Hydrophone Measurements for Biomedical Ultrasound Applications: A Review

Abstract

This article presents basic principles of hydrophone measurements, including mechanisms of action for various hydrophone designs, sensitivity and directivity calibration procedures, practical considerations for performing measurements, signal processing methods to correct for frequency-dependent sensitivity and spatial averaging across the hydrophone sensitive element, uncertainty in hydrophone measurements, special considerations for high-intensity therapeutic ultrasound, and advice for choosing an appropriate hydrophone for a particular measurement task. Recommendations are made for information to be included in hydrophone measurement reporting.

I. Introduction

The field of biomedical acoustics encompasses a wide array of diagnostic, therapeutic, and surgical applications [1]. Diagnostic uses include imaging the fetus, measuring blood flow, visualizing organs in 2-D or 3-D, and characterizing mechanical properties of tissue. Therapeutic and surgical uses include physiotherapy (mild hyperthermia), lithotripsy (comminution of kidney stones), transcranial neuromodulation, and thermal or mechanical ablation of tumors.

In all these applications it is important that the acoustic output be characterized accurately, not only to assess the safety and effectiveness of the medical procedures but also to meet the requirements of regulatory bodies. For example, before the US Food and Drug Administration clears a new diagnostic ultrasound device, acoustic output data must show substantial equivalence to a legally-marketed device [2]. Miniature ultrasonic hydrophones (that convert incident pressure into voltage) historically have played a vital role in these measurements.

International standards and technical specifications describe relevant acoustic measurements to characterize medical ultrasound devices (e.g., [3–6]). These measurements include the instantaneous pressure waveform at a particular point in the field, p(t) (see Fig. 1); peak compressional and peak rarefactional pressures, p_c and p_r (shown in Fig. 1), root-mean-square (rms) pressure, p_rms; and the pulse-pressure-squared integral, ppsi, which is the integral of p²(t) over the pulse waveform. (Note: Waveforms can differ from the diagnostic imaging pulse depicted in Fig. 1. For example, lithotripsy waveforms have much faster rise times, therapeutic pulses can have hundreds of cycles, etc.)

Fig. 1.

Typical diagnostic ultrasound imaging pulse (normalized) showing peak compressional pressure (p_c) and peak rarefactional pressure (p_r).

Additional quantities that can be computed from these hydrophone measurements include pulse duration, acoustic-working frequency, spatial-peak temporal-average intensity, I_spta, and spatial-peak pulse-average intensity, I_sppa.

Regarding spatial quantities, hydrophone scans of the ultrasound field provide information such as focal length, depth of focus, and beam cross-sectional dimensions at various points along the beam axis. From these scans, the spatial-average temporal-average intensity, I_sata, can be derived.

These quantities have been used to evaluate the biological effects and safety of ultrasound in two general categories: thermal and nonthermal. The temporal-average and rms quantities have been associated with thermal effects, and the temporal-peak and pulse-average quantities with nonthermal effects. However, because ultrasound-induced bioeffects are complex functions of the acoustic field and exposed tissue, these quantities do not provide a complete description of exposure. Therefore, additional bioeffects indices based on hydrophone measurements have been developed: the thermal index (TI) and the mechanical index (MI) [7, 8]. These indices are displayed on most modern diagnostic ultrasound imaging devices [9].

The purpose of this review article is to describe various hydrophone designs for biomedical applications. From the International Electrotechnical Commission’s definition of a hydrophone, a “transducer that produces electrical signals in response to pressure fluctuations in water” [10], several transduction mechanisms have emerged. These include capacitive, in which the pressure amplitude can be calculated from ultrasonically induced changes in the gap width of a parallel-plate, air-dielectric capacitor; piezoelectric, in which the sensor material develops an electric charge in response to an applied mechanical stress; and fiber-optic, in which laser light is directed along an optical fiber, and optical reflections at the fiber tip are sensitive to acoustic variables at the fiber/fluid boundary.

Piezoelectric and fiber-optic sensor types have found the most widespread use, and their designs are discussed in the next section. Subsequent sections cover hydrophone sensitivity and directivity calibrations, measurement system basics, signal processing and measurement uncertainty, challenges for measuring high intensity pressure fields, considerations for choosing the right hydrophone for a particular application, and recommendations for reporting hydrophone measurement results. Table I provides a glossary of useful terms.

TABLE I.

Glossary

Acceptance Angle	The full-width half maximum of the directivity if less than 180° or 180° otherwise (Sec. III).
Bandwidth	Frequency range over which hydrophone sensitivity is approximately flat (e.g., ±1.5 dB) (Sec. III)
CH	Capsule hydrophone (Sec. II, Fig. 3)
Diffraction (edge) wave	Pressure wave near hydrophone is superposition of wave received by hydrophone, reflected plane wave, and diffraction (edge) wave (Sec. II).
Directivity	Hydrophone response to a plane wave, as a function of frequency and angle of incidence. At each frequency, directivity is normalized to the value at normal incidence (Sec. III, Fig. 5)
FOH	Fiber-optic hydrophone (Sec. II)
Frequency Response	See sensitivity (Sec. III).
Maximum Pressure	Maximum pressure that a hydrophone can measure while retaining a response that is linearly related to incident pressure and not suffering damage (Sec. VII).
MH	Membrane hydrophone (Sec. II, Figs. 2, 9)
MI	Mechanical Index (Sec. I)
NEP	Noise-equivalent pressure (Secs. VI, VII)
NH	Needle hydrophone (Sec. II, Fig. 4)
p c	Peak compressional pressure (Sec. I, Fig. 1)
p r	Peak rarefactional pressure (Sec. I, Fig. 1)
ppsi	Pulse-pressure-squared integral (Sec. I)
PVDF	Polyvinylidene difluoride, a piezoelectric polymer used in hydrophone sensors (Sec. II)
Sensitive (active) element	Area on hydrophone that detects incident pressure (Sec. II)
Sensitive (active) element diameter, geometrical, dg	Physical diameter of sensitive element as defined by the electrode structure (Sec. II, III, V)
Sensitive (active) element diameter, effective, deff(f)	Frequency-dependent sensitive element diameter inferred from directivity measurements (Sec. III, V, Fig. 6)
Sensitivity	Hydrophone response to a normally incident plane wave, as a function of frequency. (Sec. III)
Spatial Averaging	In the context of hydrophones, spatial averaging refers to the hydrophone being sensitive to the spatial average of the incident pressure across the surface of its effective sensitive element (Sec. V).
TI	Thermal Index (Sec. I)

Relevant sections and figures from this article are shown in parentheses.

II. Ultrasonic Hydrophone Designs

Within the field of biomedical acoustics there is a wide variety of hydrophone designs. The most common types are:

• Membrane hydrophone (MH)

• Capsule hydrophone (CH)

• Needle (probe) hydrophone (NH)

• Fiber-optic hydrophone (FOH).

All hydrophones have a response that depends on both frequency and direction of insonation. Whilst directivity is largely determined by the size and shape of the active element relative to wavelength, frequency response depends mostly upon thickness, and radially diffracted, resonances. The hydrophones listed are ordered (top to bottom) to indicate increasing influence of diffraction phenomena on their lower frequency response. Table II compares properties for the different hydrophone types.

