Pressure measuring system with ultrasonic wave

A system for measuring from the outside of a living body the pressure within the heart or the pressure of any portion which does not allow a measurement by the direct insertion of a pressure measuring sensor. This system provides a method of measuring the pressure of the object by generating fine bubbles through cavitation, applying a low frequency ultrasonic wave to the medium, and then detecting the bubbles which are generated with a system for detecting the high or low frequency harmonics due to the bubbles or a higher frequency ultrasonic wave applied to the medium.

BACKGROUND OF THE INVENTION 
The present invention relates to a liquid pressure measuring system. 
Particularly, to a system for non-destructively measuring fluid pressure 
from outside the body containing the fluid, by the use of ultrasonic 
waves. For example, blood pressure in the heart of a living body or the 
pressure of flowing liquid used in the chemical industry to show 
particularly high temperature and pressure, the high probability of a 
chemical reaction or the existence of solid particles or fibres etc., in 
cases where it is difficult to insert a pressure gauge directly into a 
measuring object. 
Currently, a catheter equipped with a pressure sensor has been inserted 
into the blood vessels or heart in order to measure blood pressure. This 
method is accompanied by the disadvantages of creating pain and a risk to 
life by the unexpected misoperation or by infection. The methods of 
acoustically detecting a blockage of the blood flow in the arm or 
detecting the start of a pulse by winding a cuff around the arm and 
changing the air pressure are also known. This method using a cuff can be 
applied to the arms and legs but cannot be used for measurement of 
internal organs such as the heart. 
In the field of industrial systems operating at a high temperature, a low 
temperature, and/or in a strong radiation field, and the handling of 
liquids which are chemically active, of high viscosity, or of a dense 
mixture of grain particles, chips of wood and fibres etc. pressure sensors 
have often been damaged by temperature, radiation, chemical reaction or by 
external force due to solid materials. Therefore, it is desired to measure 
pressure from outside of a vessel or system but there has been no adequate 
method. 
SUMMARY OF THE INVENTION 
An object of the present invention is to provide a method of 
non-destructively measuring the pressure of a desired region within a 
substance, from outside, by the following process. Ultrasonic waves are 
applied to the desired region within a substance, to generate bubbles 
within the liquid existing in the region during the negative pressure 
cycle of the ultrasonic waves and, thereafter, the generation of bubbles 
is detected by harmonic or subharmonic ultrasonic waves which accompany 
such bubbles and/or by the echo of other ultrasonic waves of higher 
frequency applied to the region. 
Generally, ultrasonic waves are sound waves of a frequency higher than the 
audible frequency (16 kHz) but the present invention is meant to include 
audible sound waves and ultrasonic waves in the sense mentioned above. 
The present invention utilizes cavitation, a method by which the gas and/or 
water content of the blood, lymph and cell liquids etc. existing in the 
heart, blood vessles or organs of a human body are isolated or vaporized 
by the negative pressure of the externally applied ultrasonic waves to 
generate micro bubble nuclei which then grow in to larger bubbles.

DESCRIPTION OF THE PREFERRED EMBODIMENT 
The critical pressure of bubble formation is a function of the ambient 
pressure (about 1 atm at sea level), the temperature, the frequency of 
ultrasonic waves applied and the type of waves, for example progressive or 
standing waves. It is also required that the liquid to be measured has 
been sufficiently degassed or on the contrary has been exposed to or 
absorbed the gas sufficiently. 
As an example, FIG. 1 shows the measurement of the critical amplitude of 
ultrasonic waves used for generating bubbles under an ambient pressure of 
1 atm at room temperature. FIG. 1(a) relates to sufficiently degassed 
water and FIG. 1(b) relates to sufficiently airated water. The horizontal 
axis indicates the frequency, while the longitudinal axis indicates the 
sound pressure (amplitude). The profile changes in the frequency zone of 
about 10.sup.4 to 10.sup.5 Hz. At frequencies under 10.sup.4 Hz, the 
critical pressure for generating bubbles does not depend on the frequency, 
but it is highly dependent on the frequency at 10.sup.5 Hz or higher. This 
indicates that a time of about 10.sup.-4 seconds is necessary for 
formation of nuclei and growth of bubbles. 
