Bedside Ultrasound : Principles & Instrumentation - SD

Introduction to Bedside Portable Ultrasound

Hi, I am Dr. Christopher Merritt, professor of radiology at the Thomas Jefferson University Hospital in Philadelphia.

As you know, ultrasound has been in use clinically for many years and has proved to be invaluable for the evaluation of many conditions in the abdomen and pelvis.

One of these is localization of fluid collections for aspiration or drainage such as paracentesis and thoracentesis.

The recent development of small handheld units has made these devices available to the clinician for use at the bedside.

In order to use these effectively, it's important to understand some of the basic principles of ultrasound and some of the tips regarding the application of these devices.

This presentation deals with bedside portable ultrasound.

I'm Dr. Christopher Merritt from the Department of Radiology at the Thomas Jefferson University Hospital in Philadelphia.

Common Indications for Bedside Ultrasound

Common indications for bedside ultrasound include the identification of fluid and localization of fluid collections such as ascites and pleural effusion and guidance for drainage of fluid from the abdominal cavity or pleural cavity.

Training for the Use of Portable Ultrasound Devices

The training for the use of portable ultrasound devices for bedside ultrasound includes first an overview of the principles of ultrasound and the instrumentation used for ultrasound imaging.

This should be followed by hands-on training and the use of the specific scanner and the identification of anatomic landmarks.

Useful and guidance for thoracentesis and paracentesis.

Finally, before credentialing, monitoring of several procedures is essential.

The process of training concludes with a self-assessment examination.

Basic Principles of Ultrasound

In order to best use ultrasound, it's important to understand the basic principles that underlie ultrasound as an imaging technology.

The Ultrasound Transducer and Piezoelectric Effect

The ultrasound transducer is based on the piezoelectric effect.

This is the property of certain materials to change their shape when stimulated electro electrically.

The ultrasound transducer when a voltage is applied to it changes its shape slightly.

This produces pressure which can be propagated into tissue as a mechanical pressure wave.

Similarly, the transducer when struck by the pressure wave of a reflected ultrasound echo is slightly deformed.

This produces a voltage across the transducer, which can be detected, amplified, and ultimately used to display the echo information in the form of an ultrasound image.

The typical ultrasound transducer consists of a number of small Paso electric elements linked together electrically and fired and precise timing sequences to generate a two dimensional image.

Ultrasound Energy and Wave Characteristics

The ultrasound energy, which is pressure or mechanical energy, consists of changes in pressure elevation and rare affection, which are transmitted from molecule to molecule throughout the tissue.

The pressure wave moves rapidly through the tissue.

The characteristics of the pressure wave can be described in two fashions.

The spatial distance between corresponding portions of the wave can be measured usually in millimeters, and this is the wavelength Or the time that it takes for a complete pressure cycle to pass.

A given point can be measured.

This is measured in seconds and it's called a period.

A typical ultrasound transducer has a very short period and this example, 5000000th of a second.

More commonly the inverse of the period is referred to this is the frequency, and in this example, the period of one 5000000th of a second corresponds to a frequency of 2 million cycles per second or two megahertz.

Acoustical energy is part of the continuum of sound, which includes audible sound.

Industrial applications of ultrasound and medical ultrasound.

Audible sound is at the low end of the acoustic spectrum in the range of 100 to 20 kilohertz, whereas medical ultrasound is at the high end of the spectrum with most devices operating in the range between one and 20 megahertz,

the ultrasound pressure wave is propagated at a constant velocity through tissue.

The propagation velocity is proportional to the frequency and the wavelength, and in this example is equal to the product of the frequency and the wavelength.

Therefore, as the frequency increases, the wavelength decreases and conversely, as the frequency decreases, the wavelength increases.

Propagation Velocity in Different Tissues

The propagation velocity of sound is assumed to be constant by the ultrasound device.

However, this is not completely true.

For example, air within gas in the bell or within the lung propagates sound at a relatively slow rate of about 330 meters per second.

Fat and fluid propagate sound more rapidly at around 1,450 to 1,480 meters per second.

Soft tissue has an average propagation velocity of 1,540 meters per second, and it is this velocity that the ultrasound machine uses in calculating the position of echoes and displaying them in the ultrasound image.

Okay, Blood propagates sound slightly more rapidly and bone propagates sound extremely rapidly at a speed of approximately 4,000 meters per second.

These slight variations in propagation velocity are responsible for some of the artifacts that are seen with ultrasound.

