PRINCIPLES OF ULTRASOUND APPLICATION IN ANIMALS

PRINCIPLES OF ULTRASOUND APPLICATION IN ANIMALS*?

NORMAN W. RANTANEN. DVM. MS. A N D ROBERT L. EWING 111, BS$

Diagnostic ultrasound, a non-invasive mode for imaging soft tissues, requires for its use an understanding of sound and sound-tissue interaction physics. Ultrasound frequencies from 1 .O to 10.0 MHz are created by electrical stimulation of piezoelectric crystals. These crystals are housed within transducers, which, when applied to the body surface with a coupling agent will produce sound waves, referred to as the sound beam, which are propagated through the soft tissues of the body. When the sound beam encounters tissue interfaces of differing acoustic impedance, a portion of the sound beam is reflected back to the transducer which also acts as a receiver. Echoes returning from soft tissue acoustic interfaces are converted to electrical im- pulses and displayed on an oscilloscope screen as a cross section of the tissue. Lower frequency sound beams penetrate further into soft tissue, but have poorer resolving capabilities, than higher frequency sound beams.

A, B, and M-modes are the three basic forms of ultrasound used in soft tissue imaging. A-mode ultrasonic imaging is a one-dimensional display of echo amplitudes versus distance. B-mode ultrasonic imaging produces an accurate two-dimensional cross sectional image of soft tissues. M-mode ultrasonic imaging is an adaptation of B-mode to evaluate moving structures of the heart.

Fluid-filled cystic structures have characteristic clear (anechoic) central areas with acoustic enhancement of the back wall of the cyst and deeper structures. Solid masses have echoes in their central portion with resultant poor accentuation of deeper structures.

Application of ultrasound to animals requires hair removal since trapped air is a barrier to transmission of the sound beam. Gas-filled bowel and bone are effective barriers to ultrasonic imaging because of their large acoustic impedance differences compared to soft tissues.

The position of the focal point of a focused transducer relative to tissue interfaces is important to accurately depict tissue character. For example, the focal point of the transducer should be superficial to the back wall when scanning cystic structures. When solid lesions, such as liver metastases, are scanned a focal point that lies deep to the lesion should be selected in order to accentuate sound beam attenuation.

Time-gain compensation (TGC) settings are important to produce a balanced scan with equal echo production within like tissues throughout the depth of ultrasound tissue penetration. Initial TGC settings can be made from knowledge of the focal point of the particular transducer, but may have to be adjusted during scanning to produce a balanced image.

Ultrasound is an attractive imaging modality in animals since it is noninvasive and presents no known hazard to the operator or patient. Veterinary Radiology, Vol. 22, N o . 5 , 1981; p p I96 -203.

Key words: echocardiography, ultrasonics (physical principles and technics), animal diseases, veterinary medicine.

EPORTS of diagnostic ultrasound (US) use as an R alternative imaging modality for soft tissues have appeared in the veterinary literature. US is defined as sound above an audible frequency. The three principle modes of ultrasound are amplitude mode (A-mode),

* This work was supported in part by the Thomas Autzen and Stan Adler Funds.

t This paper is the third in a series of continuing education papers dealing with the subject of medical imaging.

$ From Washington State University, College of Veterinary Medicine, Department of Veterinary Clinical Medicine and Surgery, Pullman, WA, 99164 where Dr. Rantanen is Associate Professor of Radiology and Mr. Ewing is A Clinical Laboratory Technician.

brightness mode (B-mode), and time motion or motion mode (T-M or M-mode).

A recent review reports the widespread use of A- mode ultrasound in evaluating fat and lean proportions of meat animals since 1956.' In 1966, A-mode ultrasound was used as an aid in pregnancy diagnosis in ewes.2 With the development of grayscale ultrasound, however, US use has extended in other species to the a b d ~ m e n , ~ , ~ . ~ urinary tract,6 and heart.7 While US is a well accepted diagnostic modality in humans, its value remains to be proven in animals. Diagnostic ultrasound has been applied to animals at Washington State University for approximately two and a half

0 196-3627/81/0010/0196/$00.90 @ American College of Veterinary Radiology

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years and has proven valuable as an adjunct in the diagnosis of several diseases. Understanding the basic principles of sound propagation and interaction with body parts is required for a general understanding of US use in diagnosis. Various scanning technics and instrumentation as well as cross sectional anatomy must be understood to apply this imaging modality effectively as a diagnostic aid in animals.