In past methodologies, it has been impossible to accurately measure the 
critical pressure for formation of bubble nuclei. It has been measured by 
optically recognizing the bubbles or by the accoustic sound generated when 
the bubbles grow and break. These methods result in a time lag during 
growth of the bubbles between the application of pressure and the 
detection of bubbles, and therefore, fluctuation of critical pressure 
measurement is induced and some delay in the response time is caused. In 
the case of the present invention, bubbles are detected in the early stage 
of growth as bubble nuclei thereby resulting in an improvement of 
measuring accuracy and a decrease in the response time. 
The bubble nuclei differ from the liquid in accoustic impedance and give an 
intensive reflection and scattering of ultrasonic waves. The bubble nuclei 
are generally equal to or smaller than the wavelength of the applied 
waves, in the frequency range from 1M to 10 MHz, and generate a Rayleigh 
scattering. The wave energy of ultrasonic waves is proportional to the 
square of frequency. Therefore, the higher the frequency is, the higher 
the sensitivity is. But, the ultrasonic waves are exponentially attenuated 
as they are transmitted into a living body and the attenuation coefficient 
of a living body is almost proportional to the frequency. If the frequency 
is high, the ultrasonic waves suffer the same attenuation during forward 
and backward transmission for the detection of nuclei bubbles at the 
deeper region of the human body. Therefore, ultrasonic waves of 1 MHz to 
10 MHz are suitable for detection of bubbles within a human body. 
Reflected waves can be easily distinguished, in the case of measuring blood 
pressure, within the heart or within large blood vessels because the waves 
reflected from the blood are weak but intensively reflected waves appear 
due to the bubble generation. It is, however difficult to distinguish the 
bubble generation because intensive waves reflected from the structural 
tissues coexist with the bubbles such as the regions near blood vessel 
walls, small blood vessels, lymph vessel and tissue fluid. Even in such a 
case, if the liquid is flowing, the reflected waves have a Doppler shift 
due to the flow of the bubble nuclei. An embodiment of the present 
invention provides a system of eliminating the effect of reflections from 
the structual organs by extracting such Doppler shifts a method of 
detecting bubble generation with a Doppler signal and also, as a result, a 
highly sensitive measurement. This system simultaneously measures both 
flow rate and pressure, resulting in the attainment of highly accurate and 
detailed data. 
As explained above, measurement of the critical pressure for bubble 
formation according to the present invention becomes more accurate and the 
time delay, when an applied pressure is swept, is almost eliminated. 
Therefore the sweep speed can be increased while still detecting the 
generation of bubble nuclei with a high sensitivity. 
A method of sweeping the applied pressure can be selected freely, but it is 
easy to utilize a pressure amplitude which changes in relation to a sine 
wave. For this method, a continuous wave or a burst of waves illustrates 
in FIGS. 2A, B, and FIG. 2 or a pulse wave, illustrated in FIG. 2C, can be 
used. In this case, wider band widths are necessary in the sequence of A, 
B and C around the center frequency f. Moreover, it is also possible to 
perform the sweep using a saw-tooth wave as illustrated in FIG. 2D. 
It is desirable to set the center frequency f, from FIG. 1, at 10 kHz or 
less because the bubble nuclei generating pressure can thereby be lowered. 
The high sensitivity nuclei bubble detection of the present invention 
allows the use of higher frequency ultrasonic waves for nuclei bubble 
generation but it is desirable to be able to ignore the attenuation within 
a human body and consequently it is desirable to select a frequency of 
1000 kHz or less. Attenuation by tissue is of the order of 1 db/MHz.cm 
and, therefore, an attenuation of 0.2 db is expected when ultrasonic wave 
of 10 kHz are irradiating an area located 20 cm from the surface of a 
human body and such attenuation can almost be ignored. 
In the following explanation, blood is used as the measuring liquid. The 
blood pressure in the heart rapidly changes and this change is denoted as 
P.sub.p (t) for each pulsation Reference is made to the ambient pressure 
Pa (generally atmospheric pressure) and the current object is to measure 
P.sub.p (t) non-invasively, i.e., from the outside of body. Generation of 
nuclei bubbles in blood by the cavitation can be thought of as depending 
on the absolute pressure when the rate of gasification, such as 
dissolution and absorption of gas is constant and the temperature is 
constant. Such a critical pressure is denoted as Pc. The absolute pressure 
of blood is denoted as P(t), which is always changing due to pulsation. 
The following relation can thus be obtained. 
EQU P(t)=P.sub.p (t)+Pa 
Where, Pa is changing very little over several hours and P.sub.p (t) 
indicates the change by pulsation which is slower, on the order of msec, 
for each pulse. Here it is supposed that a sound field denoted as Q(t) is 
applied to the measuring area, which is given by the next equation. 
