Resolution enhancement in scanning of tissue

Fig. 4. Continuous

record of blood volume and

blood pressure after injeation of norepinephrine. Scale of minutes begins at an arbitary zero

2

0

. -

i

Time

6

.-... [min]

--

8

.“_.._.“___,._ .“___.. _“___ 10

REFERENCES I. PAGE, I. H., “Vascular mechanics of terminal Clinical shock”. Cleveland Quarterly, 13, 1 (1946). 2. BRILL? N. R. and SHOEMAKER, W. C.,

“Studies on the hepactic and visceral microcirculation during shock and after epinephrine administration. A preliminary report,” Surgical Forum, II, 119 (1960).

3. FELL? C. and RUSHMER, R. F., “Anatomic distribution of induced changes in blood volume, evaluated by regional weighing,” Journal of Applied Physiology, 16, 85 (1961). 4. JOHNSON. P. C. and SELKURT. E. E., “Intestinal weight changes in hemorrhagic shock,” AmericanJourt~alofPhysiology, 193,

reference probe was filled with fat from the omentum, free of blood. The difference in transit time is then given by

AT = ~~~~ rjat or

ATCollecting portional obtain :

(gb-

y)

- 6

..(3)

terms and solving for d (proto the blood volume) we

d = $&!$

AT-

~,dT......(4)

Equation 4 states that distance n is simply a constant multiplied by the measured difference in sonic transit time between the mixture and a reference probe. Temperature variations can be partially compensated for by placing the reference probe within the body, so that both probes are subjected to the same temperature variations. Alterations in blood hematocrit change ultrasonic

RESOLUTION by F. L. Thurstone*

velocity and necessitate monitoring. Frequency response was intentionally limited to 30 c/s by filtering. Both noise and drift of the measuring instrument are under 1 ns, which calculation shows to be good enough to allow 0’2 % changes in blood volume to be detected with a I .5 cm probe. A short delay line inserted in series with the reference probe makes it possible to calibrate the time difference in cubic centimetres of blood. Fig. 4 is a record of mesenteric blood volume with blood pressure measured simultaneously from the femoral artery. Small increases in blood volume occurred with systolic increase in blood pressure. When 3Opg of norepinephrine were quickly injected, the blood pressure increased while the blood volume decreased, which is the usual reaction to this drug. The apparatus has also been used to study pooling of blood in mesentery during haemorrhagic shock”. This work has been supported by Public Health Grants HE07158 and HE03998.

ENHANCEMENT and W.

A

technique has been developed which improves the lateral resolution of a pulseecho equipment by means of a highly focused transducer array which provides a very small beam at a focal point. The medical usefulness of two-dimensional images of tissue cross-section depends upon the resolution obtained in the image and upon the dynamic range of the presentation. In pulse-echo ultrasonic scanning systems, the resolution and the dynamic range of resolution are primarily determined by the lateral extent of the ultrasound beam produced by the transducer. An additional restric-

* Bowman Gray School of Medicine, Wake Forest College, Winston-Salem, North Carolina, U.S.A. This project was aided by a grant from The National Foundation

135 (1958). 5. MELLANDER, S., “Comparative

studies on the adrenerpic neurohormonal control of resistance aid capacitance blood vessels in the cat,” Arta Physiologica Scnndinavica, 50 (Supplement 176) : 1 (1960). 6. PINARDI, G., ADOLPH, R. .I. and RUSHMER, R. F., “Intestinal blood volume measurements in intact unanaesthetized dogs,” Physiologist, 4, 87 (1961). 7. REYNELL, P. C., MARKS, P. A.,, CHIDSEY, C.

and BRADLEY,S. E., “Changes m splanchnic blood volume and splanchnic blood flow in dogs after haemorrhage,” Clinical Science,

14, 407 (1955). 8. GIBSON, J. G., 11, SELIGMAN, A. M., PEACOCK, W. C., FINE, J., ALJB, J. C. and EVANS, R. D., “The circulating red cell and

plasma volume and the distribution of blood in large and minute vessels in experimental shock in dogs, measured by radioactive isotopes of iron and iodine,” Journal

of CIinicnl Investigation, 26, 126 (1947). 9. MULLINS, G. L. and GUNTHEROTH, W. G., “A pulsed ultrasonic transit time instrument,” Proceedings qf the 16th Annual Conferewe ou Engineering itI Medicine and 10 II

Biology,

146 (1963).

MULLINS, G. L., “Sonodistometer,” Digest 4th Intemational Conference on Medical Electronics. New York (1961). LINTERMANS, J. P., APPEL, A. J., BLOOM, R. S., MULLINS, G. L. and GUNTHEROTH, W. G., “Mesenteric flow and volume in

hemorrhagic cation.

IN SCANNING

shock.”

