required for stable cavitation to become effective at 1 MHz, and under these conditions DNA is degraded by intensities of a few watts per square centimetre. I n solid tissues, however, it is clear that stable cavitation has a much higher intensity threshold, if indeed it occurs a t all. 8.3. Frequency dependence of biological effects

It is convenient to make a broad distinction between the mechanisms of the biological effects induced by low-kilohertz ultrasound, where cavitation and streaming are of major and obvious importance, and those responsible for the effects of low-megahertz ultrasound, where many of the interactions are at present not well understood. In the low-megahertz frequency range, some effects, such asdamagetothespinal cord (Taylor andPond 1972) become less important with increasing frequency. I n contrast, the reverse is true with other effects, such as the reduction in cell electrophoretic mobility (Taylor and Newman 1972). 8.4. Curnulafive effects

Biological effects which are apparently similar may be produced by ultrasound delivered under different conditions of pulsing. For example, a t low megahertz frequency, a t a fixed pulse intensity, the integral exposure time to produce nerve block increases rapidly as the pulse duration is reduced from about 7 S to about 3 S, if there is a long interval (about 250 S) between pulses, but if this interval isreduced to 400 ms, pulses of about 100 ms become effective (Fry et al. 1950). With a pulse duration of 10 ms, the threshold for damage to spinal cord associated with constant integral pulse exposure a t a fixed pulse intensityisindependent of the interval between pulses in the range 100-400 ms (Taylor 1970). At low megahertz frequencies, cell killing due to stable cavitation is most marked when the pulse interval is around 100 ms, and killing is suppressed by shortening the pulse duration to 1 ms or less, even when the pulse intensityis increased (Clarke and Hill 1970). It is clearly impossible to make generalizations about the cumulative effects of ultrasound. 8.5. Other biological effects 8.5.1. Effect on blood $ow. Blood flow in small blood vessels may be arrested by irradiation with ultrasounda t less than about 0.5 W cm-2 a t 1 MHz (Dyson, Woodward and Pond 1971). The blood cells aggregate in bands separated by

half-wavelength intervals, and i t seems that theeffect is due to standing waves (Gould and Coakley 1974). 8.5.2. Effectsonmalignanttumours. Although Woeber (1965) reporteda striking increase in the regression rate of experimental tumours simultaneously treated with X-rays and ultrasound comparison in with those treatedby X-rays alone, Clarke, Hill and Adams (1970) concluded that theeffect is not significant.

660

P. N . T.Wells

The heat produced by the absorption of ultrasound might be applicable to treatment of tumours by hyperthermia. There has been a suggestion (Heimburger, Fry, Franklinand Eggleton 1975) that ultrasoundmayactsynergistically to improve the effectiveness of chemotherapy. 8.5.3. Effect on tissue regeneration. The stimulation of tissue regeneration has been reported by Dyson, Pond, Joseph andWarwick (1968). The rate of repair of experimental wounds in the ears of rabbits was increased by up to 30% as a result of a course of 5 min irradiations on alternatedays, using 3.6 MHz , duration 2 ms, duty ultrasound with a pulse intensity of 0-5 W C M - ~pulse cycle 0.2. Heat is almost certainly not responsible for this effect. 8.5.4. Evidence of genetic effects. Thedegradation of DKA in solution is mentionedin section 8.2.1. It is unlikely that this could happen to DNA within the cell. Thacker (1973b) reviewed many of the papers dealing with chromosome damage andultrasoundand concluded that ultrasound causes only lethal lesions on chromosomes ; genetic damage does not occur. Where mutations due to ultrasound do occur, they are likely t o be due to thermal shock. There is evidence, not statistically significant, that 5 h irradiation a t 40 mWcm-2with 2.25 MHz ultrasoundmayinduceteratogenesis(Shoji, Momma, Shimizu and Matsuda 1972), and this possibility should be further investigated.Ultrasonicirradiationinhibitsgrowthby an all-or-none effect on the cell (e.g. Bleaney and Oliver 1972), which seems to be more susceptible during mitosis (Clarke and Hill 1969); there is no recovery process, such as occurs following damage by ionizing radiation. 8.6. Possibility of hazard in diagnostics