Physical Principles of Ultrasound

Ultrasound is defined as any sound frequency above the normal hearing range of the human ear, i . e . , greater than 20,000 hertz.$ Frequencies commonly used in diagnostic US range from 1.0 to 10 megahertz (MHz) (million cycles/sec).

Sound beam intensity (amplitude) is usually mea- sured in decibels (dB) (0.1 bel). A be1 is defined as the logarithm to the base ten of the ratio of two levels of power, voltage, current, or sound intensity. The dB compares one sound intensity to another. In ultrasound, a sound beam intensity is expressed in watts per square centimeter (W/cm2) and is the rate of flow of energy per cm2. The power of the sound beam refers to the time rate at which work is done. Ultrasonic intensity and power are used interchangeably by some authors .8

Diagnostic ultrasound is produced by transducers housing crystals with piezoelectric (pressure-electric) properties. The term piezo is derived from the Greek word meaning pressure, hence the term pressure- electric. When piezoelectric crystals are deformed by pressure, electricity is produced. Conversely, when an electric current is applied to them the crystals will de- form. This is the process by which ultrasound is gener- ated and received by the transducer. Pulsed electrical deformation of the crystal produces small sound waves which impart kinetic energy to tissue molecules. These pressure waves are propagated through the tissue by the transfer of energy to adjacent molecules. There is a constant relationship between propagation velocity, wave length, and frequency. V = fh, where V is the velocity (meterdsec) , f is the frequency in hertz (cycles/sec), and A is the wave length (in meters). When reflected sound returns to the transducer, a slight deformation of the crystal is produced which generates an electric current. This current is displayed on an oscilloscope as an image of the tissue interfaces.

Transducers vary in size from 3.0 to 19.0 mm in diameter. Each transducer is manufactured to emit primarily one ultrasonic frequency, however, some lower and higher frequencies are also emitted. Ab- sorbing material surrounding the crystal within the transducer removes unnecessary sound emitted in di-

8 One hertz is a frequency of one cycle per second.

Two Field Expand i ng Boundary \ Field \

FIG. 1. Diagram of the two sections of the sound beam from a nonfocused transducer. The near field (Fresnel zone) is approximately the width of the transducer. In the far field (Fraunhofer zone) the beam diverges. (Diagram courtesy of Technicare Corporation, Ultrasound Division, Denver, Co.)

rections away from the transducer face. The pulsed directional pressure waves produced by the transducer are referred to as the sound beam.

Transducers can be focused or unfocused. The ultrasound beam produced by an unfocused transducer has two regions (Fig. 1). The near field (Fresnel zone), the section close to the transducer, is about the same diameter as the transducer. The region beyond the near field is called the far field (Fraunhofer zone) in which the ultrasound beam becomes divergent. The location of the transition between the near field and far field depends on the diameter of the transducer in re- lation to the frequency (wavelength). The relationship d = r2/A expresses the length, d, of the nearfield, where r is the transducer radius, and A is the ultrasonic wavelength. The wide field produced by an unfocused transducer limits the ability to define small structures. This is referred to as lateral resolution, which is the ability to separately define two structures at right angles to the sound beam (Fig. 2). The near field beam width can be narrowed to improve lateral resolution by decreasing the diameter of the transducer. It can, however, be seen from the above*equation (d = r2/A) that when transducer size is decreased the boundary between the near and far field (d) will be closer to the transducer. To alleviate this problem, transducers can be acoustically focused. This can be done by using a

A

FIG. 2. Diagram of the sound beam from a nonfocused transducer demonstrating problems encountered with lateral resolution. The small object, A, in the near field will produce an echo anywhere inside the sound beam. Therefore, when an object smaller than the beam width is scanned it will appear as large as the beam width. A similar problem occurs in the far field where the same size object at B will appear as wide as the divergent beam. (Diagram courtesy of Technicare Corporation, Ultrasound Division, Denver, Co.)

