Ultrasound Principles

A form of mechanical energy transmitted as a pressure wave through a given medium by the vibration of molecules. It is transmitted as a longitudinal wave, made up of molecules oscillating in the direction of propagation. Longitudinal waves create compressions and rarefactions within the air. The alternating configuration of compressions (crests) and rarefactions (troughs) can be plotted on a graph as a sine wave.

Sound wave is a longitudinal wave where the particles vibrate in the direction of propagation. As sound waves propagate, alternate regions of compression (high pressure) and rarefaction (low pressure) are created. It is often easier to visualise sound wave as a sinusoidal wave with the peaks and troughs representing regions of compression and rarefaction, respectively.

High-frequency sound waves greater than 20 kHz. Medical ultrasound uses waves of very high frequency, in the range of 2-18 MHz. The human audible range is 20-20 000 Hz.

The frequency of diagnostic ultrasound is typically in the range of MHz. For echocardiography, the frequency is 2–4 MHz.

There are several important characterising properties of a sound wave. Amplitude (A): the maximum disturbance of the medium from the equilibration point; represents the energy of a sound wave; units mm or m. Frequency (f): the number of cycles completed per second; units Hertz (Hz), i.e. 1/s. Wavelength (λ): length of space over which one cycle occurs; units mm or m. Period (T): the time required to complete a single cycle; units microsecond or second. Velocity (v): the rate that sound travels through a medium; units metres per second (m/s) or mm/µs.

Sound waves are characterised by: Wavelength (λ), the length of one cycle; Frequency (f), the number of cycles (vibrations) per second [unit: Hertz (Hz)]; Velocity (c), the speed of the sound in a particular medium and depends on the 'stiffness' (B) and density (ρ) of the material; Amplitude (A), which is proportional to the number of particles displaced and is related to loudness in sound. The relationship between f, λ and c is c = f × λ. While λ and c change with the medium's density and 'stiffness', f remains constant regardless. Amplitude tends to decrease with the distance the wave travelled due to the dissipation of energy.

Power: the energy transported per unit time by the oscillations of a particular wave; units Watt (W). The power is the sum of the kinetic and potential energy in a wave; generally proportional to the amplitude squared and frequency squared. Intensity: the power of a wave present per unit area; units W/m². Intensity varies throughout a diagnostic ultrasound beam; it is highest at the centre of the beam, tapering off at the periphery; it also varies along the path of the beam due to attenuation and focus.

Frequency and wavelength are inversely related. A short wavelength has more cycles per second and hence is a higher frequency.

Velocity (propagation speed) is the speed of sound through a given medium. It is determined solely by the characteristics of the medium: it does not depend on the source of sound or its frequency; the density and compressibility of a medium are the key determinants.

The velocity of ultrasound depends on two key properties of the medium: Stiffness — directly proportional (↑ stiffness ↑ sound velocity); Density — indirectly proportional (↑ density ↓ sound velocity). Stiffness has the greater influence overall. Sound is generally slowest in gases, faster in liquids, fastest in solids.

Propagation speed of ultrasound would most likely be maximum in a low density, high stiffness material. The speed of sound is directly proportional to the stiffness of a medium and inversely proportional to its density. Interestingly, sound often travels faster in denser materials because they also tend to be stiffer. However, variations in stiffness between materials generally have a greater effect on sound speed than differences in density.

Each body tissue has a unique velocity determined by individual characteristics. Machines commonly use an average value of 1,540 m/sec for soft tissue in depth calculations. Air 331 m/s; Fat 1,450 m/s; Soft Tissue (average) 1,540 m/s; Muscle 1,585 m/s.

The average velocity of ultrasound in soft tissue is about 1540 m/s. The speed of sound is a factor of the physical characteristics of the medium in which it is traveling. In general, velocity is faster in denser tissues. For example, in air sound travels at 345 m/s whereas it is much faster in water (1430 m/s), and up to 4080 m/s through bone. Calculate wavelength in the heart using a 4 MHz transducer: ultrasound velocity, frequency and wavelength are inter-related as λ = c/f. Using a 4 MHz transducer, wavelength would be 1.54/4 = 0.385 mm.