TABLE II.

Hydrophone Properties

Type	Sensitive Element Size (tun)	NEP (kPa)	Maximum Pressure (MPa)	Bandwidth (MHz)
MH	75–1000	1–65* [23, 24, 50, 55, 127]	75* [126]	~100* [126]
CH	85–400	-	10 [102]	~30–40 [34, 66]
NH	40–4000	≤130* T50, 55, 1271	48* [117]	~20 [66, 126]
RT FOH	100	300–500* [60,127]	100* [132]	~100* [126]
I FOH	10–125	10–15 [50, 58]	37–60* [126]	~40–50* [126]

NEP: noise-equivalent pressure. RT: Reflectance-type I: Interferometric. Values are approximate and very among different individual hydrophones within each type. Values are drawn from publications, but modern hydrophones may attain superior performance using proprietary, unpublished technology. NEP values depend on metric (e.g., standard deviation vs. peak-to-peak value) and whether signal averaging is performed. For a given hydrophone type, NEP tends to be inversely related to sensitivity and sensitive element area. NEP values may include contributions to noise from typical preamplifiers. Values for maximum pressure and bandwidth correspond to demonstrated values. Higher values may be potentially achievable.

Values at high ends of ranges marked with * correspond to robust hydrophones designed specifically for use in HITU. See Sec. VII.

A. Membrane Hydrophones (MH)

The concept of an MH was originally proposed in the 1980s [11–15]. A thin film of a piezoelectric polymer (usually polyvinylidene difluoride, a.k.a. PVDF) is held taut over a supporting ring and selectively polarised in the area of the active element only. Early devices had active elements with diameters ≥ 0.5 mm [11, 14, 16, 17]. However, with a move towards higher frequency applications, MHs with active elements of 0.2 mm diameter were developed [18]. Very high frequency (> 20 MHz) applications drove a requirement for yet smaller active element diameters, and sub-40-μm MHs have been reported [19, 20].

Fig. 2 displays four membrane hydrophones. A single layer of film with electrodes on either side is sometimes termed a coplanar MH [17]. Early devices of this form had relatively low sensitivity, especially if no integral preamplifier was included, and were susceptible to electromagnetic noise. Bilaminar MHs were developed [11, 13, 17, 21] to address both these issues. Bilaminar MHs exhibit improved shielding, provided by two ground planes, resulting in a reduction of electromagnetic noise [13, 21]. The advent of broad bandwidth differential amplifiers which could exploit common mode rejection [22, 23] enabled development of coplanar devices with bandwidths up to 140 MHz. The application of absorbing backings to MHs enhances robustness and facilitates their use in quantification of high intensity therapeutic (HITU) fields [24] (see Sec. VI). Both single-layer and bilaminar MHs employing these features are now commercially available.

Fig. 2.

Membrane hydrophones. a-c: coplanar designs with differential amplifiers, d: bilaminar design. A typical membrane hydrophone frame diameter is approximately 10 cm.

MHs exhibit a broadband response that is primarily dependent on the thickness of the film. This can be modelled either with a systems-level approach [25] or with a comprehensive analytical model [26]. The directional response (which is a function of frequency and angle of incidence) decreases smoothly and monotonically for incident plane waves that have nearly normal incidence but becomes complicated at low MHz frequencies and when incidence angles are beyond the critical angle for a water/PVDF interface (37°) [14, 27]. MHs are routinely used to disseminate primary pressure standards [21,23] and rapidly became the gold-standard for quantification of output from diagnostic ultrasound equipment, with excellent long-term stability [15, 17, 28–30].

B. Capsule Hydrophones (CH)

CHs were developed to provide a physically smaller device than MHs whilst reducing some diffractive artefacts (see [31–33] and Table I) that are present with NH or FOH designs[34]. CHs incorporate a membrane of PVDF held taut at the end of an acoustically absorbing (yet impedance matched to water) polymer structure The ellipsoidal shape of these devices reduces radial diffractive modes and thus whilst their low MHz frequency response is not quite as uniform as MHs, it is much smoother than would be encountered with NH designs [34]. The flatness of response of CHs has been attributed to the reduction of radial diffraction artifacts [33] by extending the hydrophone geometric profile over a relatively large area (compared to NHs and in terms of acoustic wavelength) and smoothing-out edges [34]. CHs are available with active diameters in the range 0.085 mm to 0.4 mm. See Table II. An example of the 0.2 mm element CH is shown in Fig. 3.

Fig. 3.

Capsule hydrophone. The length is approximately 6 cm.

C. Needle (probe) Hydrophones (NH)

NHs were also developed in the early 1980s and are based upon a piezoelectric element mounted on the tip of a co-axial conductive shaft. Active elements may be piezo-polymer [35–37] or piezo-ceramic [38, 39]. The frequency response of NHs is influenced by a number of factors [32]. At low frequencies diffraction (see [31–33] and Table I) dominates; artefacts influenced by the scattering cross-section of a cylindrical rod [40] are evident in the NH frequency response. At higher frequencies, thickness resonances of the sensing element and capacitive divider effects along the co-axial shaft introduce a higher frequency roll-off. NH designs are available with sensor diameters in the range of 10s of microns to several millimetres, as shown in Fig. 4. Overall device sensitivity, location of features in the frequency response, directional response, and spatial averaging effects are all dependent upon sensor diameter/area. A large range of available devices enables the user to select the most appropriate compromise among different performance metrics. Like CHs, NHs have compact form factors, which make them more convenient than MHs for some measurements (e.g., with their ability to get closer to the radiating surface of focusing transducers). Currently, NHs are available in a wider range of sensitive element sizes than CHs. See Table II. NH designs incorporating protective metallic coatings to enhance robustness have also been developed [39,41] (see Sec. VI).

Fig. 4.

Needle hydrophones.

D. Fiber-optic Hydrophones (FOH)

There is a wide variety of FOHs incorporating extrinsic (built on to the distal tip of the optical fiber) and intrinsic (constructed within the length of the fiber) sensing mechanisms. A detailed review is found elsewhere [42]. Only the most common types of extrinsic FOH are discussed here.

In a reflectance-type FOH, laser light is transmitted along an optical fiber whose tip is immersed in a fluid where pressure is to be measured [43]. In addition to transmitting light, the FOH hardware also receives light that reflects at the fiber tip and propagates along the fiber back toward the optical source. The reflection coefficient at the fiber/fluid interface depends on the refractive indices of the fiber and the fluid. Ultrasonic waves incident upon the FOH tip cause localised changes in density of the fluid and thus its refractive index. Consequently, there are ultrasonically induced perturbations to the intensity of the reflected light detected by the FOH. These refractive index changes are small and thus reflectance-type FOHs require higher power (Class 3B) lasers and are typically used to measure high pressure fields such as from shockwave [44–46] or HITU sources [47].