198 RANTANEN A N D EWING 198 1

FIG. 3. Diagram demonstrating two ways to focus a sound beam. The sound beam can be focused by physically bending the crystal as in the upper half of the diagram (internally focused) or reflected with a mirror as in the bottom half of the diagram (externally focused). The sound beam in both instances is narrowed throughout the focal zone, thus improving lateral resolution. (Diagram courtesy of Tech- nicare Corporation, Ultrasound Division, Denver, Co.)

lens or mirror (externally focused), or more recently by producing crystals with a curved surface (internally focused) (Figs. 3 and 4). Focusing a transducer changes the shape of the sound beam and produces a narrow section of the beam referred to as the focal zone, the midpoint of which is called the focal point (Fig. 4). Lateral resolution is best in the focal zone where the beam is narrowest.

The ability to differentiate two separate structures along the axis of the beam is also of concern. This is referred to as axial resolution and is a function of the ultrasonic pulse width. If pulse width is large, structures less than one pulse width apart will appear as a single structure (Fig. 5). The smaller the pulse width, the better the axial resolution. Regardless of the frequency of the transducer, the same number of oscilla-

Crystal Lens I Focal I -- FIG. 4. Diagram of a sound beam externally focused with a lens.

The narrowest portion of the beam is called the focal zone, the midpoint of which is called the focal point. Because the beam nar- rows as the focal point is approached this is the region of the greatest beam intensity (W/cmz). (Diagram courtesy of Technicare Corpora- tion, Ultrasound Division, Denver. Co.)

\ Reflecting H Surf aces

Axial Resolution

4

/ Ultrasound

FIG. 5 . Diagram of the problem encountered in axial resolution. In the upper half of the diagram, the reflecting surfaces will produce one echo because, due to the width of the pulse, the pulse will encounter both at the same time. In the lower half, the shorter pulse will encounter each reflecting surface separately producing two separate echoes. Axial resolution will be greater with the shorter pulse. Pulse width can be shortened by using a transducer of a higher frequency. (Diagram courtesy of Lynn Spolek, Senior thesis pre- sentation, Washington State University 1979.)

tions are required in a pulse produced by electrical stimulation of the crystal. This means that by increasing the frequency (0, the pulse length can be shortened (A = v/Q.

As the sound beam passes through body tissue, a portion of the beam is reflected back to the transducer. Reflection occurs a t tissue interfaces of differing acoustic impedance. Acoustic impedance, Z (g/cm2s), is equal to the product of the density of the tissue, p (g/cm3), and the velocity of sound, V ( c d s ) , in that tissue. Thus, Z = pV. The amount of the beam reflected back to the transducer is directly proportional to the difference in acoustic impedance at the tissue interface. Because of the large difference in acoustic impedance between soft tissue (Z = 1.63 x lo5 g/cm2s), bone (Z = 7.8 X lo5 g/cm2s) and air (Z = 0.0004 X lo5 g/cm2s), most of the sound beam is reflected when it strikes bone or gas.g Thus tissue interfaces deeper than bone or gas are hidden. It must be stressed that the absolute value of the acoustic impedance of any tissue is relatively unimportant, because it is the magnitude of the difference in acoustic impedance at tissue interfaces that determines the amount of reflection of the beam.

Tissue interfaces with large differences in acoustic impedance causing almost total reflection of sound, e . g . , soft tissue and gas, may cause the sound to reverberate between the crystal and the gas. This is commonly seen when scanning over lung tissue or gas-filled bowel. As sound reflected by the gas returns to the transducer face, a portion of the reflected beam is again reflected back to the gas interface because of the difference in acoustic impedance of the body sur-

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FIG. 6. Ultrasound scan of a normal adult horse. The top of the figure is dorsal and the horse is viewed from behind. The scan was produced by passing the transducer over the 11th intercostal space from a point dorsal to the lung border to a point ventral to the costal cartilage. Reverberation of sound between the transducer contacting the thoracic wall (T) and the lung surface (P) produced multiple concentric echoes (R). As the transducer passed ventral to the lung, the diaphragm (D) and spleen (S) allowed penetration of the beam to the medially located gas-filled bowel (G). The open arrow indicates an acoustic shadow produced by the costal cartilage as the transducer passed over it. Because of the large difference in acoustic impedance between the soft tissues of the intercostal muscles and the dense cartilage, the sound did not penetrate the cartilage. Marker dots are 1 cm apart.

face and transducer face. Reflected sound striking the transducer produces a recordable echo. The portion reflected may reverberate between the gas-transducer interface a number of times. Each time sound returns to the transducer it produces an image on the oscilloscope screen at a location equal to the delay during its travel time between the reflecting surface and transducer (Fig. 6).