##EQU1## 
Where, the frequency f is selected, for example, about 10 kHz so that Q(t) 
changes at a faster rate than P.sub.p (t). 
The combined absolute pressure (t) is denoted as follows. 
##EQU2## 
When Q.sub.0 is selected properly in order to cause (t) to decrease to 
P.sub.c (critical pressure) or less in the negative cycle of Q(t), nuclei 
bubbles are generated in the range where (t).ltoreq.P.sub.c and at the 
time t.sub.c, when (t)=P.sub.c, the following relations can be obtained. 
##EQU3## 
Accordingly, P.sub.p (t.sub.c) can be obtained when the values of Q.sub.0, 
t.sub.c, P.sub.a, P.sub.c are known. Here, Q.sub.0, t.sub.c, P.sub.a can 
be measured and P.sub.c can be obtained by calibrating the result of a 
measuring method as explained later. 
The critical pressure P.sub.c can be obtained previously in the case of 
industrial systems, but it may change when chemical reactions continue 
during a step by step process or as in the human body, the temperature and 
the degree of gasification may change over time due to changing 
conditions, for example, exercise, sleeping, etc. or between individuals. 
Explained below is a method for determining P.sub.c of the blood of a 
human body, as an example. 
In the case of blood flow within a human body, which circulates in the 
closed loop, blood having the same temperature and rate of gasification as 
the main measuring area can be found in other areas. For example, the 
blood in the left ventricle of the heart can be though of as flowing into 
the arteries of the upper arm and the blood in the veins of the upper arm 
can be assumed to have the same temperature and rate of gasification as 
that of the right ventricle. However, the blood exchanges substances at 
the vasal capilaries of the lungs and tissues and therefore its 
characteristics change as it passes through them. For this reason, when 
the critical pressures of the arteries and viens of the upper arms can be 
obtained, the pressures in the left atrium and the left ventricle of the 
heart (artery blood) and the right atrium and the right ventricle (vein 
blood) can be measured. The artery pressure P.sub.p ' of the upper arm is 
often measured by the following procedures. A rubber air tube is wound 
aroud the arm to temporarily block the flow of blood by increasing the air 
pressure in the tube. The pulsating condition is then monitored with an 
acoustic receiver while the air pressure is gradually decreased and the 
maximum blood pressure P.sub.p ' max is measured by accoustically 
detecting the sound generated when the peak pulse flow returns. The 
minimum blood pressure, P.sub.p ' min, is found by detecting the sound 
generated when the lower limit of the pulse flows returns. Generally, the 
maximum blood pressure can be measured with higher accuracy and it is 
desirable to calibrate the critical pressure with this value. 
The vein pressure and tissue liquid pressure can be measured directly by 
inserting a pressure sensor into the blood vessel or tissue; this is much 
safer than insertion into an artery. 
First, the P.sub.p ' max is measured at the upper arm portion and 
thereafter the same arm portion is measured under the same condition by 
the method of the present invention. Thus, a value of P.sub.c can be 
obtained by assuming that the result of the measurement is equal to 
P.sub.p '. In the case of industrial systems, the value of P.sub.p for the 
measuring can be determined by calibrating P.sub.c at a more feasible and 
safer region. 
Moreover, the necessary applied pressure Q.sub.0 can be lowered, thereby 
improving the response time by previously dissolving or absorbing harmless 
gas which easily forms bubbles, such as a noble gas like helium or carbon 
dioxide into the blood. These gases can be dissolved into the blood during 
breathing by placing the body in an atmosphere where part of the nitrogen 
is replaced by helium or carbon dioxide or by increasing the pressure of 
the ambient mixture. It is also possible to introduce gas directly into 
the blood vessels by injecting well gasified or volatile liquid. As 
explained above, measurements can be made with a small Q.sub.0 if the 
critical pressure P.sub.c is increased. Moreover, the response time is 
improved and the number of repetitions of the measuring sweep can be 
increased. Reduction of Q.sub.0 not only makes easy and economical the 
designing of unit, but also minimizes the effects of the ultrasonic waves 
on a human body. 
The desired timing and area of measurement can be obtained, and noise and 
unwanted signals can be eliminated, by synchronizing the application time 
and the area and measuring time and area so that the ultrasonic waves for 
detecting bubbles are applied to the measuring area only when the negative 
sweep of pressure is interacting with the measuring area. 
An embodiment applied to a human body is explained hereunder. 