Interferometric FOHs employ a structure fabricated on the tip of the sensor. This could be a Fabry-Perot interferometer with a polymeric spacer layer [48–51] or could be formed via a stack of dielectric layers [52–55]. The interferometric FOH uses a photodiode to detect the interference between reflections at the near and far boundaries of the structure at the tip. Compression of the interferometric structure by an incident ultrasonic wave leads to modulation of the light reflected from the tip; under some conditions, the reflected light intensity is proportional to instantaneous pressure. However, compression/expansion of the interferometer could also arise from thermal effects, and interferometric FOH sensors have been shown to respond to thermal change [51, 56]. Simultaneous measurement of ultrasonic pressure and ultrasonically induced heating are reported [57, 58]. Interferometric FOHs have a higher sensitivity than their reflectance-type counterparts and can measure pressure signals below 50 kPa with lower power lasers (Class 1M).

The probe-like nature of FOHs, combined with their high levels of circular symmetry leads to radial diffraction artefacts influencing their frequency [33, 54, 59] and directional response [60]. If this can be accurately modelled [33] then corrections can be applied. Alternatively, geometric modification of the FOH tip can reduce the diffraction artefact [51, 61–63]. The acousto-optic transduction mechanism is also largely immune to electromagnetic interference. Thus, FOHs are well suited to applications where separation of the acoustic and electromagnetic noise signals would otherwise be a challenge.

III. Sensitivity and Directivity Calibrations

Hydrophones must be calibrated to allow quantitative acoustic field measurements. The key quantity is the hydrophone sensitivity expressed in units of V/Pa. As illustrated in Fig. 5, the sensitivity is defined as the ratio of the output voltage to the input acoustic pressure existing at the spatial location of the device in its absence [6]. The full transfer function describing the relationship between these two quantities is typically frequency-dependent and complex valued. Full calibration for deconvolution techniques (Section V) ideally includes both the magnitude and phase response [64–68]. However, for practical purposes, voltage-to-pressure conversion has historically been carried out using the magnitude response without resorting to phase calibration. Alternatively, the minimum phase principle has been used to estimate the phase response from the magnitude response [6, 66, 68, 69].

Fig. 5.

Measurement of directional response of membrane hydrophone by rotating about two orthogonal axes.

Metrology (the science of measurement) plays a key role in establishing the framework for quantification that underpins consistency, comparability, and confidence. Whilst several hydrophone calibration techniques have been reported [70], primary methods of calibration applicable to medical ultrasound address realisation of the acoustic pascal in water. At a handful of global National Metrology Institutes (NMIs), the primary standard for realising this quantity involves application of optical interferometry to measure the acoustic displacement or velocity of a thin optically reflective and quasi-acoustically transparent membrane (thickness <5 μm) generated by a plane-progressive acoustic wave [71–74]. By placing the membrane within the far field of a transducer, the acoustic displacement (d) can be related to the acoustic pressure (p) using:

$$p=\rho c\omega d$$ (1)

where ω is the angular frequency and ρ and c are the density and speed of sound in water. The SI-derived unit of the acoustic pascal – dynamic force per unit area – is represented dimensionally as M L⁻¹ T⁻² and is linked to three of the SI base units existing prior to redefinition of the International System of Units. For an acoustic pressure amplitude of 5 MPa at 5 MHz, the displacement generated is almost 110 nm. For equal acoustic pressures, acoustic displacements are inversely proportional to frequency, leading to significantly increased calibration uncertainties at higher frequencies that arise from the challenges of measuring displacements of 10’s of picometres.

As with many areas of metrology, global consistency in realising important units must be verified internationally. Carried out through practical comparison using a strict protocol, for the acoustic pascal in water, this involves circulating stable hydrophones to several NMIs and systematically comparing results. The last Key Comparison covered the frequency range 0.5 to 20 MHz, with a typical primary calibration uncertainty over the four participating NMIs (at 5 MHz) being ±3% (95% confidence level) for an ideal case of bilaminar membrane hydrophones with 1 mm active element diameter, high sensitivity, and low noise [75].

Relative or comparative hydrophone calibration is less costly, less time-consuming, and more accessible to the end user. However, these benefits come at the cost of greater calibration uncertainty (see next paragraph). Relative calibration is performed by comparing the response of an unknown hydrophone with a second device whose calibration is traceable to national standards. The applied acoustic field can be adjusted to customize the calibration. This relates to the broadness of the acoustic field (to reduce the influence of spatial-averaging effects [76]) and extends to the type of acoustic excitation applied (continuous, tone-burst or pulsed). Temporal characteristics of the applied waveform affect the frequency range of the calibration and the frequency resolution. Calibration over a broad frequency range is required to define the hydrophone bandwidth (Table I). Importantly, requirements imposed on the relative calibration systems are much less exacting than for primary level calibration and mainly involve ensuring that the two hydrophones are positioned for maximum signal at exactly the same spatial locations in the acoustic field. A comparison of primary and relative calibration techniques applying different excitation schemes has been reported [77].

Relative calibration uncertainty includes uncertainties at both primary and secondary stages. In one investigation, relative calibration uncertainties quoted by an NMI were 8% (1–16 MHz), 12% (17–30 MHz), 15% (31–40 MHz), 18% (41–50 MHz), and 22% (51–60 MHz) for five membrane hydrophones with sensitive element diameters ranging from 200–1000 μm [78]. Uncertainties for both primary and relative calibration may depend on hydrophone type, sensitive element size, frequency, and other parameters.

An important tenet of any calibration is to ensure that the conditions pertaining to the calibration are relevant to the end user measurement application. This can relate to calibration temperature but might also extend to the type of acoustic field being measured and the observation of any preconditioning (soaking effects) which can be important for some device types. Calibrations are generally performed in water. If the hydrophone is used in other fluids, it may be necessary to account for altered sensitivity due to different acoustic impedance matching conditions.

As Fig. 5 illustrates, the hydrophone response is typically highly directional with regard to the incoming acoustic wave. This arises from phase-cancellation across the hydrophone active element as the device is progressively rotated [79]. The directional response can be used to derive the effective diameters of the hydrophone active element [80, 81]. Effective diameters are commonly frequency-dependent (see Fig. 6) and are required to establish spatial-averaging corrections, particularly for tightly focused acoustic beams (see Sec.V). When undertaken through several axes of rotation, the directional response provides useful information on the poled region of the device as well as the degree of asymmetry [82] or additional actively poled areas [83] that might impact measurement accuracy.

Fig. 6.

Beamwidths (full width half maxima) at fundamental and harmonic frequencies for Sonic Concepts H101 Transducer (3.3 MHz, F/1) (black asterisks). Frequency-dependent effective sensitive element diameters deff (f) for five membrane hydrophones withgeometrical sensitive element diameters d_g of 200 through 1000 μm (dashed lines) are also shown. Reprinted from [118].

The acceptance angle θ_A is defined as the full-width half maximum of the directivity if less than 180° or 180° otherwise. For the product of frequency (MHz) and sensitive element geometrical diameter (mm) = [0.2 0.5 1 2 5] MHz· mm, θ_A ≈ [100° 85° 75° 40° 20°] for typical MHs [27] and θ_A ≈ [180° 180° 120° 60° 25°] for typical CHs, NHs, and reflectance-type FOHs [81, 84].