It is possible to define pairs of zones on charts on intensity and time, such that in one zone the conditions have been shown to produce biological effects and, in the other, they have not. The results of three literature surveys are shown in fig. 16. Once it has been recognized that aparticularirradiation condition does not fall in the region where biological effects have not been reported,many considerations need to be takeninto accountin assessing ‘hazard’. Clearly there are insufficient data t’o state with any certainty that ultrasonic diagnosis is ‘safe’, but it is certainly justifiable when the information which it might provide could help in the care of the patient. 9. Therapeutic and surgicalmethods 9.1. Physiotherapy Ultrasonic physiotherapy is widely practised. Its scientific basis is poorly understood. Generally the rationale of the method is that the absorption of ultrasound is associated withheatdistributionsin tissues which in some circumstances are beneficial. Moreover, ultrasound may have other desirable biological effects in addition to heat. Why heat should be beneficial has not been properlyexplained, butit does modify blood flow, and everyone is

Ultrasonics in Medicine and Biology

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66 1

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Fig. 16. Three pairs of exposure zones, showing 'safe' and potentially hazardous conditions. Each line separates one pair of zones, the zone below the line being 'safe'. (i) Data collected by Ulrich (1971, 1974); the intensity is the average intensity (equal t o the product of the peakpulse intensity and the duty cycle), and the time is the total timeof the exposure, including intervals between pulses if appropriate. (ii) Data collected by Edmonds (1972), based on the envelope threshold of his figs 1-6; intensityandtime generally defined as in (i). (iii) Data collected by Wells (1974); the intensityis the peakpulse intensity, and the time is the integrated pulse time.

familiar with the relief which locally applied heat-for example, from a hotwater bottle-can bring to aches and pains. According t o Lehmann and Guy (1972)) the t'emperature range which is effective is 40-45 "C. The frequency used in physiotherapy is usually within the range 1-3 MHz, and intensit'ies of between 0.25 and 3 W cm-2 are commonly applied from a transducerwith anarea of about 5 cm2. Theapplicator is either held in contactwiththeskin(with some oil or jelly as a couplant), or boththe applicator and the affected part are immersed in water, or a water-filled bag is placed between the applicator and the skin. The irradiation usuallyinvolves acontinuousmovement of the ultrasonicbeamoverthetreatedsurface. Each application may lastfor 3-50 min, and typically this is repeated ten times during a course of treatment. According to Patrick (1971), acute back pain, tenosynovitis, scar tissue and acute shoulder pain, all respond to ultrasonic therapy; really chronic conditions generally do not. The sooner that therapy is begun following trauma causing soft-tissue lesions, the better the result. 9.2. Neurosurgery

The first irradiations of nerve tissue were done with unfocused beams (Lynn and Putnam 1944, Lindstrom 1954). The literature is none too clear about the reasons and the results. The most important potential value of ultrasonic neurosurgery is that it is the only method (apart from some forms of radiotherapy) by which damage can be induced in a deep part of the nervous system without injury to the 2'7

P.N . T.Wells

662

overlying tissues. Conditions in a focused ultrasonic beam can be arranged so that it is only at the focus that the intensity is sufficient to cause damage. A typical arrangement is illustrated in fig. 17. The threshold relationshipbetween

R-"

E l e c t r i c a l connection

Metal case Transducer

iP

F o c u s e d beam f l o w out region

Fig. 17. Ultrasonic signal commissurotomy. The transducer is a section of a hollow sphere, and theultrasonic beam is contained within the metal cone through which saline is flowing. The focal region falls in the grey matter of the spinal cord (shown here in transverse section, having been exposed by laminectomy), and the application of a pulse of about 2 S duration with a peak space-time intensity of about 200 W cm-2 causes an elliptical lesion to develop, with a length of about 2 mm and a diameter of about 1 mm. Approximately 40 lesions are positioned close together along the axis of the cord, to interrupt the y, 6 and C fibres, thus controlling lower abdominal and pelvic pain without loss of motor or sensory function.