FIG. 3 and FIG. 4 show the methods of forming the ultrasonic wave amplitude 
which is applied to areas which do not prevent transmission of the 
ultrasonic waves, such as the abdomen and arms or legs. FIG. 5 shows the 
method which can be applied to the case where the ultrasonic waves cannot 
pass through the body because a lung, containing air, exists behind the 
heart. The lung becomes an intensive reflector due to the large difference 
in the accoustic impedance between the air and tissue, and therefore, the 
ultrasonic waves can not pass through the body. 
The transducer 1 forming the sweep pressure is driven, for example, at a 
center frequency of 10 kHz. If the ultrasonic waves are applied to a human 
body, a transducer diameter of 50 to 200 mm is most desirable. A hole of 
about 15 to 25 mm in diameter is provided at the center in order to mount 
the bubble detection send/receive transducer 5. A human body 2 and a 
particular tissue 3 such as the heart, liver or artery surround the 
measuring area 4. The dimension of the measuring area 4 is determined by 
the beam diameter of the bubble detection ultrasonic waves and the drive 
pulse length or the gate width for extracting the measuring signal from 
the reflected signal with the timing gate. The the bubble detection 
send/receive transducer 5 has a center frequency of, for example, 3.5 MHz. 
The diameter required for obtaining a sufficiently converged beam is about 
10 mm. For humans, the size of the measuring area 4 can be set to several 
millimeters. 
FIG. 3A shows the example where 1, 5, and 2 are respectively arranged 
within water 90 which is a sound conductive medium. A reflector 6 
consisting of a metal plate, having an acoustic impedance which is greatly 
different from that of water or a living body is arranged within the water 
opposite to the transducer 1 with the distance of n.times.half-wavelength 
(.lambda./2). Simultaneously the transducer 1 transmits a continuous wave 
of wavelength .lambda.. Thereby creating a resonant condition between the 
transducer 1 and the metal plate 6, thus forming a standing wave as shown 
in FIG. 3B. 
When the frequency is 10 kHz, the wavelength in the water or a living body 
is 15 cm. If, for example, n is 4 the distance between the transducer 1 
and the metal plate 6 becomes 30 cm (15/2.times.4). This is sufficient for 
placing the abdomen of a human body between the transducer 1 and the plate 
6. The vibration loop center of the standing wave can be set to the 
measureing area by shifting the transducer 1 and the plate 6 in relation 
to the abdomen 2 while keeping the distance between the transducer 1 and 
the plate 6 at a constant value. As the pressure in the area changes with 
the sine wave of 10 kHz, the maximum pressure amplitude of which is the 
amplitude of loop, generation of nuclei bubbles can be obtained by using a 
half cycle of negative swing for the pressure sweep. 
FIG. 4 is an example of the use of a progressing wave. A plastic bag 7 
containing water is used in place of the water in FIG. 3. This plastic bag 
exists between the transducer 1 and the body 2 such as the abdomen of a 
human body. Jelly or oil is applied at the contact surface in order to 
obtain transmission of the ultrasonic wave by eliminating the air. Another 
plastic bag 8 containing water, and a non-reflective absorber of 
ultrasonic waves 9 consisting of a plastic or rubber containing metallic 
powder or bubble corpuscles is placed on the other side of the body. is 
placed on the other side of the body 8 and the absorbant material 9 are 
integrated and the jelly or oil is applied at the contact surface between 
the body 2 and the water bag 8 in order to eliminate any air gaps. When 
the transducer 1 transmits the pulse wave as shown in FIG. 2C, the pulse 
wave progresses into the absorbant material 9 from the transducer 1 at the 
velocity of sound in water (about 1500 cm/sec) and is absorbed. FIG. 4B 
shows the of pressure value at a particular moment during its 
transmission. When looking at the particular measuring area, for example, 
the area 4, the pressure varies with the same waveform as the waveform 
transmitted by the transducer 1, but it is delayed by a time obtained by 
dividing the distance between the transducer 1 and the measuring area 4 by 
the velocity of sound and changes with time. Namely, a pressure sweep is 
carried out. 