IV. Measurement Basics

While much of this review is focused on hydrophones and their characteristics, a typical measurement system also comprises: a water tank of suitable dimensions; a means to position the hydrophone in space relative to the ultrasound source; a waveform data capture device; and a control computer to orchestrate the measurement process [85–87]. The hydrophone converts the ultrasound pressure fluctuations into electrical signals which the measurement system then analyzes to determine pressure amplitude, intensity, beam patterns, total power, and other relevant field characteristics. While this section emphasizes automated systems, many of the concepts also apply to manual positioning systems, which are also useful for performing measurements.

A. Scanning Tanks

There are two basic arrangements. The source acoustic axis (the nominal Z axis) may be horizontal, permitting relatively shallow (~20cm in depth) water tanks. This approach is especially useful for sources with long (> 20cm) focal lengths; it also allows materials to be conveniently placed between the ultrasound source and the hydrophone, for instance, to measure acoustic attenuation properties [88, 89]. Alternatively, the source acoustic axis may be vertical. This orientation generally requires a deeper (>50cm) water tank but alleviates the need to immerse the source completely into the water, which is useful when testing sources which are not fully waterproof.

Other considerations for the tank itself include reflective surfaces, temperature control, and water conditioning [85, 90]. For measurements of continuous waves, tank walls should be treated with absorbing/scattering materials to minimize reflections. For measurements of pulsed waves, wall treatments may also be helpful, but reflections may often be avoided simply by time-gating of signals. Water surface reflections can be mitigated somewhat using plastic floats (~2cm diameter), which serve to both break up the smooth boundary, and to minimize evaporative water loss and temperature variation. In addition, they reduce re-gassing rate of degassed water by reducing the surface area exposed to air. Temperature monitoring or control is required because both the sound speed in water and, in some instances, hydrophone sensitivity, are temperature dependent [90]. Water temperature is typically maintained a few degrees above ambient (for easier control). Finally, the water in the tank is typically filtered and deionized (close to 5 μS/cm). The deionization requirement originally stemmed from using single-layer unshielded coplanar membrane hydrophones, wherein the water around the hydrophone provided a conductive path across the elements, altering sensitivity in a frequency-dependent manner. Bilaminar membrane and fiber optic hydrophones, on the other hand, are immune to water conductivity issues. Typically, hydrophones should not be immersed longer than 24 h, but occasionally 48–72 h immersion can be tolerated. Further application-specific guidance [3–6] and techniques for water conditioning [91] are provided in IEC documents. Additional environmental considerations for specific hydrophones (e.g., handling instructions, minimum and maximum recommended soaking times, etc.) may usually be obtained from hydrophone manufacturers. Soaking allows microbubbles possibly adhering to the hydrophone structure to dissipate.

B. Mechanical Scanning Considerations

The mechanical hydrophone scanning mechanism has two primary design requirements: accurate alignment of the hydrophone and source acoustic axes, and positional resolution and repeatability [85]. The first requirement entails angular control of either the hydrophone, the source, or both. This can depend upon the type of hydrophone used. While NHs have directivity patterns that match the theory for a given active element size [35, 92], they can demonstrate a slight angular misalignment between their physical orientation and the direction of maximum sensitivity. This can lead to discrepancies in measurements, especially at higher source frequencies (or higher harmonics of non-linear waveforms) where the hydrophone directional response is narrower. For MHs, the direction of maximum sensitivity is exactly perpendicular to their surface, allowing the membrane to establish the XY plane of the measurement setup, and therefore the Z axis direction. It must be noted that MHs can have a non-circular effective aperture, due to fringing field effects, which again, may affect measurement accuracy.

A key component of overall measurement accuracy is the resolution and repeatability of the X-Y-Z positioning system. Generally, the positioning system step sizes should be smaller than one half wavelength of the highest frequency of significance in the source spectrum (including harmonics). However, in many cases (e.g., far field measurements), this guideline may be relaxed, as discussed in application-specific guidelines [3–6]. Another important issue to consider is whether backlash compensation is required. Backlash refers to positioning errors (due to slack in threads/gears) when an axis changes direction. (Constant force springs are one remediation). Finally, when the hydrophone is scanned in discrete spatial increments, it will shake slightly with each step. If the data capture system is set up to take multiple waveforms and time average them for noise reduction, the resultant data will be distorted until the positional “ringdown” has stopped. Thus, the hydrophone mounting arm must be as lightweight and stiff as possible, with the smallest possible moment arm between the drive mechanism and the hydrophone.

C. Waveform Capture

A digital oscilloscope is the preferred means of capturing the hydrophone output, rather than a computer board level or custom solution, because it is more easily calibrated by an outside traceable facility. Many current digital oscilloscopes have sufficient bandwidth (>100MHz), bit depth (>8bit) and record length (>2k) for ultrasound measurements and are then interfaced to the control computer. Oscilloscope waveform acquisition is best triggered by a signal from the device under test, rather than from the hydrophone signal. One recurring issue is the oscilloscope input impedance, typically either 1MΩ‖10–35pF or 50Ω, which must be accounted for to insure proper hydrophone calibration values [36, 93]. Piezoelectric hydrophones without preamplifiers specifically require sensitivity correction to compensate for the input electrical impedance of the oscilloscope [94, 95].

D. Computer Control

The scanning process is managed by a control computer which is interfaced to the motion control system and the oscilloscope [85–87]. The computer is programmed with key functions such as data capture, voltage and time scale adjustment and autoscaling, waveform analysis, and beam scanning. It also uses calibration data to establish the appropriate hydrophone sensitivity to convert voltage to pressure [96]. Although it slows the beam scanning process, autoscaling each waveform (automatically adjusting the vertical scale of the oscilloscope) can maintain optimal dynamic range for each measurement.

E. Measurement Process

It is important to verify that the central beam (Z) axis of the positioning system is co-linear with the transducer axis. This can be done by 1) finding a field maximum in the focal plane by 2D scanning, 2) finding a field maximum in another plane with evident maximum, 3) comparing X and Y coordinates of the two maxima, and then 4) iteratively correcting the position and/or orientation of the transducer if necessary. The hydrophone is then scanned along the Z axis to determine the depth(s) associated with the spatial peak pressure or intensity, either derated (i.e., compensated for attenuation that would be assumed to happen in tissue) or non-derated. The hydrophone then is scanned laterally (X and Y directions) in the focal plane to determine the beam widths. The spatial sampling interval is typically one-half wavelength at the measured center frequency. In addition to lateral beam scans, a full two dimensional “raster” scan may be performed to determine the total power emitted from the source and/or to obtain acoustic holography data for full 3D field reconstruction [6, 97].

V. Hydrophone Transfer Function Deconvolution and Measurement Uncertainty

A. Signal Processing to Deconvolve for Sensitivity

For a narrowband, quasi-planar beam, pressure may be computed by dividing hydrophone output by hydrophone sensitivity (e.g., in units of V/Pa) at the appropriate frequency. However, biomedical pressure waves are often either broadband or focused or both.