intensity and exposure time is shown in fig. 15. The size of the lesion is proportional to thelogarithm of the irradiation timefor lesions of less than 2.5 mm diameterandintensities of 0-2-2.5 kW cm-2 (Warwick andPond 1968). Typically, the lesion becomes visible 4-9 min after irradiation, increases in size until 15 min after irradiation andremains constant in size for 18 d ; it then diminishes as part of a recovery process. I n summary, focused ultrasonic neurosurgery has the following characteristics (Meyers, Fry, Fry,Eggleton and Schultz 1960) : (i) Permanent lesions of any size, shape and orientation can be produced a t any site; no disruption of intervening tissue occurs. (ii) Changes in function can be achieved, either temporarily or permanently. (iii) The vascular system can be left intact and functional in cerebral regions in which all neuronal elements are destroyed. (iv) White matter can be selectively damaged without involving grey matter. (v) Mortality and morbidity instances are low. The method has been used to control some of the symptoms of Parkinson's disease (Meyers, Fry, Fry, Dreyer, Schultz and Noyes 1959) and to destroy the pituitary gland (Hickey, Fry, Meyers, Fry, Bradbury and Eggleton 1963).

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Nowadays, better treatments are available. Apart from research into neuroanatomy (Fry 1972), it is only in spinal commissurotomy for the control of intractablepainthatthemethod is a t present being developed (Griffith, Brownell, Bowes, Halliwell and Wells 1973). 9.3. Vestibular surgery

Direct ultrasonic irradiationof the lateralsemicircular via asurgical approach through a thin layer of bone is quite effective in treating MBniBre’s disease and 1955, James 1963). An alternativeand relatedvestibulardisorders(Arslan rather simpler operation is based on irradiationthrougharound window (Kossoff, Wadsworth and Dudley 1967). Theresults are compatiblewitha mechanism whereby the heat due to the absorption of ultrasound causes a permanent shunt between the endolymphatic and perilymphatic systems by localized destruction of the function of the membranous labyrinth (Wells 1970). The chief hazard of the operation is that the facialnerve may be paralysed, but thiscan be avoidedby careful control of the ultrasonic intensity to ensure that the temperature does not rise excessively. 9.4. Treatment of juvenile laryngeal papillomatosis

Direct ultrasonic irradiation of the remnants of juvenile laryngeal papillomata following surgical excision reduces the recurrence rate (White, Halliwell and Fairman 1974). A likely explanation is that the virus (which is believed to cause the disease) is destroyed by the resultant heat. 9.5. Surgical procedures using direct mechanical effects Low-frequency ultrasound (about 25 kHz) is used to drive fine-pointed tools for dental applications (Balamuth 1967), to emulsify cataracts (Kelman 1967) and in bone surgery (Volbov and Shepeleva 1974).

Conclusions Ultrasonic methods are used routinely in the investigation and treatment of many pat’ients. The surgical applications of ultrasound are rather specialized, and, although widely used, neither the mechanisms nor the effectiveness of ultrasonicphysiotherapyhaveyet been thoroughlystudied.There is scope for much research into ultrasonic biophysics. The development of ultrasonic diagnostic methods has been a steady process with no individual dramatic breakthroughs. X-ray computerized tomography, on the other hand, is recognized as having revolutionized the investigation of the brain, and the method has become widely accepted in a very few years. The application of OT to abdominal studies has eclipsed ultrasound scanning, but, like any other eclipse, this situation is probably transient. Apart from the limitation due to gas inthe abdomen, ultrasonic scanninghas several important advantages over O T : it is faster; it has better resolution;it has better contrast between solid tissues and liquids such as blood; it is apparently safe; and it is much cheaper. 10.