FIG. 5 shows the case where a strong absorber or reflector, like a lung, 
exists behind the measuring area 4, such as the heart. In the case where a 
strong absorber exists behind the measuring area, the situation is similar 
to the case of FIG. 4. But if there is an strong reflecting surface 10 
behind the measuring area, the sweep pressure at the measuring area 4 
becomes uncertain because the pressure field is formed at the area 4 by 
both the reflected waves from the reflecting surface 10 and the field of 
progressing wave from the transducer 1. In order to prevent such 
overlapping, the width of the progressive wave must be shortened by 
setting its center frequency at 100.about.1000 kHz and the wave 
propagation direction (incident direction) must also be changed as shown 
in FIG. 7. The existence and location of a reflecting body can be detected 
by using the transducer 1 as the receiver or by the transducer 5. FIG. 5B 
shows the pressure distribution at a particular moment. 
In any case, the size of the transducer 1 (diameter) cannot be too large 
for practical use and therefore the diameter becomes almost equal to the 
wavelength and, as a result, the wave generated becomes similar to a 
spherical wave. 
In FIG. 3, the transducer 1 must supply a drive energy large enough to 
compensate for the energy spherically diverging in directions other than 
towards the plate 6 in order to obtain a resonance between the transducer 
1 and the plate 6. In FIG. 4, and FIG. 5, the transmitter surface of 1 is 
not required to be flat and can be formed as a concave surface in order to 
converge the energy in the required direction. In any case, the pressure 
amplitude along the axis perpendicular to the face of the transducer 1 
changes as a function of distance z and therefore it is necessary to 
determine the function, by setting the transducer 1 in water without a 
human body 2 and measuring the pressure as the function of the axial 
distance z. 
The bubble detection transducer 5 can be flat or concave and can also be a 
phased array type of multielements. As the material, structure, circuit 
etc., those used by the so-called A mode, M mode, B mode and Doppler 
measurement can be used. 
As shown in FIG. 7, the measuring area can be determined while observing 
the sectional view of the B mode. For this purpose, the method similar to 
the well known Doppler measurement combining the B mode can be used. 
In this case, since the applied pressure sweep frequency is sufficiently 
low, simultaneous operation with the bubble detecting system can be 
realized without any interference between them. On the contrary, the B 
mode and bubble detection frequencies are sufficiently high, so any 
influence or interference on the low frequency critical pressure does not 
occur. This can be understood from FIG. 1. 
The transducer 5 can be mounted in a port of transducer 1 as indicated in 
the figure, or it can be mounted in a location other than the port of that 
transducer 1. If the B mode is used in combination the sector scanning can 
be done by 5 itself and the scanning for detecting bubbles passing the 
measuring area 4 may be done during the scanning. Moreover, it is also 
possible to use another B mode probe as shown in FIG. 7. 
If the measuring area 4 is located in the tissue cell, the Doppler effect 
cannot be used because the measuring area 4 will have no blood flow and 
detection must be made by extracting the change of the reflection 
intensity. For example, when a burst of waves (having a duration of about 
1 .mu.s) with a center frequency of 3.5 MHz is transmitted from the 
transducer 5, the pulse becomes a burst wave of about 1.5 mm in length and 
progresses at a rate of about 1.5 mm/.mu.s. As the pulse progresses, 
reflected waves are sent back from each point in accordance with changes 
in the accoustic impedance. Therefore, the received waveform at transducer 
5 is continuous and complicated. But, observations can be made by only 
receiving the waveform reflected from the position of the measuring area 4 
by extracting that signal with a timing gate. This is ordinarily well 
known. The one reflected signal from the measuring area 4 can be obtained 
for a single scanning as explained above. Namely, when the measuring area 
4 is located at the depth of about 20 cm within a body, the time required 
for the forward and backward transmission of ultrasonic wave is 266 .mu.s 
and measurements can be made 3760 times per second. As explained above, 
when the measuring area 4 is located in the flow of the heart or a blood 
vessel not only the simple reflection intensity but also the Doppler shift 
caused by blood flow can be analyzed and detected by a well known method 
such as the doppler method which is very effective for eliminating waves 
reflected from structural tissues. 
If the applied pressure has a center frequency of 10 kHz and is swept to 
the negative direction sinusoidaly, the negative half cycle is about 50 
.mu.s. Therefore, a single detection can be obtained in a single sweep. 
The critical pressure can be detected by several sweeps and detections 
where the phases of the sending and receiving waves of the ultrasonic 
waves for both pressure sweep and detection are shifted a small amount 
each time. Details of these procedures are shown in FIG. 6. The heart 
pulsation is 1 to 2 times per second and therefore, it has a sufficient 
frequency for following dynamic changes of pressure. In order to detect, 
in detail, the status wherein the heart pressure rapidly changes, 
measurement can be made by shifting the phase so that the measuring points 
are sequentially placed in a rapidly changing period by synchronization 
with an electrocardiograph signal. 
