The principle of ultrasound – echopedia types of electricity consumers

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Propagation speed in human soft tissue is on average 1540 m/s. It is defines as to how fast the ultrasound can travel through that tissue. It is determined by the medium only and is related to the density and the stiffness of the tissue in question. Density of the medium is related to its weight and the stiffness of the medium is related to its “squishability”. As the medium becomes more dense, the slower is speed of ultrasound in that medium (inverse relationship). The stiffer the tissue, the faster will the ultrasound travel in that medium (direct relationship). There are tables where one can look up the velocity of sound in individual tissues.

So far we have defined the ultrasound variables and parameters. In the next section will talk more about pulsed ultrasound. Pulse Duration is defined as the time that the pulse is on. It is determined by the number of cycles and the period of each cycle. In clinical imaging, a pulse is comprised of 2-4 cycles and the pulse duration is usually between 0.5 to 3 microseconds. Pulse duration does not change with depth, thus it cannot be changed by the sonographer. Pulse Duration (msec) = # of cycles x period (msec). Since Wavelength (mm) = Propagation speed in tissue (mm/microsecond) / frequency (MHz), this can be rewritten as 1/frequency = wavelength / propagation speed. mp electricity bill pay indore And since period = 1/frequency, then the Pulse Duration = (# of cycles x wavelength) / Propagation speed.

A related parameter to PRP is the Pulse Repetition Frequency or PRF. PRP and PRF are reciprocal to each other. PRF is the number of pulses that occur in 1 second. This parameter is not related to the frequency of ultrasound. PRF can be altered by changing the depth of imaging. It is measured in Hertz (Hz). PRF = 77,000 / depth of view (cm). As evident from the equation, as the location of the target gets further away, the PRF decreases. PRF is related to frame rate or sampling rate of the ultrasound.

I would like to talk about Duty Factor (DF) here. This parameter is related to ultrasound bioeffects, but since it is also related to pulsed ultrasound it is reasonable to introduce it in this section. DF is defined as a percent of time that the ultrasound system is on while transmitting a pulse. DF = pulse duration (sec) / pulse repetition period (sec) x 100. It has units of % and ranges from 0 (the system is off) to 100 (the system is on continuously). Typical valued of DF in clinical imaging are 0.1% to 1% (usually closer to 0), thus the machine is mostly listening during clinical imaging. Another interesting point to note is the fact that since the sonographer changes the PRF by changing the depth, they indirectly change the duty factor. And lastly, one must realize that an anatomic image cannot be created with a continuous wave ultrasound. Since one must listen for the return signal to make an image, a clinical echo machine must use pulsed signal with DF between 0.1 and 1%.

Back to propertied of pulsed ultrasound, we need to discuss spatial pulse length. Up to now we introduced properties that were related to timing. Spatial Pulse Length is the distance that the pulse occupies in space, from the beginning of one pulse till the end of that same pulse. It is measured in units of distance with typical values from 0.1 to 1 mm. SPL (mm) = # cycles x wavelength (mm). Axial or longitudinal resolution (image quality) is related to SPL. Axial resolution = SPL/2 = (# cycles x wavelength)/2.

We will now talk about interaction of ultrasound with tissue. As we discussed in the section of amplitude, the energy of ultrasound decreases (attenuation) as it travels through tissue. The stronger the initial intensity or amplitude of the beam, the faster it attenuates. electricity austin Standard instrument output is ~ 65 dB. So for a 10 MHz transducer, the maximum penetration would be as follows: 1 dB/cm/MHz x 10 MHz x (2 x max depth) = 65 dB. Max depth = 65/20 = 3.25 cm. If we use a 3.5 MHz transducer and apply the same formula for max depth, will get Max depth = 65/7 = 9.3 cm. Attenuation of ultrasound in soft tissue depends on the initial frequency of the ultrasound and the distance it has to travel. As we saw in the example above, in soft tissue the greater the frequency the higher is the attenuation. So we can image deeper with lower frequency transducer. The further into the tissue the ultrasound travels, the higher the attenuation is, so it is ultimately the limiting factor as to how deep we can image clinically relevant structures.

Temporal resolution implies how fast the frame rate is. FR = 77000/(# cycles/sector x depth). Thus frame rate is limited by the frequency of ultrasound and the imaging depth. The larger the depth, the slower the FR is and worse temporal resolution. The higher the frequency is, the higher is the FR and the temporal resolution improves. Sonographer can do several things to improve the temporal resolution: images at shallow depth, decrease the #cycles by using multifocusing, decrease the sector size, lower the line density. electricity news philippines However one can realize quickly that some of these manipulations will degrade image quality. And this is in fact correct: improving temporal resolution often degrades image quality. M-mode is still the highest temporal resolution modality within ultrasound imaging to date.

Before we talk about Doppler Effect, let us discuss the ultrasound transducer architecture and function. The current transducers became available after the discovery that some materials can change shape very quickly or vibrate with the application of direct current. As important is the fact that these materials can in turn produce electricity as they change shape from an external energy input (i.e., from the reflected ultrasound beam). This effect of vibration form an application of alternative current is called a piezoelectric effect (PZT).