Energy is removed from the sound beam as it passes through soft tissues. This energy removal is referred to as attenuation. Although attenuation is caused by several processes, two predominate.'O The first process is absorption which is the conversion of ordered motion of ultrasound into the disordered motion of heat. The amount of absorption increases with the frequency of the sound beam. The second process is scattering of the sound beam by small (less than a wavelength) tissue interfaces. This is called Rayleigh scattering. It occurs in all directions and results in energy loss from the sound beam. The intensity of the scattered sound increases with its increasing frequency and is proportional to the square of that frequency.s

Ultrasound attenuation by soft tissue is at a rate of approximately 1 dB/crn/MH~,~." while the attenuation

Signal Level

lll A 1

I I

I

I l l , . I D i sta nce 1

B Signal Level

Distance

I

Gain I

I

I

I

I

I 1

Distance FIG. 7. Diagram demonstrating the effect of attenuation of the

sound beam with penetration into soft tissues. (A) The signal level decreases as the sound beam is attenuated and energy is lost. Therefore, superficial echoes are higher in amplitude than echoes from deeper interfaces. (B) The middle diagram, represents the ideal situation where echo amplitude (signal level) is balanced from SU- perficial to deep interfaces. This is accomplished by electronically increasing the gain of deeper echoes so equal reflectors have an equal amplitude. The bottom one-third of the diagram represents the time gain compensation (TGC) utilized to achieve this equality of echo amplitudes. (Diagram courtesy of Technicare Corporation, Ultrasound Division, Denver, Co.)

rate of US by skeletal muscle is approximately 3.0 dB/cm/MHz." Because the factors causing absorption and scattering of the beam are frequency dependent, lower frequency sound will penetrate further into soft tissue than higher frequency sound. This contrasts with electromagnetic radiation used in conventional radiography where higher frequency radiation pene- trates deeper than radiation of lower frequency.

Because of sound wave attenuation, echoes from tissue interfaces near the crystal will be higher in amplitude than those from more distant tissue interfaces. To compensate for this, the amplitude of echoes from tissue interfaces near the crystal are electronically

UNIRAD Application Seminar, Denver, CO., Oct, 1979.

200 RANTANEN A N D EWING 1981

suppressed and the amplitude of echoes from more distant tissue interfaces are progressively increased. This is referred to a s t ime gain compensat ion (TGC)y,'",'2 (Fig. 7). The goal of TGC is to receive echoes of the same amplitude from similar tissue interfaces down to the effective depth of the transducer. Generally, the effective depth of a focused transducer is equal to two times its focal point (midpoint of focal zone). This is the approximate depth to terminate the TGC. A 40 dB initial suppression of the near echoes (starting point of the TGC curve) is generally recom- mended for instruments with a 40 dB dynamic range.'" This TGC curve can be used as a starting point, however, adjustments may have to be made based on the appearance of the scan. Overall gain can be increased or decreased to affect all echoes similarly.

Based on tissue attenuation rates of 1 dB/cm/MHz it is often possible to gain information from tissue interfaces beyond the theoretical limit of the transducer. One explanation for this is that the sound beam may pass through tissues or fluids that cause less than theoretical or expected attenuation of the beam. An- other explanation for this is the presence of sound frequency lower than the primary frequency which the transducer was designed to emit. Although a primary frequency predominates in the beam, lower and higher frequencies are present. The lower frequencies will penetrate to greater depths because of lower attenuation rates.'"

Ultrasound Display Modes

Most US equipment is calibrated at lo3 energy pulses per second and an average tissue velocity of 1540 m/s at 37°C."13 The connection of the transducer to the console has a series of encoders which spatially orient the returning echoes to display accurately the acoustic interfaces of the tissue on the display screen.