FIG. 6A shows how the sweep of the pressure is formed at the measuring area 
4. The vertical axis indicates the absolute pressure (t), which is given 
as a sum of the atmospheric pressure P.sub.a, the heart pressure P.sub.p 
(t) with reference to the atmospheric pressure and applied sweep pressure 
-Q.sub.0.sin (2.pi.ft). In this figure, P.sub.c is the bubbleforming 
critical pressure. The horizontal axis indicates the time t. When (t) is 
lower than the P.sub.c, the bubbles are generated. 
FIG. 6B shows the reflected signal extracted by the timing gate. The 
vertical axis indicates the amplitude, while the horizontal axis, the 
time. T is the time when the ultrasonic wave pulse is transmitted. M is 
the time when the send pulse reaches the measuring area 4. R is the time 
when the reflecting signal is received. If the distance between the 
transducer 5 and the measuring area 4 is l, the sound velocity is V, thus 
the time interval between T and R is given by 2 l/V and the following 
relation is obtained. T-M=M-R. Along axis 11 of FIG. 6B, the sending time 
T.sub.1 is synchronized to the applied pressure waveform so that the 
measuring time M.sub.1 coincides with the time T1 of the sweep pressure. 
The waveforms 12, 13 . . . are slso obtained in sequence, similarly 
shifting the time of T.sub.2, T.sub.3, . . . as shown in the figure. In 
the case of the waveforms 11 and 12 where (t) does not exceed P.sub.c, 
the reflected signals R.sub.1, R.sub.2 are low in amplitude, but when (t) 
exceeds P.sub.c, the waveforms 13, 14, and 15 give the intense reflected 
signals R.sub.3, R.sub.4 and R.sub.5 because generated bubbles have very 
different accoustic impedances. FIG. 6C shows the waveforms obtained by 
extracting only the Doppler shift signals from the reflection signals of 
FIG. 6B. The signals R.sub.1 ', R.sub.2 ' are sufficiently small as 
compared with the R.sub.3 ', R.sub.4 ', R.sub.5 ', and improve the bubble 
detection accuracy. 
In any case, when a pressure (-Q.sub.0.sin 2.pi.ft) at each point (t.sub.1, 
t.sub.2, . . . ) of the pressure sweep waveform is previously known, 
P.sub.c can be obtained from the point where bubble generation starts. 
In other way, P.sub.c can also be obtained from the minimum Q.sub.0 for 
detecting bubble generation which can be obtained by adjusting such 
Q.sub.0. In FIGS. 6B and C, the detecting waveforms 11, 12, . . . are 
overlapping in time in order to make clear the phase relation with the 
pressure sweep waveforms. In practice, the waveforms 11, 12, . . . are 
sent and received for different sweep cycles. 
As another method, the condition of measuring point 4 can be measured 
continuously by sending the bubble detection ultrasonic wave as the 
continuous wave in place of the pulse as shown in FIG. 6. This embodiment 
is explained below, upon reference to FIG. 7. 
In FIG. 7, 2 is a human body, 3 is the heart and 4 is the measuring area, 
selected in the blood flow in the left atrium in the figure. Numeral 1 is 
the ultrasonic transducer for pressure sweep and it is driven at a center 
frequency of 10 to 1000 kHz. The waveform is generated in the waveform 
generator 22. The waveform generator 22 digitally stores the waveforms 
which are converted from A to D in series and generates wave forms with a 
D/A conversion by sequentially reading the stored data. The center 
frequency can be changed by changing the period of the read clock. The 
amplifier 21 drives the transducer 1 through the power amplification of 
the waveform obtained from the waveform generator 22 and forms the 
necessary sweep negative pressure. Numeral 5 is the bubble detection 
probe. In this embodiment, the transmitting unit 5' and receiving unit 5" 
are provided individually and the continuous wave with M sequence 
modulation is transmitted and received. The base frequency generator 24 
utilizes a crystal oscillation unit and it operates, for example, at 2 
MHz. The M sequence modulation circuit, 25 sequentially reads the M 
sequence codes previously stored in the ROM in accordance with the clock 
from the timing control circuit 23 and, for example, phase-modulates the 
base sine signal. The power amplifier 26 drives the transmitter 5'. The 
receiving unit 5" receives signals which are amplified by the receiving 
amplifier circuit 27. The M sequence codes from the sequence modulation 
circuit 25 are sent to the multiplication circuit 30 via the variable 
delay circuit 28 which provides a delay equal to the traveling time of the 
sound forward and backward from transmitter to measuring area and back, by 
reffering the preset value specified with the ten-key in the depth setting 
circuit 29 and is compared with the output of the amplifier 27 by making a 
correlation between them. This is realized by the multiplication circuit 
30. An output of the multiplication circuit 30 is orthogonally detected by 
31 through comparison with the original oscillation signal. The real and 
imaginary part are sent to the amplitude circuit 32, where the square 
value of the amplitude is obtained by the integral circuit having a time 
constant shorter than the M sequence code length but longer than the code 
interval, for example, several tenths of a code length. The square-sum 
circuit, and the squared amplitude are used as the Y.sub.1 signal. 