Acoustic impedance is defined as the resistance to propagation of sound waves. It is a physical property of the medium which varies according to its density. Acoustic impedance (Z) is a physical property of the medium which can be calculated by the product of its density (ρ) and acoustic velocity (V): Z = ρV.

Reflections of ultrasound, commonly known as echo, occur where there are changes in tissue acoustic impedance (Z), that is, impedance mismatch. As Z is proportional to tissue density (ρ), reflections occur at tissue interfaces (boundaries) where there are differences in ρ. The percentage of ultrasound reflected (R%) when travelling from tissue 1 to tissue 2 is R% = [(Z2 − Z1)/(Z2 + Z1)]² × 100, where Z1 and Z2 are the acoustic impedances for two adjacent tissues. R% only depends on the difference of Zs between the two tissues and not on the direction of ultrasound. The percentage of ultrasound left for penetration is (1 − R%).

An effect exhibited by certain quartz crystals (such as lead zirconate titanates) known as 'piezoelectric crystals'. The piezoelectric crystal will change shape when an electrical potential is applied across it: dimensions change slightly depending upon the polarity of the applied voltage; change in shape causes local changes in pressure, leading to mechanical waves (ultrasonic radiation) being generated at the exact frequency of the applied voltage. The converse effect occurs when pressure is applied to a piezoelectric crystal: compressive stress causes a change in dimensions of the crystal; generates an output voltage at the frequency of the compressive waves. Allows piezoelectric crystals to act as transducers: direct effect converts mechanical energy into electrical energy; converse effect converts electrical energy into mechanical energy.

The piezoelectric effect is the property that allows ultrasound technology to convert electrical energy into mechanical energy. In the ultrasound system, the sound source is a piezoelectric crystal, such as quartz. The piezoelectric effect allows for crystals to vibrate when an electrical voltage is applied across it and subsequently creates sound waves. Conversely, piezoelectric crystals also can convert sound waves back into electrical energy. In the absence of media (i.e., a vacuum), sound cannot propagate.

Ultrasound waves are produced by an ultrasound probe. An electric current is passed through a piezoelectric crystal: act as transducers converting electrical energy into mechanical vibrations at high frequencies; generates ultrasound waves at the frequency of the voltage applied. Ultrasound waves travel through a given medium. When a structure is encountered some waves are reflected back towards the probe. The probe detects returning waves: the pressure effect distorts the piezoelectric crystals producing an output voltage; converts the mechanical energy into electrical energy.

Output signals from the probe are then converted into an image: the time between the waves being sent out and returning is calculated; depth of a structure is determined by the time taken for a wave to return; pixels are created at the appropriate depth for returning waves; the brightness of the pixel correlates with the strength of returning wave. Multiple crystals are located within a probe: positioned adjacent to each other in an 'array'; connected electrically to generate a 2D image. As the ultrasound probe both emits and receives the signal they are known as 'transceivers'.

The transducer alternates through phases of generating and receiving ultrasound waves. The 'Transmission Phase' is the period of wave generation: usually very brief (0.5–3 µs). Followed by a 'Receiver Phase' for wave detection: much longer (up to 1 ms) than the transmission phase; allows echoes from a range of depths to be detected. The combined durations of the transmission and receiver phases is the pulse repetition period: determines the frame rate of an ultrasound generated image. Shallower depths allows for a shortened receiver phase and thus a higher frame rate.

In 2D imaging, ultrasound is transmitted in pulses, each consists of a defined number of cycles. The length of a pulse can be measured by the time needed to produce the pulse; this time is known as the pulse duration and is equivalent to transmission time. High frequency requires less time to produce one pulse; hence it has shorter pulse duration. After a pulse is transmitted, the transducer spends the rest of the time listening for echoes (listening time) until the next pulse is sent. The frequency by which the pulses are sent is known as pulse repetition frequency (PRF) and has nothing to do with ultrasound frequency. Frequency 2–20 MHz; pulse duration typical 1–2 µs; pulse repetition period typical 250–500 µs; pulse repetition frequency = 1/PRP, typical 2000–4000 Hz.