Submitted

for publi-

OF TISSUE

M. McKinney*

tion on range resolution is imposed by the minimum duration of the pulse. In the present system the receiver is time gated so that only echo information returning from the focal point is accepted: this information is used to generate a single point in the image display. A focusing system has been developed by the authors which produces a coherent focus of the sonic energy generated by a plane transducer. That is, energy from the transducer face is focused to a point in space and the length of the propagation path to the focal point is the same from every point on the transducer. In order to avoid internal reverberations which would lengthen the pulse, the focusing system incorporates onlyreflecting surfaces. Its numerical aperture is

C

large so that it will illuminate and receive echoes from irregular surfaces. A cross-section of the focusing system is given in Fig. 1. The basic reflecting surface is generated by the rotation of an ellipse about its major axis. Energy emanating from the first focal point of the ellipse is reflected from this spheroid surface to the second focal point of the ellipse. In order to illuminate the spheroid surface, a small paraboloid of revolution is placed at the first focus of the ellipse. The focal points of the spheroid and the paraboloid are made to coincide. A plane transducer is directed towards the paraboloid along the major axis of the spheroid. The paraboloid converts the plane wave from the transducer into a diverging spherical wave and the spheroid reflects this as a

uLTRAsoNrcs/January

1966

25

converging spherical wahe concerging on the target; the second focal point in space. Thus a coherent reflecting system is achieved with a large numerical aperture. Several prototype transducer arrays based on this focusing principle have been built, one of which is illustrated in Fig. 2. This array, machined from brass stock. has a spheroidal reflector of 25 cm diameter. The distance between its foci is 20 cm, yielding a numerical aperture of 0.53. The small paraboloidal reflector is supported at the first focal point of the spheroid by three struts. Initially the focusing system was aligned by replacing the transducer by a collimated light source: final adjustment used echo ranging techniques. The resolution of a scanning system based on a focused radiator and a timegated receiver is determined by the focusing characteristics of the array and by the duration of the pulses transmitted and gated on reception. The transducers used in testing our prototype scanning system have been heavily damped to provide a short pulse with a fundamental frequency of 2.25 MC/S. The shortest echo obtained from a plane surface is shown in Fig. 3. Thus the major portion of the echo energy is in a 1 ps period to the left of centre with an appreciable amount of ringing in a subsequent 1 PS period. On this basis, a 1 ps gate was chosen. This corresponds to an ideal target volume of 0.75 mm diameter in tissue or water. The ringing within the transducer limits the dynamic range of maximum resolution to approximately 20 dB in range and approximately 40 dB laterally. In order to generate a two-dimensional image, the focused transducer system is transported within the water bath in two dimensions by a modified radiographic scanning frame. This moves the array linearly at constant velocity back and forth, indexing in the second dimension at the completion of each sweep. The transducer position is transferred to the beam position on a storage oscilloscope by means of linear potentiometers. Echo information returning within the 1 ps time gate is used to intensify the beam, storing the image of echo from the focus point by point on the oscilloscope screen. A two-dimensional image is generated corresponding to echoes from a thin cross-sectional plane. Lateral resolution tests have been performed on this scanning system with a series of parallel wires lying on a plane as the target. The scanning system can resolve 0.5 mm diameter wires spaced I mm between centres with a dynamic range of approximately 40 dB. That is, the echo amplitude when the array is directed at a wire is 40 dB above the echo amplitude when it is directed at a point between the wires. A more difficult target and the scan obtained from it are

_

/.-

/

’

Elhpse

Fig. I. Cross-scc~~on of the focusing system. Outer reflector i% a degenerate ellipsoid, more preasely a spheroid

illustrated in Fig. 4. Here the wires are 0.25 mm in diameter and spaced at 0.5 mm. The scan shows these wires clearly resolved but the dynamic range of resolution was reduced to approximately 3 dB. The best possible resolution

Fig. 3. Echo pulse duration Fig. 4. Resolution

must therefore be only slightly bcttcl than this 0.5 mm spacing which corrcsponds to 0.75h at the fundamental. Velocity measurement is being de\eioped so that the velocity in the coupling medium can be accurately matched to

from a plane reflector

test model and an ultrasonic

scan obtained

therefrom.

Wires are 0.5 mm

anan

that of the tissue under study. The chief limitation on image resolution will be the changes of velocity within complex The dynamic range of the tissues. display will be limited by ringing of the transducer and reverberation within the tissue. The improvement in resolution over conventional B-scanning techniques pro-

vided by this highly focused C-scan method is at the expense of an increased time required to generate the scan. The immediate cause of delay is mechanical, the transport system for the reflecting array, but an ultimate limit is imposed by the maximum permissible pulse repetition rate and the number of points wanted on the display.

REFERENCES

KOSSOFF, G. er al., “Ultrasonic two-dimensional vtsualization techniques,” lnsrirute of Electrical nerd Electronics Engbreers Transactions on Sonics and Ulrmsortics, SU-12. 31 (1965). 2. HERTZ,C. H., “Ultrasonic heart investigation,” Medico1 Electronics nnd Biological Engirreering, 2, 39 ( 1964). 3. OLOFSSON, S., “An ultrasonic optical mirror system,” Acrrsfica, 13, 371 (1963). 1.