The real or imaginary parts of the output of 31 are sent to the Doppler 
extraction circuit 33. This signal is detected after the band pass filter 
31 which functions to limit higher frequencies, allowing the Doppler shift 
frequency to pass but not allowing the original oscillation frequency to 
pass. The bandpass filter 31 also functions to limit lower frequencies by 
not allowing the lower Doppler shift frequency to pass due to their 
stationary or almost stationary speed. This signal is then sent to Y.sub.2 
as the Doppler signal. If it is necessary to judge the direction of blood 
flow, both real and imaginary part are used. 
The timing control circuit 23 generates the required clock signals from the 
original oscillation frequency of the base frequency generator 24, and 
also generates the control signals using the built-in program for each 
portion. 
On the other hand, the actual pressure waveforms at each position of the 
measuring area 4 within the water caused by the transducer 1 are 
previously measured. These are stored, after the A/D conversion, into the 
sweep waveform storing circuit 34. The digitized waveforms are selected 
from the waveform storage circuit 34 corresponding to the depth preset by 
the depth setting circuit 29. These are read by the clocks which control 
the read start timing and the read speed provided by the timing control 
circuit 23, and the data obtained is D/A converted by the D/A converter 35 
and is used as the X axis deflection signal (negative sweep pressure) for 
the CRT of the display unit 36. Numeral 36 is the two-channel 
synchroscope, giving respective outputs of the amplitude circuit 32 and 
the Doppler extraction circuit 33 to the Y.sub.1 and Y.sub.2 axis. From 
the Y.sub.1 -X and Y.sub.2 -X curves, the critical pressure P.sub.c can be 
confirmed, and the pressure P.sub.p can be obtained by reading the X 
values of the rising and falling points of the displayed curve. In this 
example, a manual judgment is made from the curve but is can be made 
automatically by the electronic means. Of course, P.sub.p can be digitally 
or analogously displayed and recorded continuously. 
The B mode sector scan probe 50 is independent from said transducers 1 or 
5. This is mechanically combined with 5 by the links and joints 51, 52 and 
53 which joints have potentiometers giving the angular data. The position 
calculating circuit 54 calculates the relative position of transducers 5 
and scan probe 50, and the data calculated is sent to the B mode display 
unit 55. The beam location of the transducer 5 is displayed as a line on 
the B mode display unit 56 and the position corresponding to the measuring 
area 4 is marked by increased brightness or by a marker in accordance with 
the depth information given by the depth setting circuit 29. This 
information is used for assigning the measuring area required on the 
sector scanned sectional view of a human body. 
Moreover, the method of the present invention can also be used as a unique 
inspection or diagnostic means, very useful for tissue characterization 
and early detection of disease by executing measurements at each point in 
a bidimensional plane and by displaying the result on the display unit as 
a plane image. 
In this case, it is easier and more effective to get the relative 
distribution of the critical pressure P.sub.c of bubble generation than it 
is to measure the absolute pressure of each point. Namely, in each organ 
of a living body, the composition and temperature of the cell liquid is 
different and the critical pressure P.sub.c is also different in each 
organ. But, it is not easy to obtain the critical pressure P.sub.c by 
calibrating it by using a separate measuring method at another measuring 
area as shown in the case of blood. In such a case, contrary to the 
measurement of blood pressure, the critical pressure P.sub.c can be 
measured relative to the atmospheric pressure, on the supposition that the 
absolute pressure is almost constant with respect to the atmospheric 
pressure (this condition is almost always true except for the area near 
the heart), and the tissue characterization can be made by observing the 
distribution in the bidimensional plane. 