There are 3 possibilities for an ultrasound wave that enters tissue. Transmission: wave and its energy pass through the tissue in its propagation direction. Attenuation: wave loses energy due to interaction with tissue. Reflection: wave changes direction and travels back towards the probe. Ultrasound needs to be transmitted far enough into the tissues in order to image them but must be reflected back to be received by the probe.

When an ultrasound pulse encounters a reflector, three things can happen: Scattering, Reflection, Transmission. Scattering: when the size of a reflector is smaller than λ (e.g. RBC), the ultrasound pulse is scattered in all directions; as a result, only a very small portion of echoes return to the transducer; scattering gives rise to the characteristic acoustic appearance of the tissue. Transmission: ultrasound that is not reflected will continue to travel across the tissue boundary; tissues of similar densities favour transmission than reflection.

Reflection occurs when waves encounter an interface or boundary between materials with different acoustic impedances. The difference in impedance between mediums is known as the reflection coefficient. The greater the reflection coefficient, the greater the size of the wave (and percentage of energy) that is reflected at the interface. For example, a large difference in impedance is present between air and tissue: therefore, almost all of the energy is reflected back towards the probe; almost no energy is transmitted through the air to allow image generation for deeper structures; this explains why it is vital to use ultrasound gel when scanning to eliminate air tissue interfaces. The angle of reflection of a wave equals the angle of incidence: when a wave strikes a tissue interface at 90 degrees a very strong reflection travels back toward the ultrasound probe generating a bright image; waves striking tissue at other angles are reflected away from the probe leading to loss of wave energy.

Reflection occurs at the interface of tissues where the densities are different. The larger is the difference, the more ultrasound is reflected. Near total reflection occurs at air-tissue interface explaining why lung, with multiple air-tissue interfaces, is a poor ultrasound conductor.

Attenuation can be defined as the net energy loss from the interaction of tissues with the ultrasound beam. This results in the amplitude and intensity of ultrasound waves decreasing as they travel through tissue. Attenuation affects high frequency ultrasound waves to a greater degree than lower frequency waves: lower frequency transducers are used for scanning deeper structures. Due to 4 key mechanisms: 1. Reflection — waves may be reflected away from the transducer when they encounter an interface at an angle other than 90°. 2. Refraction — describes the bending of the ultrasound beam when a beam strikes an interface at an angle other than 90° causing it to deviate (analogous to a drinking straw bending when resting in a glass of water). 3. Scatter — describes the dispersal of a beam when it strikes an interface that is the same size or smaller than its wavelength. 4. Absorption — waves are absorbed as heat, reducing the energy of the beam; degree of absorption is directly proportional to frequency.

Significant attenuation through bone and air explain why ultrasound is poor at visualising these tissues. The degree of attenuation through muscle explains why transoesophageal echocardiography can produce superior images to transthoracic echocardiography. Attenuation means that weaker signals are received from deeper structures as compared to identical structures positioned more superficially. Time-gain compensation (TGC) is an adjustment that ultrasound machines perform to differentially amplify signals depending on the depth they were received from.

Attenuation (Echocardiography in ICU): Ultrasound energy, hence amplitude, gradually diminishes as the ultrasound pulse penetrates deeper into the tissue. The loss of energy can be due to three main reasons: 1. Scattering where energy is reflected in all directions; 2. Overcoming tissue viscosity (stiffness or resistance) where energy is lost as heat; 3. Reflections hence less energy available for penetration. How to improve brightness at deeper structures: 1. Increase the far-field gain using time gain compensation (TGC). 2. Position the focal point at the far field. 3. Lower the frequency to minimise energy loss and to increase penetration. 4. Activate tissue harmonic imaging (THI).

Resolution can be defined as the detail of the image determined by the ability to distinguish two points as separate in space. Two components of resolution can be considered: Axial — the ability of a probe to distinguish between two structures along the beam length. Lateral — the ability to distinguish between two adjacent objects side by side in the ultrasound beam.