For broadband beams, hydrophone sensitivity might vary substantially with frequency throughout the pressure spectrum. In these cases, pressure should be computed by deconvolving hydrophone output with hydrophone sensitivity [43, 59, 62, 64, 65, 98–104], as shown in Fig. 7. This is usually done in frequency domain by dividing the spectrum of the hydrophone output voltage by the frequency-dependent hydrophone sensitivity. (See Sec. V-C.) The challenge of uncertainty determination for waveform parameters, including uncertainty propagation, through deconvolution has been addressed [99, 100, 104–106]. For many applications, deconvolution with sensitivity magnitude performs nearly as well as deconvolution with complex sensitivity [59, 65, 98]. However, methods exist for deriving sensitivity phase from sensitivity magnitude to enable complex deconvolution when phase is not directly measured [6, 66, 68, 69, 106]. The importance of sensitivity deconvolution has been recognized in an international standard [6].

Fig. 7.

A pressure waveform measured using membrane (NT, S5, ST), needle (ON, DI), capsule (GL), and Fabry-Perot fiber optic (PA) hydrophones before and after sensitivity deconvolution. Some signals are shifted in time, but this does not affect measurements of peak compressional pressure or peak rarefactional pressure. Reprinted (after reformatting) from [98],

B. Signal Processing to Deconvolve for Spatial Averaging

For focused beams, the incident pressure might vary substantially with position across the hydrophone sensitive element surface. In these cases, pressure measurements must account for the fact that hydrophones average the incident pressure over the sensitive element surface.

Fig. 6 illustrates how spatial averaging effects can arise. Fig. 6 compares frequency-dependent harmonic component beamwidths to hydrophone geometrical and effective sensitive element diameters d_g and d_eff(f). Effective diameters d_eff(f) are frequency-dependent and derived from directivity measurements [80, 84].

Most spatial averaging correction methods assume circular transducer geometry and are based on either empirical or numerical methods [90, 107–111]. Recently, analytic forms for filters that model spatial averaging have been derived and validated for beams with circular [112, 113] and rectangular [114–116] symmetry. These spatial averaging filters can be used in inverse filtering approaches for spatial averaging correction [78, 116, 117]. One analytic form has been incorporated into an international standard [6, 112]. Fig. 8 shows an example of spatial averaging correction resulting in reduction of error in peak compressional pressure from 23% to 4% [117].

Fig. 8.

A pressure waveform measured using reflectance-type fiber optic (HFO) and needle (HNA) hydrophones before and after spatial averaging corrections. Reprinted from [117].

Spatiotemporal deconvolution theory has been used to derive simple formulae, given in Appendix A of [118], to allow quick estimation of spatial averaging correction factors, without inverse filtering. Originally developed and validated for MHs, NHs, and FOHs, they may be extended to CHs [84].

C. Filtering

When performing inverse filtering, erratic results may arise from division in frequency domain by small values. These artifacts have been avoided by using Gaussian [98], Wiener [100], and Butterworth [78, 117] filters in sensitivity deconvolutions and Wiener filters [119] in spatial averaging deconvolutions.

D. Criteria for Maximum Appropriate Hydrophone Sensitive Element Size

Spatiotemporal deconvolution theory has been used to derive criteria for maximum appropriate hydrophone geometrical sensitive element diameter d_g [118]. Unlike the commonly cited IEC criterion [6], these criteria were derived for focusing rather than planar transducers, are applicable to nonlinear signals in addition to linear signals, and are expressed in terms of the easily accessible d_g instead of d_eff(f).

E. Measurement Uncertainty

Even when variabilities related to factors such as sensitivity, directivity, and spatial averaging are either negligible or corrected for, comparisons of measurements of pressures and intensities from a single source performed with multiple (at least 4) hydrophones have exhibited coefficients of variation on the order of 10%–20% in the range of 1 to 3.5 MHz [59, 98, 102, 120]. Much of this variation is likely due to uncertainty in hydrophone calibration (which is usually relative instead of primary; see Sec. III). Hydrophone-to-hydrophone measurement variation depends on measurand (e.g., p_c, p_r, pii, or beam width) and tends to increase with frequency and intensity. Investigators who perform measurements with a single hydrophone would be prudent to acknowledge in their reports that their measurement uncertainty may be at least 10%–20%.

Responsible reporting of hydrophone-based pressure, intensity, and beamwidth measurements should always consider potential effects of frequency-dependent hydrophone sensitivity, spatial averaging, and measurement uncertainty.

VI. Considerations for High Intensity Therapeutic Ultrasound

Reliable characterization of HITU (a.k.a. high intensity focussed ultrasound, HIFU [121]) fields is important to demonstrate safety and effectiveness of an application, for instance, tumor ablation [122–125]. However, the required hydrophone measurements at clinical driving levels pose several metrological challenges [126]. Extreme focal pressure amplitude and intensity values several orders of magnitude larger than those used in diagnostic ultrasound or ultrasound physiotherapy may easily cause damage to the sensors commonly used for acoustic field measurements. Furthermore, the pressure waveforms are strongly nonlinearly distorted and comprise many higher harmonic frequency components. This results in additional demands on the bandwidth of the measurement system [6]. Finally, a small sensing element size is desirable to be able to correctly determine the spatial peak pressure and intensity values as well as the small beamwidths produced by the strongly focusing transducers.

Care must be taken with respect to any disturbance possibly caused by acoustic reflections from hydrophones. Particularly for hydrophone measurements close to the focal distance of the transducer, reflections are likely to be collected by the source. If any instabilities of the acoustic output of the HITU transducer are registered, the burst length should be reduced to separate in time the transmitting period from the period of incident reflections and multipath reflections. Furthermore, it should be kept in mind that HITU fields at clinical driving levels are likely to introduce damage to any polymeric materials, like sensor housing or MH rings, exposed to the focal region due to sound absorption and heating. HITU fields may lead to de-electrodization and hence alter the sensitivity of the hydrophone. Degassing the water is recommended to avoid cavitation.

Laboratory-internal comparison studies have been performed to investigate the applicability of different sensor types and the implications on the variation of HITU field characterization [102, 127].

FOHs have been increasingly used for HITU field measurements [47, 128–131]. However, FOHs are relatively expensive, difficult to use, not widely available, and in some cases have bandwidth limits that preclude detection of all significant harmonics (e.g., 30 MHz for some measurements in [47] and [129]). Positioning and position stability of the sensing fiber tip are difficult to assure when measuring in the focal region. For example, acoustic radiation forces have been observed to possibly displace fiber tips during measurements [130]. Also, disturbance of measurements caused by mechanical reflections from the fiber mounting can become an issue. A substantial limitation of FOHs applicable for very high-pressure amplitude measurements seems to be a comparatively high noise equivalent pressure (NEP) of several MPa [131]. In this case, a second more sensitive hydrophone may be needed for field characterization outside the focal region [132].

Robust piezoelectric NHs comprising a metallic coating to protect against cavitation damage of the electrode have been developed [39, 41, 133] and have been applied to measurements up to 48 MPa peak compressional pressure, providing a bandwidth of at least 30 MHz [117]. CHs have been applied successfully to measure moderate HITU fields and have a reported peak compressional pressure limit of approximately 10 MPa for frequencies of 1–3 MHz [102].