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Probably the most importantfactor which presentlylimits therate of introduction of ultrasonicdiagnosticmethods into hightechnologyhealth care systems is the lack of doctors and technicians trained in interpreting and obtaining data. Hopefully, during the next few years training courses will be made widely available. It is also importantthat scientific and industrial resources should be devoted now tothe development of new ultrasonic instruments designed to make it easier to obtain still more clinically valuable information, so that in the longer term the benefits of ultrasonic diagnosis can be given to more people by fewer and less skilled medical and paramedical staff.

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CARSTESSES, E. L., and SCHWAN, H. 1959, P., J . Acoust. Soc. Am., 31, 305-311. CHAPMAS, I. V., 1974, BT.J . Radiol., 47, 411-415. CHESLER, E., JOFFE, H. S., BECK,W . , and SCHRIRE,V., 1971, Pediat. Clin. S . Am., 18, 1163-1190. CHIVERS,R. C., and HILL,C. R., 1975, Ultrasound X e d . Biol., 2, 25-29. CLARKE, P. R., and HILL,C. R., 1969, Expl Cell. Res., 58, 443-444. CLARKE, P. R., and HILL,C. R., 1970, J . Acoust. Soc. Am., 47, 649-653. CLARKE, P. R., HILL,C. R., and ADAMS, K., 1970, B r . J . Radiol., 43, 97-99. COGHLAN,B., and TAYLOR, M. G., 1976, Ultrasound Med. Biol., 2, 181-188. COGHLAN, B. A., TAYLOR, M. G., and KING, D. H., 1974, in Cardiovascular Applications of Ultrasound, Ed. R. S. Reneman (Amsterdam: North-Holland) pp. 55-65. COLOMBATI, S.,and PETRALIA, S., 1950, Ricerca Scient., 20, 71-78. CRAVEN,J. D., CONSTANTINI, M. A., GREENFIELD,M. A., and STERN,R., 1973, Investve Radiol., 8, 72-77.