'Spatial pulse length' is the product of the wavelength and the number of cycles in a pulse. A shorter wavelength (and thus higher frequency) results in a shorter spatial pulse length. Resolution is determined by the 'spatial pulse length' of the wave: the minimum distance that can be reflected between two points is equal to half the 'spatial pulse length'; therefore resolution is high when the 'spatial pulse length' is short and the wave frequency is high. Higher frequency waves have greater resolution but at the expense of deeper penetration.

Axial resolution is determined as one half the Spatial Pulse Length (SPL/2). Spatial pulse length is the product of wavelength and number of cycles in the pulse; the shorter the spatial pulse length, the better the axial resolution. It is dependent on transducer frequency, transducer bandwidth and pulse length (preset by the manufacturer); better with higher frequencies and shorter pulses; independent of imaging depth, and so is fixed throughout the image. Lateral resolution is determined by beam width; narrower beams have superior lateral resolution; lateral resolution (also called angular, transverse or azimuthal resolution) is best at the preset focal length of the transducer. M-mode imaging has the greatest favorable effect on temporal resolution because the transducer only sends and receives along a single scan line. Good temporal resolution is important for distinguishing moving structures. (Resolution in ultrasound imaging — Ng & Swanevelder: spatial resolution (axial + lateral), temporal resolution, and contrast resolution.)

Beam Focusing and Lateral Resolution (Echocardiography in ICU): Electronic focusing converges the beam toward a focal zone, which has the narrowest beam width and the highest beam intensity. To resolve two closely placed reflectors, the beam width needs to be narrower than the distance between them, such that each reflector sends an echo back at different times. If two reflectors are not lying in the focal zone, lateral resolution is compromised. Axial resolution depends on the pulse duration (or pulse length); if the PD is long, the second echo may catch up with the first and merge into one; if the PD is short, the two echoes are separate. With the same number of cycles per pulse, high frequency results in shorter pulse duration.

Increasing imaging depth → increases PRP (sound takes longer to return). Increasing imaging depth → decreases PRF (fewer pulses per second). Unaffected by depth (inherent properties of the transducer and remain constant regardless of imaging depth): Pulse duration (PD); Amplitude; Spatial pulse length (SPL).

Increasing the depth of interrogation from 10 to 20 cm during transthoracic echocardiography will most likely cause a decrease in Pulse Repetition Frequency. Increasing imaging depth lengthens the pulse repetition period (PRP), which in turn reduces the pulse repetition frequency (PRF). Parameters such as pulse duration, amplitude, and spatial pulse length are intrinsic to the transducer and are not affected by imaging depth.

The absorption of an ultrasound wave is proportional to the frequency. When more energy is absorbed the wave is less able to penetrate deep into tissues. Therefore lower frequency waves penetrate deeper but at the expense of resolution. Higher-frequency probes provide less penetration of the US waves through the tissue planes but generate higher-resolution images; high-frequency probes (linear, intracavitary) are used to visualise superficial structures, while lower-frequency probes (curvilinear) are used to visualise deeper structures at the expense of resolution.

Artefact can be defined as an appearance that does not accurately correspond to the true image of the examined area. Artifacts can often be recognized by altering the image plane, depth, or frequency. Unusual object should be viewed from multiple directions to ensure that it is anatomic rather than artefactual.

Types of ultrasound artefacts. Acoustic Artifact — Axial Direction: Reverberation, Comet Tail, Mirror Image, Acoustic Enhancement, Acoustic Shadowing, Lateral Shadowing. Acoustic Artifact — Lateral Direction: Side lobe, Beam width. Interference Artefact — External Equipment: Unshielded Electrical Equipment, Aliasing, Click; Devices: Pacemakers, Implants. Technical Artefact — Scanning Technique: Contact / Interface; Interpretation: Anatomic / Pitfall error.

Reverberation: the ultrasound beam is reflected between two highly reflective interfaces; causes the same interface to be represented multiple times in the image at equally spaced positions; each representation is weaker and deeper due to the greater time it takes for the wave to return and the energy lost. Mirror Image: occurs due to the interaction of waves with a strong reflector; gives the appearance of two images — one real and one artifact; the artefactual image always appears deeper than the true anatomy and is an equal distance on the other side from the strong reflector. Acoustic (post-cystic) enhancement: occurs due to waves passing more easily through fluid filled structures with only a minimal amount of the beam attenuated; this leaves greater wave energy to be reflected after the fluid filled structure which appears brighter in the image due to time gain compensation.