There are three main display formats or modes" (Fig. 8):

(a) Amplitude mode (A-mode) ultrasonic imaging is a one-dimensional display of returning echo amplitude and distance. Each peak represents a returning echo, the height of which is proportional to its amplitude. Most instruments have an A-mode display.

(b) Brightness mode (B-mode) ultrasonic imaging is a two-dimensional display of dots. The transducer is moved across the surface of the body, and cross sectional anatomy is depicted. The position of the dot on the screen is determined by the time it takes for an echo to return to the transducer. The brightness of the dots is proportional to the amplitude of the returning echoes. Fig. 9 is a diagram of the articulated arm of a B-mode scanner. Encoders are located in the arm to spatially orient the returning echoes. Real time ultrasonic imaging is a form of B-mode used to record movement of structures. In real time, imaging, echoes

0 . . . . 0 - - . . . .

FIG. 8 . Composite diagram of the three basic modes of ultrasound. (1) A-mode (Amplitude mode) is a one dimensional display of echo amplitude versus distance. The sloped line terminating at A represents the time gain compensation (TGC) curve used to produce echoes of like amplitude from like interface reflections (see Fig. 7). The transducer is located on the left of the display. (2) B-mode (brightness mode) depicts a two-dimensional cross section of the tissue. As the transducer is moved over the surface of the body the dots (3) form lines in the tracing conforming to tissue interfaces. (4) M-mode (motion mode) is similar to B-mode in principle but the transducer is held stationary and the dots are recorded on moving light-sensitive paper. The central organ in parts two and three, in the case of M-mode, could represent cardiac chamber wall motion. M-mode is used in echocardiography.

are recorded continuously on a nonstorage cathode ray display screen which is analogous to image intensified fluoroscopy except that cross sectional anatomy is displayed. This image may be frozen and photo-

FIG. 9. Diagram of the articulated arm used in B-mode ultrasound. Encoders at one and two of the arm and three at the transducer attachment spatially orient the returning echoes on the display screen to depict tissue interfaces accurately. (Diagram courtesy of Technicare Corporation, Ultrasound Division, Den- ver, Co.)

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graphed or recorded on videotape. The transducer head of real time units is attached by a flexible cable. Encoders are contained within the moveable head to allow rapid transducer movement from one area to another in contrast to B-mode scanning arms.

(c) Motion mode ultrasound (time motion; M or TM-mode) is a one-dimensional format displaying dots, as in B-mode. With M-mode, however, the transducer is held in place over moving organs and the display is printed on an oscilloscope or moving strip of light-sensitive paper. M-mode is used primarily in echocardiographic studies to measure cardiac wall motion and valve excursions. An actual image of the heart is not produced.

Biological Hazards

Diagnostic ultrasound is considered safe in its present manner of medical application. Sound pressure waves produce heat which can cause cellular physiologic changes. Diagnostic US is pulsed, however , and has a low average intensity (3 t o 10 milliwatts/cm2). Its effects are momentary, as heat is dissipated in the fluids of the body. In contrast to diagnostic US, therapeutic ultrasound employs higher in- tensities (0.5 to 3.0 watts/cm2). No deleterious effects are found when therapeutic ultrasound is used prop- erly.I3

B-Mode Application to Animals

Because US is poorly propagated in a gaseous medium, hair must be removed from the area of interest and a coupling agent (mineral oil or aqueous gel) used between the skin and transducer. This applies to all modes of ultrasound.

Reduction of patient motion is important. Chemical or physical restraint is most often advisable.

Transducers producing low US frequency should be selected in order to allow maximum sound penetration into the patient. Higher frequency transducers can then be employed to improve resolution of tissue interfaces. The TGC curve must be changed when different transducers are used. Pattern of movement of the transducer is important in producing good quality scans. A single pass over the area to be scanned is preferable, with the transducer held perpendicular to the surface of the underlying organs whenever possible. This favors recording echoes that would otherwise be lost by reflection in a direction away from the transducer. Sector scanning, performed by pivoting the transducer in a single motion in the plane of the scan instead of passing it over the body surface, is often advantageous depending on the surface contour and the organs of interest. This technic works well between ribs.