Since the machine assumes that the sound is propagated and all tissues at the same velocity.

Namely 1,540 meters per second,

Continuous Wave vs. Pulse Ultrasound

ultrasound can be produced from a transducer in a continuous fashion, which is called continuous wave ultrasound.

This method is utilized in simple handheld pocket doppler devices for the identification of blood flow and pressure measurement.

The rest of ultrasound involves short pulses of ultrasound delivered by the transducer.

Ultrasound imaging and doppler are all based on the use of pulse ultrasound.

In these cases, the ultrasound pulses are quite short consisting of a few cycles of sound separated by quite long intervals in which no sound is propagated.

The ultrasound transducer produces pulses of sound which propagate within the tissue until they strike a reflecting interface.

When an interface is encountered, some of the energy is reflected back to the transducer where it is detected and ultimately displayed as a part of the image.

Echo Ranging and Image Generation

The principle by which the ultrasound machine determines where an echo is located is called echo ranging.

In a sense, the ultrasound machine is nothing more than a fancy stopwatch.

If we can imagine a transducer at the surface of the liver With a stopwatch attached, when we fire a pulse, we start the stopwatch.

It takes a finite amount of time for the pulse to reach the echo of interest and the same amount of time for the echo to return from the interface.

In this case, the example shows that the time from transmission of the ultrasound pulse until the echo returns is 0.145 milliseconds.

Once we've measured the time from the transmission of a pulse until the echo returns, it becomes a matter of simple arithmetic to calculate the distance that the sound has traveled.

And this example, the total distance is 22.4 centimeters.

Since this represents the time from the transmission of the pulse until the return of the echo, we divide by two to determine the depth of the reflecting interface producing the echo from the transducer.

In this case, the spot that we've marked lies at a depth of 11.2 centimeters from the transducer.

By repeating this process, literally millions of times, it's possible to generate a two dimensional image of all of the reflecting structures, within the ultrasound field, and this in turn can be used to generate the two dimensional realtime ultrasound image.

With B Mode imaging, multiple pulses of ultrasound are sent down and a series of scan lines this quickly builds up a two dimensional image of the echoes in the underlying tissue.

This process is repeated many times per second, as many as 20 or 30 frames per second are generated rapidly in this session, giving the appearance of motion or real time Transducers can be designed to produce a rectangular image format by transmitting parallel lines of ultrasound energy to build up the image or the ultrasound beams can be steered to produce a sector format.

These are generated at a frame rate of up to 28 frames per second.

Attributes of Excellent Imaging

To produce a real time gray scale image, excellent imaging requires a number of attributes.

These include high spatial resolution, excellent contrast resolution, good temporal resolution, and freedom from artifacts.

I'd like to explore each of these attributes in a little bit more detail as they're important in understanding the ultrasound image and in optimizing the ultrasound image.

Resolution Components

First, let's talk about resolution.

The ultrasound image is a two dimensional image, but there are three components to the resolution displayed in the image.

The first is the axi resolution, the resolution along the axis of the ultrasound beam.

The axial resolution determines the rehabilitate to differentiate as separate two objects that lie along the path of the ultrasound beam.

In this example, the arrows point to two tiny bright echoes which are microcalcifications in a breast cancer.

These are identified as separate calcifications rather than merging together as a single echo as a result of the high axial resolution of the system used for this image.

Axial resolution is related to the wavelength or frequency as we can see in this graph.

As the frequency increases from one megahertz to 15 megahertz, the wavelength associated with that frequency decreases at one megahertz.

The wavelength is approximately 1.5 millimeters, whereas at 10 megahertz, the wavelength is approximately 0.15 millimeters.

Axial resolution is determined by the length of the ultrasound pulse, which typically typically consists of two to three wavelengths.

Therefore, the pulses will become shorter as the frequency increases.

If we look at an ultrasound pulse as it passes between two nearby objects along the path of the ultrasound beam, we can see that if the pulse is longer than the distance between the objects, the echos produced by these objects will merge and appear on the image as a single object rather than as two separate objects.

If on the other hand, a shorter pulse is used by using a higher frequency transducer, then this pulse will produce two discrete echoes.

The actual resolution for the P 10 system is estimated to be in the order of 1.1 to 2.5 millimeters.

Another component of resolution is called a lateral resolution.

This is the resolution in the plane perpendicular to the ultrasound beam and is indicated by the L in this image.

Notice that the lateral resolution changes with position from the transducer.