Acoustic shadow: occurs when a wave is unable to pass through a structure that is deep to a highly reflective surface; leaves a dark shadow behind where no waves are being reflected back to the transducer. Lateral shadowing: due to refraction at the curved edge of a structure; causes the beam to be directed away from the probe at the edges and are therefore 'lost' from the image. Side Lobe: the beam contains waves fanning out from the apparently infinitely thin central beam which forms the main image; if a strong reflector is encountered in these side beams they can be reflected back and interpreted as part of the main image. Beam Width Artefact: the beam is not equally wide along its course, having a narrower focal zone; artefact occurs when two structures in a wider part of the screen are displayed as overlapping when the machine displays reflections.

US artifacts (Feldman, Katyal & Blackwood, RadioGraphics 2009): the beam width, side lobe, reverberation, comet tail, ring-down, mirror image, speed displacement, refraction, attenuation, shadowing, and increased through-transmission artifacts are encountered routinely in clinical practice; some artifacts are avoidable and due to improper scanning technique, while others are generated by the physical limitations of the modality.

In M-mode imaging the transmit time/receiver cycle is very rapid because the transducer only sends and receives an ultrasound signal along a single scan line; this optimizes the temporal resolution of the image and allows evaluation of rapidly moving structures. The frame rate in M-mode is much higher than in 2D or 3D because M-mode does not sweep through the imaging sector. Early closure of the aortic valve in hypertrophic cardiomyopathy is best demonstrated by M-mode.

Probes can be characterized according to their: shape; arrangement of the piezoelectric crystals; frequency range; footprint.

Phased array probes use the concept of constructive interference to 'steer' soundwaves to focus on specified regions of interest; they feature a smaller curved surface which emits diverging soundwaves producing a field of view wider than the footprint of the probe itself, ideal for maneuvering between ribs; probe frequencies span 2–8 MHz allowing use for both higher-resolution cardiac imaging and deeper structures. Curvilinear array probes have a curved surface with crystals arranged in a curved fashion resulting in a diverging sound wave propagation; this diverging feature results in a field of view which is wider than the footprint of the probe itself.

The probe can be manipulated in a number of ways to optimise the image. Can be considered using the mnemonic PART (Pressure, Alignment, Rotation and Tilt).

Pressure: improves the image by reducing the distance to target structures through compression of subcutaneous tissues; can be used to compress veins or move structures out of the way; care needs to be taken not to use excessive pressure which causes discomfort for the patient. Alignment: refers to the movement of the probe on the skin to move the target structure into the field of view; sliding the transducer through different alignments can be useful in tracing structures along their course to help with identification. Rotation: refers to the twisting of the probe to fine tune the view of a target structure; allows the true axial view of structures to be obtained; rotation through 90° can change an image from a long axis to a short axis view. Tilt: assists in bringing the face of the probe, and thus the direction of the ultrasound beam, into a perpendicular arrangement with the target structure; this maximises waves reflected back to the probe and thus provides a better image. (Rock also described.)

'Rocking' the ultrasound probe is used to improve the image of structures: this is known as anisotropy. The sciatic nerve is highly anisotropic and will be bright at a 90° angle of insonation but virtually disappear at 80° or 100°.

Dynamic range (Echocardiography in ICU): A 2D ultrasound image is represented by shades of gray. Dynamic range (DR) describes how the original range of gray scale is displayed. If the range is 'compressed' (reduced DR), the structure with lighter gray and darker gray will appear as white and dark, respectively; this results in an image with higher contrast. A wider DR image appears 'softer'. Narrow DR: eliminates low-level background noise but enhances the cardiac structures (increases contrast); good for border detection. Wide DR: brings out the weaker signals; results in softer images; best for detecting structures with little echo variation, such as thrombus, vegetation and tumor. Some machines use the term 'compression' instead of 'dynamic range'. Changes to frame rate: narrowing sector width, decreasing image depth and decreasing the number of focal zones increase the frame rate and temporal resolution.