The critical pressure P.sub.c of each tissue changes as time ellapses, 
depending on total body activity, such as exercise, eating and sleeping 
etc. and this variation also appears in the blood. A clearer tissue 
characterization image, eliminating said aging variation, can be obtained 
by simultaneously measuring the critical pressure of the blood and 
displaying the critical pressure of each tissue with the relative value 
obtained from said critical value of blood. 
FIG. 8 shows the block diagram of an embodiment for such object. The 
measuring system 80 is similar to that in FIG. 7. The only difference from 
FIG. 7 is that the low frequency and high frequency ultrasonic wave 
transducers 1 and 5 can move in the vertical direction by means of the 
pulse motor 81. In addition, the measuring location scanning control 
circuit 82 operates in such a manner as to sequentially advance y of the 
bidimensional coordinate (x, y) and also advance x at a high speed for 
each y. 
The drive circuit 83 operates the pulse motor 81 and generates the 
specified number of pulses for each advance of y, shifting the transducers 
1, 5 to the specified pitch. The bubble generating detection circuit 84 
monitors the reflection of the high frequency ultrasonic waves sent from 
the orthogonal detection circuit 31 in FIG. 7 and detects the rise time. 
The sampling circuit 85 samples the low frequency ultrasonic wave 
amplitude sent from the sweep waveform storing circuit 34 in FIG. 7, 
namely the relative sweep pressure value as compared with the atmospheric 
pressure at the time when bubble generation is detected by said bubble 
detection circuit 84. Numeral 86 represents the measured-value temporary 
holding circuit, and numeral 89 represents the subtraction circuit. The 
bidimensional memory 88 to receives an output value of the subtraction 
circuit 89 which is then written at the address x, y. The display unit, 87 
displays the measured value at each coordinate (x, y) stored in the 
bidimensional memory 88 in the brightness or color tone in accordance with 
the values as the plane image. A value of x from 82 is used to set the 
measuring area depth for 29 in FIG. 7. 
In this embodiment, the transducers 1 and 5 are first applied to the upper 
part of arm in order to measure the critical pressure P.sub.c of the 
artery blood with the procedures explained for FIG. 7 and the value 
obtained is temporarily stored in the holding (register) circuit 86 as a 
reference. Thereafter, the transducers 1 and 5 are moved to the desired 
part of the body and the critical pressures of respective measuring areas 
are measured by sequentially changing x and y. These measured values are 
compared with the value in the register 86 by means of the subtraction 
circuit 89 and the difference obtained is written into the bidimensional 
memory. Instead of subtraction, the ratio of tissue critical pressure to 
blood critical pressue can be used. 
According to the system as explained above, measurement for one display 
format can be realized in about 25 seconds, for example, by changing x, y 
respectively from 1 to 500 in order to obtain the picture elements of 
500.times.500 and the continuous wave of 10 kHz is used as the low 
frequency ultrasonic wave. In practice, the propagation time is different 
for areas far from or near to the transducer, and some delay is caused by 
the drive of the pulse motor. Consequently, a little more time is 
required. A higher speed operation can also be realized by employing the 
phased array type transducers 1, 5. The scanning in the direction y is 
carried out electronically and measurements on individual points on the 
same line are carried out simultaneously by providing a plurality of pairs 
(for example, 500 pairs) of 27, 28, 29, 30, 31, 34, 84, 85 in FIG. 7. In 
this case, scanning is carried out for the polar coordinates in stead of 
the orthogonal coordinates and y is the deflection angle, while x is the 
distance from the center. 
Since the critical pressure P.sub.c of tissue is generally comparatively 
low, it is practical to raise the critical pressure P.sub.c by dissolving 
an inert gas such as helium, krypton and xenon or carbon dioxide gas. 
Moreover, the application field, which is similar to that of the tracer 
method using a radioactive isotope, can be developed by injecting a 
chemical substance which selectively works on the particular tissue and 
largely changes its critical pressure. 
As explained above, according to the present invention, an inner pressure 
of an industrial system or a living body can be measured non-invasively by 
detecting an ultrasonic cavitation generated by an ultrasonic wave, 
resulting in the effect of measuring internal pressure safely without risk 
and without invading a system or causing a living body to come to death, 
moreover without causing a pain and without any fear of introducing 
impurities or infectious disease. In addition, since the measuring area 
can be changed from the outside, pressure distribution can also be 
measured on a real time basis. 
It is also known that a high or low harmonic frequency ultrasonic wave is 
generated during bubble generation and/or collapsing. The high frequency 
ultrasonic wave unit 5' is no longer necessary, if such harmonic 
ultrasonic waves are used for the detection of bubbles.