MHs [17] are commonly not considered to be very robust sensors. However, the general advantages of MHs for HITU measurements are broad bandwidth [23], flat and predictable frequency response, and small sensing element size. Conventional MHs have been applied at low, quasi-linear sound propagation conditions and at medium amplitude conditions [134]. The pressure range detectable without destruction of a conventional MH was determined by means of a series of measurements at focus with increasing driving voltage [24]. A hydrophone design comprising additional protective layers has been developed to increase the robustness against cavitation (Fig. 9). Such devices have been used for measurements at clinical pressure amplitude levels up to 80 MPa peak compressional pressure [105]. Duty cycles, however, should be reduced compared to clinical HITU applications to avoid excessive output intensities during field characterization and to enhance cavitation thresholds. Additional protection layers alter the frequency response from the common form for MHs. Typically, an overall low-pass behavior combined with a thickness-mode resonance can be observed [105, 135] (Fig. 10). Other approaches have included disposable MH elements [136] as well as reflector-type hydrophones [137].

Fig. 9.

HIFU membrane hydrophone stainless steel front protection layer and a silicone oil backing (GAMPT mbH, SI04)

Fig. 10.

Magnitude and phase of frequency response for HIFU membrane hydrophone in Fig. 9. Expanded uncertainty of the calibration data for k = 2 (95% confidence). Reprinted from [135]. ©2019 IEEE.

To enable unbiased measurements of broadband acoustic pulse waveforms, waveform deconvolution can be applied [6, 59, 64, 106] to compensate for frequency-dependent sensitivities. Calibration methods are available that provide the necessary broadband and complex-valued calibration data [64, 71, 105, 135]. Similarly, the effect of spatial averaging of finite receiving elements can be corrected for by applying standardized methods [6, 117].

VII. How to Choose an Appropriate Hydrophone for Your Application

To choose from among the wide variety of available hydrophone designs, some foreknowledge of the application is needed. Usually, users employ some combination of previous experience, experimentation, or modeling to provide estimates of pressure levels, bandwidth, and spatial variation of the acoustic signal they wish to measure. This information is then used to select a hydrophone based on the following typical considerations:

A. Minimum Pressure Level

The minimum pressure level to be measured should be comfortably above the noise equivalent pressure (NEP) of the hydrophone assembly—typically at least by a factor of ten. NEP may be calculated by dividing the noise level of the amplifier or data acquisition system (e.g., V) by the sensitivity of the hydrophone assembly (e.g., V/Pa). The noise level of the amplifier is usually provided by its manufacturer and can be further reduced with filtering and (in the case of stable repetitive signals) averaging. For a given hydrophone type, NEP tends to be inversely related to sensitivity and sensitive element area.

The minimum pressure to be measured should also be about a factor of ten higher than any synchronous electrical interference detected by the hydrophone. This problem usually occurs when measuring a long tone burst at a position close enough that the source transducer is still electrically excited at the time of acoustic arrival. In such cases, either a well-shielded high-sensitivity piezoelectric or optical hydrophone should be chosen. See Table II.

B. Maximum Pressure Level

The maximum pressure to be measured should be within the linear range of the hydrophone assembly, as defined by a 10% nonlinearity criterion detailed in [83], section 5.1.7. This threshold is typically due to saturation of the preamplifier or due to similar reversible electronic phenomena which leave no permanent damage to the hydrophone assembly. Manufacturers can be expected to provide the maximum pressure that the hydrophone assembly can measure and still be within the linear range.

The second consideration for maximum pressure is the potential for damage to the hydrophone due to cavitation, mechanical stress, or heating. Usually, such damage is not a concern for diagnostic applications. However, measurements of therapeutic fields may result in damage necessitating the repair or replacement of a hydrophone, unless it is particularly robust as described in Section VI. In cases involving high pressure or intensity the user should seek guidance from the hydrophone manufacturer on mitigating the risks of damage. See Table II.

C. Frequency Response

As discussed in Section V, deconvolution can correct for variation of the hydrophone’s sensitivity over the frequency range of the acoustic signal. However, use of a hydrophone with a flatter response may result in lower measurement uncertainty, and may not require additional broadband calibration of the hydrophone. Ref. [6], Section 5.7 provides criteria for when broadband calibration and deconvolution are required, based on the characteristics of the acoustic signal and the hydrophone’s frequency response.

For most hydrophones, frequency response is determined by hydrophone resonance frequency and associated electronics, including preamplifiers and amplifiers [36, 93]. The fundamental mode of vibration corresponds to half- or quarter-wavelength inside the sensing medium depending on the existence and properties of backing material, resulting in widest attainable bandwidths ranging from 100 kHz to 140 MHz. Amplifiers may amplify current or voltage. Current amplifiers minimize sensitivity loss due to cable length, acting as capacitive voltage divider. Voltage amplifiers are easier to implement but need to be incorporated in the immediate vicinity of the active element electrodes (ideally < 1–2 cm). Preamplifier complex input impedance also plays a critical role in determining bandwidth. See Table II.

D. Sensitive Element Size

A small effective diameter is usually desired to minimize phase and amplitude variation across the hydrophone aperture, and to reduce the sensitivity of the measurement to alignment errors. As discussed in Section V, criteria for maximum geometrical sensitive element diameter have been derived that are applicable to acoustic beams generated by focused transducers that may be linear or nonlinear (i.e., with significant energy in harmonic components).

Generally, the effective diameter is sufficiently small if it is less than one quarter of the minimum acoustic wavelength in the spectrum of the acoustic beam to be measured. However, with increasing acoustic frequency or increasing nonlinearity (i.e., harmonic content), achieving this strict criterion typically requires tradeoffs with sensitivity, bandwidth, robustness, and purchasing price; note also that above 10 MHz, hydrophones meeting this strict criterion are not currently available commercially. As discussed in Section V, the effective diameter depends on frequency, is larger than geometrical diameter at low frequencies, and may be derived from directivity measurements (see Fig. 6). A relaxed criterion can be derived for measurements of linear beams at the focus or in the far field [6, 138], but may still be difficult to satisfy. In practice it may be advisable to choose a larger aperture, in which case the analysis of Section V may be used to correct for spatial averaging and/or assess additional uncertainty in the measurement due to the non-ideal size of the aperture. See Table II.

E. Hydrophone Type

As discussed in Section II, biomedical hydrophones come in several form factors. MHs, while generally providing the best bandwidth, cannot, for example, measure close to the radiating surface of curved acoustic sources. Additionally, hydrophones that present a large reflection profile (e.g., MHs) may create standing waves, unless the reverberations are mitigated by tilting the hydrophone, which comes at the cost of directivity. Users may therefore choose NHs and CHs for some applications, at the expense of bandwidth.

To measure high pressures (e.g., HITU), FOHs and robust MHs and NHs are preferred. FOHs have high spatial resolution and can withstand high pressures but tend to be comparatively expensive and difficult to use.

VIII. Recommendations for Reporting Hydrophone Measurements

To promote reproducibility of measurements and consistency of scientific publications, several articles offer recommendations for reporting experimental conditions for investigations of potential ultrasound bioeffects [139–141]. To complement those articles, the present article provides substantially more motivation for the subset of recommendations that involve hydrophone measurements.

Considerations discussed in the present article are the basis for recommendations listed in Table III for good practice in reporting on hydrophone measurements (although it is understood that some items might not always be available). While some of these recommendations may be found elsewhere, others are unique to the present article. For supporting explanation and guidance, each recommendation is followed by a parenthetical reference to the relevant section in the present article.

TABLE III.