The lateral resolution is poor near the transducer narrows to a point indicated by the parallel lines and arrows, and then diverges again deep to this point.

Lateral resolution defines the ability to differentiate objects that lie side by side and the plain perpendicular to the direction of the ultrasound being.

Again, we see two calcifications lying side by side in these images which are defined by the lateral resolution of the ultrasound device.

A third component of resolution is the elevation resolution.

This is determined by the thickness of the ultrasound slice.

This cannot be varied by the user, unlike CT and MR in which the image is derived from measurements of physical properties of tissues.

Assumptions in Ultrasound Imaging and Resulting Artifacts

Ultrasound is quite different.

Ultrasound basically is based on a number of assumptions about how the acoustical energy interacts with tissue.

These inceptions are important because frequently they are not fully met in the generation of the image, and this can result in artifacts which may be confusing or misleading.

The first assumption is that the ultrasound beam is narrow and uniform in width.

As we saw in the preceding slides.

The lateral resolution in fact is not uniform.

It may be quite large or poor near the transducer and far away from the transducer and is optimal only in the focal zone.

In addition, most transducers produce energy which extends outside the primary beam.

In this case, a scle in photograph shows a beam profile which narrows to a focal point as well as a additional ultrasound field which extends to the side indicated by the arrow.

This is an area of relatively high acoustical uh, energy, which may interact with strong reflectors and produce important artifacts.

If, for example, a strong reflector such as gas or bone lies outside the primary beam and encounters a side lobe of high energy from the transducer, the echo from that interface may return to the transducer.

The machine assumes that the echo is coming from along the beam path and therefore displays the echo in an incorrect location within the image.

These are called side lobe artifacts and may be quite conspicuous.

The arrow here points to an artifact projected over the distended bladder and the pelvic ultrasound.

This side lobe results from the interaction of the acoustic field of the transducer off axi with gas and adjacent bowel within the female pelvis side lobe artifacts are important and may particularly pose a problem in the examination of areas in which both gas and fluid are present.

These are both conditions that exist in the evaluation of ascites and pleural effusions, which are common bedside procedures performed With handheld ultrasound devices.

Here, for example, is a view of the abdomen in a patient with ascites, the arrows point to two linear echoes extending to either side of a loop of bowel.

These in fact are artifactual rather than real and are side lobe artifacts.

By slightly changing the angle and position of the transducer, these artifacts disappear.

Another assumption, uh, involved in ultrasound is that the ultrasound beam travels directly to and from the echo producing interfaces.

This then brings us to a discussion of what kinds of structures reflect, sound and serve as acoustical interfaces.

There are two types of reflectors, diffuse scattering reflectors and large mirror-like or specular reflectors.

Diffuse reflectors arise from very small interfaces smaller than the wavelength of the ultrasound used to produce the image.

When ultrasound is interacts with these small reflectors, the sound is scattered in all directions, a small amount of which may return to the transducer to produce the image.

Diffuse reflectors produce the characteristic echo pattern associated with the parenchyma of organs such as the liver and spleen, kidneys, and uterus.

A Large flat surfaces act more like mirrors.

This view shows the liver with the bright curva, linear echo of the diaphragm along the back of the image.

Notice that a portion of the diaphragm is quite bright and well-defined, but as we go to the left side of the image, we see that the diaphragm fades from view.

This is because the portion of the diaphragm, which is nearly perpendicular to the direction of sound, reflects sound directly back to the transducer, much as a mirror would, whereas sound striking, the portion of the diaphragm to the left reflects off away from the transducer, and as a result that portion of the diaphragm is not displayed.

The strength of a ultrasound interface determines the amount of sound that is reflected to the transducer and the amount that continues on, uh, to deeper structures.

The amount of, uh, sound reflected in an interface is determined by the difference in the acoustical impedances of the tissues making up the interface.

The acoustical impedance of each tissue is in turn determined by the density of the tissue and the propagation velocity of sound within that tissue.

An interface between tissue with a typical propagation velocity of 1,540 meters per second and a density of approximately one and fluid with a slightly lower propagation velocity and similar density allows some of the sound to be reflected at the interface back to the transducer and a considerable amount of the sound to pass on to image deeper structures.

In contrast, an interface between tissue and air, which has very low density and low propagation velocity or bone, which has high density and high propagation velocity, will reflect essentially all of the incoming ultrasound preventing the visualization of deeper structures.

Therefore, when we look at the interface between the diaphragm and the lung seen on the left of this image, we see that no image is seen to the far left as it is all reflected by the strong interface formed by the diaphragm and the aerated lung.