Many materials exist in nature that exhibit piezoelectric effect. Ccommercial transducers employ ceramics like barium titanate or lead zirconate titanate. The transducer usually consists of many PZT crystals that are arranged next to each other and are connected electronically. The frequency of the transducer depends on the thickness of these crystals, in medical imaging it ranges 2-8 MHz. An ultrasound pulse is created by applying alternative current to these crystals for a short time period. Afterwards, the system “listens” and generates voltage from the crystal vibrations that come from the returning ultrasound. electricity production in india An important part of the transducer is the backing material that is placed behind the PZT, it is designed to maximally shorten the time the PZT crystal vibrates after the current input is gone also known as ringing response. By decreasing the ringdown time, one decreases the pulse length and improves the axial resolution. In addition, the backing material decreases the amount of ultrasound energy that is directed backwards and laterally.

Image production is a complex process. Echo instrumentation must generate and transmit the ultrasound and receive the data. Then the data needs to be amplified, filtered and processed. Eventually the final result needs to be displayed for the clinician to view the ultrasound information. As the first step in data processing, the returning ultrasound signals need to be converted to voltage. Since their amplitude is usually low, they need to be amplified. The ultrasound signal usually is out of phase so it needs to be realigned in time. At this point one has the raw frequency (RF) data, which is usually high frequency with larger variability in amplitudes and it has background noise. The next step is filtering and mathematical manipulations (logarithmic compression, etc) to render this data for further processing. At this stage one has sinusoidal data in polar coordinates with distance and an angle attached to each data point. This information needs to be converted to Cartesian coordinate data using fast Fourier transform functions. Once at this stage, the ultrasound data can be converted to analog signal for video display and interpretation.

Image display has evolved substantially in clinical ultrasound. Currently, 2D and real time 3D display of ultrasound date is utilized. Without going into complexities of physics that are involved in translating RF data into what we see every day when one reads echo, the following section will provide the basic knowledge of image display. gas works park If one can imagine a rod that is imaged and displayed on an oscilloscope, it would look like a bright spot. Displaying it as a function of amplitude (how high is the return signal) is called A-mode. If one converts the amplitude signal into brightness (the higher the amplitude the brighter the dot is), then this imaging display is called B-mode.

There are several parameters that make second harmonic imaging preferential. Since it is produced by the tissue, the deeper the target the more second harmonic frequency is returned. As the ultrasound beam travels through tissue, new frequencies appear that can be interrogated. Second harmonic data gets less distortion, thus it produces better picture. Also, the second harmonic is strongest in the center of the beam, thus it has less side lobe artifacts. At the chest wall the fundamental frequency gets the worst hit due to issues that we have discussed (reflection, attenuation) – if one can eliminate the fundamental frequency data then these artifacts will not be processed. One concept of eliminating fundamental frequency data is called pulse inversion technology. The transducer sends out 2 fundamental frequency pulses of the same amplitude but of different phase. As these pulses are reflected back to the transducer, because of the different phase they cancel each other out (destructive interference) and what is left is the second harmonic frequency data which is selectively amplified and used to generate an image.

Continuous wave (CW) Doppler required 2 separate crystals, one that constantly transmits, and one that constantly receives data. There is no damping using this mode of imaging. One can measure very high velocities (i.e., velocities of aortic stenosis or mitral regurgitation). The advantage of CW is high sensitivity and ease of detecting very small Doppler shifts. The disadvantage of CW is the fact that echos arise from the entire length of the beam and they overlap between transmit and receive beams. Thus one cannot determine where in the body the highest velocity is coming from – range ambiguity.

Pulsed wave (PW) Doppler requires only one crystal. It alternates between transmitting and receiving data. The transducer “listens” for the data at a certain time only, since the sampling volume is coming from the location that is selected by the sonographer (i.e., the velocity at the LVOT or at the tips of the mitral valve). This is called range resolution. The major disadvantage of PW Doppler is aliasing. In PW mode, the transducer has to sample a certain frequency at least twice to resolve it with certainty. This put a limit on the max velocity that it can resolve with accuracy. 2 x Doppler frequency (Nyquist) = PRF. If the velocity is greater than the sampling rate / 2, aliasing is produced. The following maneuvers can be performed to eliminate aliasing: change the Nyquist limit (change the scale), select a lower frequency transducer, select a view with a shallower sample volume.

Color Flow Doppler uses pulsed Doppler technique. bp gas locations The velocity data is encoded in color, and it reports mean velocities. Since it is a pulsed Doppler technique, it is subject to range resolution and aliasing. Color data is extremely complex and consumes significant computational resources, thus several assumptions are made to speed up this process. Returned echo frequencies are compared to a predetermined threshold to decide whether this is a 2D image vs Doppler shift. Once the computer decides that the frequency is low enough to be a Doppler shift data, repetitive sampling determines the mean velocity and variance. Then a color is assigned using a color look-up table rather than doing a discrete Fourier transform for each data point. Velocities that move toward the transducer are encoded in red, velocities that move away are encoded in blue. One must remember that the color jets on echo are not equal to the regurgitant flow for a number of reasons. The regurgitant flow is a three dimensional structure with jet momentum being the primary determinant of jet size. This parameter is effected by the jet velocity as well as flow rate. Blood pressure will affect the velocity and thus the regurgitant flow. Chamber constraints will have an effect on the appearance of the color jet, especially eccentric jets. Lastly, the settings of the echo machine will have an effect on how the color flow jet appears on the screen.