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DANCKWERTS, H.-J., 1974, J . Acoust. Soc. Am., 55, 1098-1099. DAVIDSON,J. K., MORLEY,P,, HURLEY,G. D., and HOLFORD, N. G. H., 1975, Br. J . Radiol., 48, 435-450. DAWES,G. S., 1974, New Eng. J . Med., 290, 557-559. DAY,T.K., FISH,P. J., and KAKKAR,V. V., 1976, B r . M e d . J., 1, 618-620. DONALD,I., 1974, Am. J . Obstet. Gynec., 118, 299-309. DUNN,F., 1962, J . Acoust. Soc. Am., 34, 1545-1547. Am., 56, 1638-1639. DUNN,F., 1974, J . Acomt. SOC. DUNS, F., EDMONDS, P. D., and FRY,W. J., 1969, in Biological Engineering, Ed. H. P. Schwan (New York: McGraw-Hill) pp. 205-332. DUP~N, F., and FRY,W. J., 1961, Phys. Med. Biol., 5, 401-410. DCSSIK,K. T.,DUSSIK,F., and WYT, L., 1947, U’ien. Med. Wschr., 97, 425-429. J. B., JOSEPH, J., and WARWICK, R., 1968, Clin. Sci., 35, 273-285. DYSON,M., POND, DYSON,M., WOODWARD, B., and POND,J. B., 1971, Nature, Lond., 232, 572-573. EDLER,I., 1965, in UltrasonicEnergy, Ed. E. Kelly (Urbana:University of Illinois Press) pp. 303-321. EDLER, I., and HERTZ,C. H., 1954, K . Fysiogr. Sallsk. Lund Forh., 24, 40-58. EDMOKDS, P. D., 1972, in Interaction of Ultrasound and Biological Tissue, Ed. J. M. Reid and M. R. Sikov, (FDA) 73-8008 (Rockville: US Department of Health, Education and Welfare) pp. 299-317. EDMONDS, P. D., BAULD, T.J., DYRO,J. F., and HUSSEY,M., 1970, Biochim. Biophys. Acta, 200, 174-177. ELDER, S. A., 1959, J . Acoust. Soc. Am., 31, 54-64. ESCHE, R., 1952, Akust. Beih., 2, 71-74. EVANS,D. S., 1971, Ann. R. Coll. Surg. Eng., 49, 225-249. FEIGENBAUM, H., ZAKY, A., and WALDHAUSEN, J. A., 1967, Am. J . Cardiol., 19, 84-90. FEIZI, and EMANUEL, R.,1975, Br. Heart J., 37, 1286-1302. FIELDS, S., and DUNN,F., 1973, J . Acoust. Soc. Am., 54, 809-812. FISH, P,, 1972, in Blood Plow Measurement, Ed. C. Roberts (London: Sector) pp. 29-32. FITZGERALD, D. E., GOSLISG,R . G., and WOODCOCK, J. P,, 1971, Lancet, 1, 66-67. FLAX, S. W., WEBSTER,J. G., and UPDIKE, S. J., 1971, Instrum. Soc. Am. Trans., 10, 1-20. FLORIANI, L. P,, DEBEVOISE, N. T., and HYATT,G. W., 1967, Surg. Forum, 18, 468-470. FRANKLIN, D. L.,BAKER,D. W., ELLIS,R. M., and RUSHMER, R. F., 1959, IRE Trans. Med. Electron., ME-6, 204-206. FRUCHT, A. H., 1953, 2. Ges. E x p . X e d . , 120, 526-557. FRY,F. J., 1972, in Interaction of Ultrasound and Biological Tissue, Ed. J. M. Reid and M. R. Sikov, (FDA) 73-8008 (Rockville: US Department of Health, Education and Welfare) p. 297. FRY,F. J., KOSSOFF,G., EGGLETON, R. C., and DUNK,F., 1970, J . Acoust. Soc. Am., 48, 1413-1417. FRY, W.J., and DUNN,F., 1962, in Physical Techniques in Biological Research, Ed. W. L. Nastuk (New York: Academic Press) vol. IV, pp. 261-394. FRY,W. J., and FRY,R. B., 1954a, J . Acoust. Soc. Am., 26, 294-310. FRY,W. J., andFRY,R. B., 1954b, J . Acoust. Soc. Am., 26, 311-31 7. FRY,W. J., TUCKER,D., FRY,F. J., and WULFF,V. J., 1951, J . Acoust. Soc. Am., 23, 364-368. FRY,W. J., WULFF,V. J., TUCKER, D., and FRY,F. J., 1950, J . Acoust. Soc. Am., 22, 867-876. GAERTNER, W., 1954, J . Acoust. Soc. Am., 26, 977-980. S,, 1968, J . Acoust.Soc. Am., 44, GIGLIO,E. J., LUDLAM, W.M,, and WITTENBERG, 1359-1364. GOLDMAN, D. E., and HTJETER, T. F., 1956, J . Acoust. Soc. Am., 28, 35-37. GOLDMAN, D. E., and HTJETER,T.F., 1957, J . Acoust. Soc. Am., 29, 655. GOLDMAN, D. E., and RICHARDS, J. R., 1954, J . Acoust. Soc. Am., 26, 981-983. GOUTD, R . K., and COAKLEY,W. T., 1974, in Finite-amplitude Wave Effects in Fluids, Ed. L. Bromo (Guildford: I.P.C. Science and Technology Press) pp. 252-257. GRAMBERG,H., 1956, Dissertation, Johann-Wolfgang-GoetheUniversity, Frankfurtam-Main. GRAMIAK,R.,and WAAG,R. C., Eds, 1975, CardiacUltrasound (SaintLouis: C. V. Mosby).

o.,

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