The doppler effect applies to all waves, both sound and electromagnetic radiation. It changes the frequency of reflected ultrasound waves (not velocity), shows a shift to a lower frequency if the source is moving away from the receiver and can indicate the velocity of the red blood cells. Doppler imaging is a technique that uses ultrasound waves to detect the motion of blood or tissue; this can be displayed as either colour or sound.

The detected Doppler shift frequency (fd) can be used to calculate the velocity using the 'doppler equation': v = (c · fd) / (2 · fT · cosθ), where v = flow velocity (m/s), c = the speed of sound in tissues (m/s), fd = Doppler frequency shift that is received (Hz), cosθ = cosine of the angle between the sound beam and moving fluid, fT = frequency of the transmitted ultrasound from the transducer (Hz). Types of doppler imaging: Spectral Doppler — velocity information from a single location displayed as a frequency shift–time plot. Colour Doppler — displays the direction and magnitude of flow as a colour image; Blue indicates flow Away from the ultrasound probe and Red indicates flow Towards the probe (BART). Duplex Doppler — combines real time colour doppler superimposed onto real-time grey-scale b-mode image. (Full detail on the Doppler Principles page.)

Image Guided Procedures: vascular access; peripheral nerve blockade; pleural / peritoneal / pericardial aspiration or drainage; assessment prior to neuraxial blockade; assessment of cricothyroid membrane prior to cannulation. Diagnostic Imaging: echocardiography; 'Point of Care Ultrasound' — lung imaging, abdominal imaging & detection of portal venous flow, trauma imaging (FAST Scan), detection of DVT. Monitoring: oesophageal doppler cardiac output monitoring; transcranial doppler assessment of cerebral blood flow; foetal heart rate monitoring. Other: cleaning of equipment; humidification devices; therapy for acute and chronic pain.

Advantages: relatively inexpensive; widely available; can be performed by the bedside; non-invasive; non-ionising; safe in children and pregnancy.

Disadvantages: operator dependent; cannot image lung or bone; not good for imaging obese patients or deep structures; specific issues with oesophageal probes — must be avoided in patients with strictures or tumours, carry risk of bleeding (particularly in varices) / perforation.

Bioeffects refer to the biological impact ultrasound pressure waves have on living tissues. Classically, these effects are categorized as either thermal or mechanical (non-thermal) in nature. Mechanical (MI) and Thermal (TI) indices are frequency-dependent calculated values which indicate the likelihood of clinically relevant bioeffects given the current ultrasound probe settings. MI or TI values <0.5 are unlikely to exert significant bioeffects, while values >1–1.5 indicate likely bioeffects and require careful attention. As part of the FDA's Output Display Standard (1992), all ultrasound machines must display the Mechanical Index (MI) when operating in B-mode and the Thermal index (TI) when performing adult orbital/cephalic or obstetrical ultrasound examinations.

The Mechanical Index (MI) is a measure of the power of an ultrasound beam and determines possibility of non-thermal bioeffects of the acoustic field, such as cavitation. Cavitation is the expansion and contraction or collapse of bubbles because of the acoustic pressure of the ultrasound beam. MI is defined as peak negative pressure divided by the square root of the frequency of the ultrasound wave. According to FDA recommendations, bioeffects of ultrasound are best avoided when the Mechanical Index is less than 1.9. Lowest thermal index can be achieved by choosing lowest frequency and lowest intensity (low intensity results in low energy transmission into the body).

Tissue harmonic imaging (THI) uses echoes of higher frequencies than the fundamental wave to generate images. These harmonic frequencies occur because the ultrasound wave propagates faster through compressed regions of tissues. Because these harmonic frequencies only have to travel one-way back to the transducer, the resulting images are clearer; harmonic imaging improves endocardial definition and decreases near field and side lobe artifacts. However, axial resolution is reduced in harmonic imaging, and there is lower penetration and increased acoustic shadowing. (ALARA: keep output As Low As Reasonably Achievable; raising receiver gain does not increase output power.)