Similarly, this view of a gallstone shows a bright echo at the surface of the gallstone and then deep to the gallstone, a dark area which represents a acoustical shadow due to the fact that essentially all of the incoming energy is reflected by the surface of the stone and therefore no energy passes through the stone to provide echoes from deeper structures.

Another assumption of the ultrasound device is that the attenuation as the sound passes through tissue is uniform.

This assumption is also incorrect.

If we look at sound passing through tissue at two different frequencies, we can see that at five megahertz after sound passes a certain distance, The energy is reduced, and this example, uh, the lowest uh level is 1% of the original energy, but if we go up to 10 megahertz, we see that the energy is relu reduced much more rapidly and at the same depth that removes 99% of the energy at five megahertz.

99.9% is removed at 10 megahertz.

This then leads us to a brief discussion of the term used to describe the process of attenuation.

We describe this in units of decibels.

If we look at two different sites within tissue and the ratio of the power at these two sites, we can see in this example that at point A we have a hundred percent of our power and at point B only 1%.

The ratio of these two powers is one to 100.

The log of this is two.

The decibel notation simply multiplies the log of the ratio by 10.

Therefore, in this case, the difference between point A and point B can be described as 20 decibels.

We can refer to point B as 20 db less than point A, or we can refer to point A as 20 db greater than point B.

In general, with ultrasound, most tissues attenuate at very roughly one decibel per centimeter per megahertz, so with increasing depth there's greater attenuation, and attenuation is more rapid with higher frequencies.

Now, if we simply pass ultrasound into tissue, we'll see that the image gradually fades as a result of this attenuation.

In this example, we see tissue, which at the surface has a normal display level, but at two centimeters we have only 1% of the energy.

We're down 20 db and there's little or no information available at this depth.

Therefore, with ultrasound we have to correct for this attenuation, and we do that by process called time gain or depth gain convers.

Essentially what we do is that as we move deeper and deeper into the tissue, we add gain our amplification to restore the echo level to a nominal level.

For example, in this slide, at one centimeter we've attenuated 10 db and therefore we correct by adding 10 DB of amplification or gain to restore the echo level to that similar to the surface, and at two centimeters we're down 20 DB and therefore need to add 20 DB of gain or amplification to restore the echo level.

This time gain compensation is important in optimizing image quality.

We mentioned that the ultrasound machine assumes attenuation is uniform and that we can correct for this by our time gain compensation, adding amplification for increasing depth, but not all things that we image have uniform attenuation.

For example, on the right we see a cyst which shows very bright, echoes deep to the cyst.

This is a result of the fact that the assumption of uniform attenuation is not met at the surface.

We have no attenuation and no need for adding amplification.

At one centimeter, we've attenuated about 10 DB and tissue and we've added 10 DB of amplification to restore the grade level.

The cyst, however, has attenuated only about three db.

Therefore, by adding 10 DB at this depth, we actually end up above the level at which we started at plus seven db, and this continues as we go deeper and deeper, uh, into the tissue.

Similarly, in this example, we have an object which is producing shadowing.

Here again, the process is the same.

At one centimeter, we've attenuated 10 DB and added 10 DB of gain.

The mass, however, has attenuated more than estimated, in this case, 20 DB rather than 10 db.

Therefore, by adding 10 DB of gain, we still haven't restored the level to the original and or down 10 db.

Finally, the ultrasound device assumes that the speed of sound in tissues is constant.

This may produce, uh, certain minor artifacts and it's not a particularly important concern for most bedside procedures.

Safety and Bio Effects of Ultrasound

Finally, a comment about safety and bio effects.

Ultrasound has been used for many years and to date, no harmful effects of ultrasound at diagnostic levels have ever been demonstrated in humans.

Nevertheless, there are mechanisms by which ultrasound theoretically can produce bio effects.

These include thermal effects due to tissue heating and mechanical effects due to the pressure wave as it moves through tissue resulting in cavitation.

The FDA requires that indices be displayed on the ultrasound display to indicate conditions that favor the deposition of heat or mechanical energy.

The TI value is a value that reflects the potential for heating, and it's called the thermal index.

The MI value is related to the potential for cavitation.

The numbers themselves are meaningless, but they allow the user to monitor changes in these values and to implement the principle of aara.

That is to keep these values as low as reasonably achievable.

This concludes discussion of the basic principles and instrumentation necessary for the understanding and performance of bedside ultrasound.