US scanning can be performed to depict the entire cross section of small animals. This is done by making

FIG. 10. Transverse compound ultrasound scan of a normal dog's urinary bladder (B). The dog is in dorsal recumbency and is viewed from behind. Therefore, the dog's left side is to the reader's right and vice versa. The transducer has been passed over the abdomen in multiple strokes (T) to produce an image of the entire cross section of the caudal abdomen. This technique is used to image the entire cross section but is prone to production of artifacts (A) since the same tissue is scanned from different angles. Marker dots are 1 cm apart.

multiple passes of the transducer over the surface around the circumference of the body. Compound movements may, however, produce unwanted artifacts because the beam passes over the same tissues more than once, and a greater chance of patient movement is possible (Fig. 10).

Organs of interest should be scanned in two planes at 90" angles to each other to thoroughly evaluate the internal organ interfaces as well as organ size, shape, and position. Anatomic points of reference should be utilized to locate areas of interest when the scans are reviewed after the examination. For instance, the film on which the scan is recorded should be marked ac- cording to the plane of the scan, e . g . , transverse or longitudinal; and the distance from a palpable or visi- ble surface landmark should be recorded. Some method of photographing the scan for a permanent record should be available. Most instruments have Polaroid or multi-format cameras.

The ability to differentiate solid from cystic structures with US is extremely valuable in characterizing mass Fluid-filled cysts characteristically are clear (anechoic) with strong accentuation of the back wall (Fig. I I ) . This is referred to as acoustic enhancement of structures deep to the cyst. This is evi- dence that the sound beam passed through the fluid relatively unattenuated compared to sound in adjacent structures. The echoes are stronger from the back wall of the cyst and deeper structures compared to adjacent structures. The sound intensity is greater in the sound beam transmitted through the cyst. Homogeneous, non fluid-filled structures, e .g . , freshly clotted blood, appear similarly anechoic, but the acoustic enhancement of deeper structures is not as apparent. This is

202 RANTANEN A N D EWING 1981

FIG. I 1 . A transverse ultrasound scan of a normal dog's liver and gall bladder made at the level of the xiphoid cartilage of the sternum. The dog is in dorsal recumbency and is viewed from behind. There- fore, the dog's left side is to the reader's right and vice versa. The gall bladder (GB) demonstrates the appearance of a cystic structure. There are no central echoes and there is acoustic enhancement of the liver and right hemidiaphragm (D) between the arrows deep to the gall bladder. This is caused by lack of attenuation of the sound beam by the fluid bile and results in higher amplitude echoes from the back wall and deeper tissues. To obtain an accurate image of the cystic appearance of the gall bladder the focal point of the transducer should be within or superftcial to the gall bladder. Hepatic veins (V) are seen passing through the liver. Marker dots are 1 cm apart.

because the sound beam is attenuated by scattering (Rayleigh scattering) and energy has been removed from the beam.

A transducer with a focal point that can be posi- tioned superficial to the deep wall of the cyst or homogeneous tissue should be used. This will accentuate homogeneity because the section of the beam encountering the cyst will be beyond the focal point and will be of a lower intensity. This is important because in the narrowest portion of the beam, before the focal point is reached, the sound intensity will be greatest (Wkm') and if this portion of the beam en- compasses the cyst, artifacts may be produced. This could produce nonstructural noise and apparent echoes within the cystic structure giving an erroneous impression of solid tissue.I0

Certain organs may act as acoustic windows. Acoustic windows attenuate little of the sound beam, thus allowing more of the sound to pass beyond it to more distant structure^.^ In man, the full urinary bladder is used as an acoustic window to more effectively scan pelvic organs. The poorly echogenic spleen of the horse has proven to be an aid in visualizing the left kidney in studies at Washington State University (Fig. 12).

Solid structures with a composite of tissue types will be echogenic with poor definition of the back wall (Fig.

FIG. 12. Ultrasound scan of a normal adult horse. The top of the figure is dorsal; the horse is viewed from behind. This scan was made by passing the transducer from dorsal to ventral over the surface of the left flank near the 18th rib. The poorly echogenic spleen (S) next to the body wall (BW) is used as an acoustic window to image the left kidney (K). Because the spleen attenuates the sound beam very little, the medial surface of the kidney can be imaged. The echoes in the central portion of the kidney are from the dense connective tissue of the collecting system. The kidney parenchyma is relatively anechoic. Marker dots are 1 cm apart.