Recommended Information to be Included in Reports of Hydrophone Measurements

hydrophone manufacturer and model (Sec. II), hydrophone type (e.g., MH, CH, NH, FOH reflectance type, FOH interferometric type) (Sec. II), sensitive element diameter (See Secs. II, III, V, and VII), conformance (if any) with IEC standards [3–6] (Sec. I) formulas for calculating derived quantities such as intensities, MI, TI, pulse duration, acoustic working frequency, and beam width (Sec. I) most recent calibration date, bandwidth, specified uncertainty, and whether calibration was primary or relative (Sec. III), manufacturers) and model(s) of components including attenuators, amplifiers, and preamplifiers (Secs. IV, VII), beam orientation (e.g., horizontal, vertical) (Sec. IV), waveform parameter for hydrophone alignment for focal point (e.g., pc, pr, ppsi) (Secs. I, IV), water conditioning (e.g., temperature, deionized, degassed) (Sec. IV), strategy for managing effects of reflections on measurements (Sec. IV), oscilloscope model and acquisition parameters including input impedance, bit depth, sampling rate, acquisition time, signal averaging, and triggering (Sec. IV) sensitivity correction procedure to compensate for oscilloscope input impedance (if any) (Sec. IV) scanning step and range (Sec. IV) whether and how sensitivity deconvolutions and/or spatial averaging corrections were applied (Sec. V). estimate(s) of measurement uncertainty, accounting for sensitivity calibration uncertainty, variability due to repositioning, expected inter-hydrophone variability, and other variables relevant for the measurement (Sec. V). strategies (if any) to mitigate HITU-related risks to hydrophone and hydrophone measurements, such as duty factor reduction, cavitation monitoring, and spot-checks of hydrophone after usage (Sec. VI) NEP (which can be estimated from parts of voltage waveforms where signal is absent) (Sec. VII)

Where appropriate, brief justifications should accompany these details. Relevant sections for explanation and guidance are shown in parentheses.

IX. Conclusion

Since the seminal Special Issue on Ultrasound Exposimetry published in IEEE T-UFFC in 1988 [142], considerable progress has been achieved regarding measurement standards and methodology for hydrophones. In addition, some progress has been made in development of hydrophone designs, especially FOHs. FOHs have high spatial resolution and can withstand high pressures but tend to be comparatively expensive and difficult to use. MHs, NHs, and CHs (based on PVDF) are adequate tools for many applications, with sufficient sen sitivity, robustness, bandwidth, and spatiotemporal resolution.

Complete characterization of hydrophone performance entails calibration of sensitivity and directivity over a broad frequency band. Accurate hydrophone measurements require careful attention to the measurement environment, including the geometry of the water tank, scanning apparatus, data acquisition, and water conditioning.

Primary sources of discrepancies among measurements performed with different hydrophones include complex electroacoustic frequency response, frequency-dependent effective sensitive element size, and sensitive element asymmetry. However, these discrepancies can be mitigated with appropriate signal processing.

Measuring high intensity ultrasound is particularly challenging because of the high nonlinearity of the signals and the potential for damage to the hydrophone.

Making an intelligent choice for a hydrophone for a specific task requires balancing various factors, including minimum and maximum pressure levels to measure, frequency response, sensitive element size, cost, fragility, and compactness.

Biographies

Andrew M. Hurrell was born in Ashford, Kent, UK in 1972. He received his BSc (Hons) in physics with modern acoustics from the University of Surrey, UK, in 1994 and his PhD in underwater acoustics from the University of Bath, UK, in 2002. From 1994 to 1996 he was a member of the Acoustics Materials Team at the Defence Research Agency (Holton Heath). In 1996 he joined Precision Acoustics Ltd, Dorchester where he is currently the Principal Research Physicist. During that time, he has authored more than 40 papers and conference presentations. He has also contributed to 3 textbooks on ultrasonics. His current interests included the design and construction of hydrophones and hydrophone arrays, development of PVDF sensors/transducers and the use of analytical and finite difference techniques to simulate ultrasonic transduction and propagation phenomena. Andrew was recipient of the IOA Young Persons’ Award for Innovation in Acoustics Engineering and is peer reviewer for numerous journals and grant funding authorities. He serves as one of the UK members of IEC Technical Committee TC87 (Ultrasonics) and chairs the British Standards Committee EPL/87 that shadows it. He is also a member of the British Standards Committee EH/1/7 (Underwater Acoustics).

Mark Schafer (S’83-M’86-SM’16) was born in Pittsburgh, PA. He received the SB degree in Electrical Engineering from the Massachusetts Institute of Technology in 1979, an MS in Acoustics from the Pennsylvania State University in 1982, and a Ph.D. in Biomedical Engineering from Drexel University in 1988. He is currently Research Professor in the School of Biomedical Engineering, Science, and Health Systems at Drexel University, with continued research efforts in biomedical ultrasound. Prior to joining Drexel, he held management positions with several medical companies as CTO or VP of R&D. He has consulted with firms worldwide on design, development, intellectual property, regulatory and clinical aspects of medical ultrasound products, including diagnostic, therapeutic and surgical applications. He is a serial entrepreneur and inventor on over 30 patents, and has authored numerous journal articles and book chapters on ultrasound measurement and applications. Dr. Schafer is a Fellow of the American Institute of Ultrasound in Medicine, the Acoustical Society of America, and the American Institute of Medical and Biological Engineering, a recipient of the Chief’s Award for Technology Transfer, U. S. Department of Agriculture, past President of the Ultrasonic Industry Association, and is currently President of the IEEE UFFC Society.

Samuel M. Howard received his A.B. in Engineering Sciences from Harvard University. He received his M.S. and Ph.D in Theoretical and Applied Mechanics from Cornell University with minors in Applied Mathematics and Physics. From 1989 to 2001 he worked for Acuson (later Siemens) in transducer R&D. From 2001 to the present he has been at Onda Corporation, where he is the CTO and directs the Acoustics Laboratory. His current interests focus on ultrasound metrology for medical and industrial applications.

Keith A. Wear (Life Fellow, IEEE) received the B.A. degree in applied physics from the University of California at San Diego, San Diego, CA, USA, and the M.S. and Ph.D. degrees in applied physics, with a Ph.D. minor in electrical engineering, from Stanford University, Stanford, CA. He was a Postdoctoral Research Fellow with Physics Department, Washington University, St. Louis, MO, USA. He is currently a Research Physicist with the U.S. Food and Drug Administration, Silver Spring, MD, USA. Dr. Wear is a Fellow of the Acoustical Society of America, the American Institute for Medical and Biological Engineering, and the American Institute of Ultrasound in Medicine (AIUM). He received the 2019 AIUM Joseph H. Holmes Basic Science Pioneer Award. He served as Associate Editor-in-Chief for IEEE TRANSACTIONS ON ULTRASONICS, FERROELECTRICS, AND FREQUENCY CONTROL (IEEE-TUFFC) (2019–2021). He has served as Associate Editor of IEEE-TUFFC (2002–2021), Journal of the Acoustical Society of America (2012-present), and Ultrasonic Imaging (2013-present). He was the Technical Program Chair of the 2008 IEEE International Ultrasonics Symposium (IUS), Beijing, China. He was the General Program Chair of the 2017 IEEE IUS, Washington, DC, USA. He has served as the chair of the AIUM Technical Standards Committee (2014–2016), AIUM Bioeffects Committee (2021–2023), AIUM Basic Science and Instrumentation Community (2004–2006 and 2014–2016), and AIUM Therapeutic Ultrasound Community (2013–2015). He is Chair of the American Association of Physicists in Medicine Task Group 333 on Magnetic Resonance Guided Focused Ultrasound Quality Assurance. His research interests include hydrophone measurement methodology, high intensity therapeutic ultrasound, photoacoustics, and quantitative ultrasound.