DOPPLER MEASUREMENT OF MITRAL VALVE AND VENTRICLE WALL VELOCITIES Y. Yoshitoshi*, Y. Kohashit.

K. Machii*, H. Sekiguchi”, S. Shimizu” and H. Kuno’.

H eart

sounds do not necessarily reflect the velocity of the opening and closing of the mitral valve and are easily affected by the physical properties of the valve, the transmission characteristics of the chest wall and various other factors. The present study was undertaken chiefly to measure the maximum velocity of mitral valve closure in normal subjects and in patients with mitral stenosis by the frequency analysis of the ultrasonic Doppler signal: maximum velocity indicates the movability of the mitral valve. At the same time, anterior ventricular wall movement in the apical area was measured. METHOD

Our ultrasonic Doppler apparatus used is similar to that reported by Satomura et aL1T2 A 3Mc/s ultrasonic wave is radiated to the heart through the chest wall at the third left interspace and over the apical area. Sound is reflected from the anterior ventricular wall, from the mitral valve and from the bloodstream. Because of the Doppler effect, the frequency of the reflected wave f,differs from the frequency of the original wave J The frequency of the Doppler signal fd, obtained by mixing f and f,, is proportional to the velocity of the cardiac movement in the direction of the ultrasound beam. The Doppler signals are recorded on magnetic tape, together with the frequency-modulated electrocardiogram, and transferred to a sound spectrograph, the output of which is a plot of frequency against time, usually known as a frequency spectrogram. In the present experiments the electric input to the transmitter of the transducer was about 2W. The piezoelectric element of the transducer was a lead zirconate disc IO mm in diameter and divided into two semicircular pieces, one for the transmitter and the other for the detector. *First Deuartment of Internal Medicine, Faculty of Medicine, University of Tokyo. TKobayashi Institute of Physical Research, Tokyo. ORion Co. Ltd., Tokyo.

MITRAL

VALVE

Y. Mishina*,

S. Ohta*,

SIGNALS

Mitral valve signals were usually obtained from the third left interspace. In the frequency spectrogram of the Doppler signals (Fig. 1) mitral valve movement appeared as sharp upward and downward deflections. In normal subjects, four signals, Mel, MO, MO’ and Mps, were recognized. According to the direction of the sound beam, MO’ and Mps could not be obtained. In mitral stenosis, only Mel and MO signals were recognized. The peak of Mel is synchronous with the sharp deflection of the first heart sound. The peak of MO appears 0.04-0.10 set after the onset of the second heart sound and is synchronous with the mitral opening snap in mitral stenosis. MO’ and Mps signals appear in diastole. In the present paper, only the Mel signal will be discussed. In Fig. 1 the peak frequency of the Mel is approximately 1,200 c/s, corresponding to the maximal velocity of the valve closure of approximately 30 cm/s. In normal subjects, the maximal velocity of the mitral valve closure ranged from 20 cm/s to 40 cm/s (average, 27 cm/s). In mitral stenosis, the velocity ranged from 30 cm/s to 80 cm/s (average, 52 cm/s), twice faster than normal. In a patient undergoing commissurotomy, the velocity was reduced from 65 cm/s to 30 cm/s after the operation (Fig. 2). In three cases of mitral stenosis with atria1 fibrillation, intensity of the first sound and velocity of the mitral closure fluctuated according to the nature of the preceding diastole. If the preceding diastole was short, a great intensity with high velocity of the mitral closure was observed, while the reverse was true when the preceding diastole was long. The maximal velocity of the mitral opening ranged from 12 cm/s to 20 cm/s in normals, and from 25 cm/s to 60 cm/s in mitral stenosis. ANTERIOR

VENTRICULAR

WALL

SIGNALS

Anterior ventricular wall signals were obtained from the apical area. Instead of the sharp deflection of the valve

and Y. Hanaoka*.

movement, the anterior ventricular movement was reflected by two triangular peaks, protosystolic and prodiastolic (Fig. 3). The protosystolic peak coincided with the maximal ejection phase of the ventricle, and the protodiastolic wave with the ventricular rapid-filling phase. The maximal velocity of the anterior ventricular movement at the protosystolic ejection phase ranged 7 cm/s to 12 cm/s, and no marked difference was found among the normals, hypertensions and patients with aortic insufficiency. BLOODSTREAM SIGNALS

Bloodstream signals were detected along the outflow tract of the left ventricle. Blood movement appeared on the spectrogram as an amorphous shadow, lying between 50 c/s and 400 c/s. It looks like a white noise and is easily differentiated from valvular and ventricular signals (Fig. 1). DISCUSSION

In 1954 Herz and Edler,3 and in 1957 Effert4 recognized by ultrasound cardio-

Time

[WC]

Fig. 1. Frequency spectrogram of the Doppler signal taken at the third left interspace. Mitral valve and bloodstream signals MCI Mitral closure MO Mitral opening BS Bloodstream

ULTRASONICS/~a?IUary

1966

27