13). The focal point should be placed beyond or at the levei of the far border of suspected solid structures (such as tumor metastases in the liver). Placing the focal point beyond or at the level of the lesion will accentuate the attenuation in the area beyond the lesion when compared to the surrounding parenchyma.'O

The sound beam will not penetrate beyond bone, calculi, or soft tissue calcification and an acoustic

FIG. 13. Ultrasound scan of preputial metastasis from a seminorna in a horse. The horse was in dorsal recumbency, its head to the reader's left. The transducer was passed over the skin surface. The scan demonstrates high amplitude echoes from the internal structure of the metastatic lesion; there is poor definition of the deepest extent of the lesion (arrows). This appearance is characteristic of a nonhomogenous, solid (echogenic) mass. Because of attenuation of the sound beam by the Lumor tissue interfaces, echoes of the deep wall of this well-encapsulated lesion are not enhanced as they are with a cystic structure (compare with Fig. 1 I ) . Marker dots are 1 cm in diameter.

VOL. 22, No. 5 ULTRASOUND APPLICATION I N ANIMALS 203

shadow is produced. This acoustic shadow appears as an anechoic area deep to the calcification. Acoustic shadows often delineate poorly mineralized calculi. l4

If the focal point of the transducer is placed beyond the calcified structure the acoustic shadow will be accen- tuated as in metastatic lesions (Fig. 6). Acoustic shadowing can create problems with scanning if areas of interest are beneath normal skeletal structures such as ribs.

REFERE

At Washington State University Veterinary Hospital US is important as an alternative imaging modality. It is particularly important in characterizing the internal structure, o.g., solid, cystic, or mixed, of abdominal masses in animals. It has applications in pregnancy diagnosis and fetal measurement in horses, characteri- zation of pleural effusions in horses,15 and the investi- gation of abdominal or subcutaneous mass lesions in several species.

NCES

I . Stoufer JR, Westervelt RG. A review of ultrasonic applica-

2. Lindahl 11. Detection of pregnancy in sheep by means of ul-

3. James AE, Brayton JB, Novak G, et al: The use of diagnostic

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trasound. Nature London 1966; 212:642-3.

ultrasound in the evaluation of the abdomen in primates.> Med Primatol 1976; 5: 160-75.

4. James AE, Osterman FO, Bush RM, et al. The use of compound B-mode ultrasound in abdominal diseases of animals. JAVRS

5. O'Grady JP, Yeager CH, Thomas W, Estra G, Findleton L. Practical applications of real time ultrasound scanning to problems of zoo veterinary medicine. J Zoo An Med 1978; 9(2):52-6.

6. Cartee RE, Selcer BA, Patton CS. Ultrasonic diagnosis of renal disease in small animals. JAVMA 1980: 176:426-30.

7. Pipers FS, Hamlin RL. Echocardiography in the horse

8. Hagen- Ansert SL. Textbook of diagnostic ultrasonography.

1976: 17:106-12.

JAVMA 1977; 170:815-19.

St. Louis: CV Mosby, 1978.

9 . Bartrum RJ, Crow HC. Greyscale ultrasound: a manual for physicians and technical personnel. Philadelphia: WB Saunders, 1977.

10. Powis RL. Transducers, tissue-ultrasound interaction, and TGC. Radiology Today 1980; 2(4):18-28.

11 . McDicken WN. Diagnostic ultrasonics. New York: John Wiley & Sons, 1976.

12. Powis RL. Ultrasound Physics for the Fun of It. UNIRAD Corp, Denver Co. 1978.

13. Sikov MR, Hildebrand BP. Advances in the study of birth defects: teratological testing. Vol 11, Persaud TUN, ed. Baltimore: University Park Press 1979; 267-91.

14. Pollack HM, Arger PH, Boldberg BB, Mulholland SG. UI- trasonic detection of nonopaque renal calculi. Radiology 1978;

15. Rantanen NW, Gage L, Paradis M: Ultrasonography as a diagnostic aid in pleural effusion in horses. Vet Radio1 1981;

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PRINCIPLES OF ULTRASOUND APPLICATION IN ANIMALS