Volker Wilkens was born in in Oldenburg, Germany, in 1969. He received the diploma degree in physics from the University of Oldenburg, Germany, in 1997, where he had worked in the field of optical metrology and spatial phase shifting techniques in electronic speckle pattern interferometry. Since 1997, he has been engaged in optical measurement of ultrasound and fiber-optic sensors at the Physikalisch-Technische Bundesanstalt (PTB) in Braunschweig. He received the Dr. rer. nat. degree in physics from the University of Oldenburg in 2001 for his work on optical multilayer hydrophones. In 2005 he became head of the ultrasonics working group at the PTB. Dr. Wilkens is responsible for customer services like hydrophone, ultrasonic power meter and reference transducer calibrations and acoustic output measurements on prototypes as well as for the maintenance of the national standard measurement setups. His interest in ultrasonic exposimetry and safety aspects of medical ultrasound including the development of ultrasound sensors and calibration techniques has been supplemented by active work for international standardization in the International Electrotechnical Commission (IEC), Technical Committee 87, starting in 2006. In 2011 he was elected to become chair of the German mirror committee GUK 821.3 “Medizinische Ultraschallgeröte” at the “Deutsche Kommission Elektrotechnik” (DKE). In 2013 he was appointed convenor of IEC TC 87 WG 8 “Ultrasonic Field Measurement”. In 2016 he became chair of IEC TC 87 “Ultrasonics” and was confirmed for a second term in 2022. Dr. Wilkens is a member of the German Acoustical Society (DEGA).

Bajram Zeqiri was born in Salerno, Italy, in October 1957. He joined NPL in 1984 following the completion of a BSC in Chemical Physics and PhD in solid-state chemistry at the University of Kent, Canterbury. His research has seen him contribute to the development of ultrasonic measurement techniques addressing a range of topic areas: acoustic property of materials determination, calibration and application of ultrasonic hydrophones, high-power cavitating acoustic fields, ultrasonic power measurement, standards for physiotherapy ultrasound equipment and the development of novel sensors for Ultrasound CT imaging. Currently Head of Science for Medical and Marine Physics at NPL, in 2019 Bajram was appointed visiting Professor within the Department of Medical Physics and Biomedical Engineering, University College London. In 2021, he was elected a Fellow of the Royal Academy of Engineering and in the same year he won the UK Institute of Physics James Joule Medal and Prize for distinguished contributions to the development of acoustic measurement techniques and sensors. He has authored around 85 peer reviewed publications and filed six patents in the area of ultrasonic measurement.

Peter A. Lewin, M.Sc., Ph.D. is R.B. Beard Distinguished University Professor of Electrical and Computer Engineering, Drexel University, Philadelphia, and Director of the Ultrasound Research and Education Center in The School of Biomedical Engineering, Science and Health Systems at Drexel University. Dr. Lewin obtained his M.S. degree in Electrical Engineering in 1969 and the Ph.D. in Physical Acoustics in 1979 in Copenhagen, Denmark. Before receiving his Ph.D. degree he was employed by Bruel and Kjaer, Denmark, where he was involved in the development of underwater piezoelectric transducers and associated electronics. From 1978 to 1983 he was associated with the Danish Institute of Biomedical Engineering (now Force Institutes) and The University of Denmark, Copenhagen-Lyngby, where his research activities primarily focused on propagation of ultrasound waves in inhomogeneous media and development of PVDF transducers. In 1983 he joined the faculty of Drexel University. Dr. Lewin was awarded several patents in the field of ultrasound and has authored or co-authored 250 scientific publications, most of them on topics in ultrasound. He is also co-editor of Ultrasonic Exposimetry (CRC Press, 1993), a landmark book in the field. His current interests are primarily in the field of biomedical ultrasonics including the design and testing of piezoelectric transducers and sensors, power ultrasonics, ultrasonic exposimetry, tissue characterization using nonlinear acoustics, biological effects of ultrasound, applications of shock waves in medicine, image reconstruction and processing, and ultrasonically assisted chronic wound healing. Dr. Lewin is elected Life Fellow of the Institute of Electrical and Electronics Engineers (IEEE). He is also a Fellow of the American Institute of Ultrasound in Medicine (AIUM), Acoustical Society of America (ASA), American Institute for Medical and Biological Engineering (AIMBE) and Elected Fellow of International Academy for Medical and Biomedical Engineering (IAMBE). In addition, he is Fellow of the College of Physicians in Philadelphia. He has also served as a Chair (1997–1999) of the AIUM’s Technical Standards Committee and the AIUM’s Board of Governors (2002–2004). Dr. Lewin serves as a consultant to the U.S. Food and Drug Administration, Center for Devices and Radiological Health and was appointed to the US Technical Advisory Group of ANSI to the International Electrotechnical Commission (IEC). He also served as Associate Editor of the peer-reviewed journal “Ultrasonics” and completed his tenure (2019–2022) as elected Editor-in-Chief of the IEEE Transactions on Ultrasound, Ferroelectrics and Frequency Control. He was elected as a resource member of the prestigious Franklin Institute Science and Awards Committee, Philadelphia and in April of 2018 received the Institute of Electrical and Electronics Engineers (IEEE) Philadelphia Section Benjamin Franklin Key Award. The award recognizes outstanding technical innovations, which contributed to intellectual, industrial, and economic development and demonstrated human benefits. Dr. Lewin is also recipient of Drexel University Provost Career Award for Outstanding Scholarly Productivity. Most recently, he was tapped by the Secretary of Health to serve on the NIH Advisory Council and received the highest recognition as a Distinguished Advocate and Fellow from The American Institute for Medical and Biological Engineering (AIMBE) for seminal contributions to the field of biomedical ultrasonics and development of new piezoelectric transducers and measurement methods.

Gerald R. Harris (M’72-S’76-M’79-SM’82-F’94-LF’11) was born in Jacksonville, NC, on November 22, 1945. He received the bachelor’s degree in electrical engineering in 1967 from the Georgia Institute of Technology, Atlanta, the M.S. degree in biological engineering in 1971 from the Rose-Hulman Institute of Technology, Terre Haute, IN, and the Ph.D. degree in electrical engineering in 1982 from the Catholic University of America, Washington, DC. He is retired from the Food and Drug Administration’s Center for Devices and Radiological Health, Silver Spring, MD, where his main activities comprised the experimental and theoretical evaluation of medical ultrasound transducers and systems. Dr. Harris is a Fellow of the Acoustical Society of America, American Institute of Ultrasound in Medicine, and a Life Fellow of the Institute of Electrical and Electronics Engineers.