WO2017051164A1

WO2017051164A1 - Method and device for acoustic particle palpation

Info

Publication number: WO2017051164A1
Application number: PCT/GB2016/052929
Authority: WO
Inventors: James Choi; Hasan Koruk
Original assignee: Imperial Innovations Limited
Priority date: 2015-09-21
Filing date: 2016-09-20
Publication date: 2017-03-30
Also published as: GB201516714D0

Abstract

The present invention relates to a method of measuring the elasticity of a material of interest, the method comprising palpation of the material of interest using at least one acoustically responsive particle which has been administered into or adjacent to said material, wherein said palpation comprises the use of acoustic waves to push the at least one acoustically responsive particle against the material of interest to cause transient deformation of the material and measuring the transient deformation of the material and/or the force response. The present invention also relates to a device for palpating a material of interest with at least one acoustically responsive particle and measuring the resulting deformation of the material of interest and/or force response, the device comprising a palpation component comprising one or more transducers configured to generate an acoustic wave suitable for pushing at least one acoustically responsive particle against the material of interest to cause transient deformation of the material of interest within a region of palpation and an active monitoring component configured to actively image the material of interest and/or the acoustically responsive particle.

Description

Method and device for acoustic particle palpation

Field of the invention

The present invention relates to methods and devices for using acoustic wave(s) and acoustically responsive particles to palpate and measure the elasticity of a material.

Background to the invention

Palpation, and the monitoring of the resulting deform ation or force response, are effective in determ i ni ng a m aterial's mecha nical properties^1-5. In the lab and the clinic, this two-part process is used in the diverse and complex range of elasticity measurement devices available^6-8.

Atomic force microscopy (AFM)¹ uses a nano-sized tip at the end of a cantilever to palpate the surface of structures. Optical⁹ and magnetic¹⁰ tweezers use particles responsive to light or magnetism, respectively, to apply stress. In the clinic, manual palpation of superficial tissue provides a qualitative assessment of stiffness and is used to diagnose diseases such as breast cancer.

In order to assess deeper tissue, an ultrasound beam can be used to palpate by exerting an acoustic radiation force (A F) in the direction of propagation¹¹. Palpation is the fundamental unit of these elasticity measurement systems and the characteristics of the stress source, for example size, distribution and magnitude, determine the capabilities and limitations of the system, which include resolution, sensitivity and depth of use.

In contrast to other stress sources, ultrasound has the unique ability to palpate material beneath the surface of materials by focussing the beam. Whereas physical objects press directly against the surface of a material, ultrasound propagates through the material while momentum is transferred from the acoustic wave onto the material through absorption, scattering, and reflection. Thus ARF-induced stress is applied not from the surface of a material, but throughout a long ell ipsoidal beam volume that is in the order of a few millimetres wide and tens of millimetres long. The size of this focal volume is limited by the wavelength of the acoustic centre frequency. Aside from the limitations on resolution, this larger stress volume makes it more susceptible to a breakdown of the assumptions of tissue homogeneity. In other words, there is uncertainty with how the ARF-induced stress is distributed within a beam, because it is dependent on the material's unknown absorption properties. Complications arise in materials with masses, layers, vessels and cavities, for example.

Although A F-based ultrasound elasticity imaging has had some success in the clinic, for example in the diagnosis of diffuse liver diseases and breast masses, the vast majority of diseases with known changes in mechanical properties remain undiagnosable¹² due to limitations in currently available in vivo stress sources.

US Patent Application No. 2004/0054357 describes a method of measuring tissue elastic properties using acoustic radiation force on laser-generated bubbles, wherein the displacement of the bubble induced by acoustic radiation force is directly related to the elasticity of the surrounding material. However, such a method has a number of disadvantages. The laser used to generate the bubble requires physical destruction of the material. When considering the construction of a full elasticity image, this would require complete destruction of the material, thereby making it impractical for clinical imaging. Furthermore, lasers cannot be focused to create such bubbles deeper than a few hundreds of microns, limiting the method to skin and other superficial targets. The use of a laser to generate the bubbles allows little control over bubble size, and the bubbles created generally have a diameter of 100 to 800μιη. Such bubbles are too large to be injected or to circulate the entire body. The use of large bubbles also presents a risk of embolism, stroke and heart attack. The ultrasound frequency used to displace the bubbles is far from the resonance size of the bubbles formed. The bubbles are therefore only moving due to reflection, limiting the resulting displacement and the force generated by the bubbles.

There is therefore a need in the art for a method of measuring the elasticity of a material which provides a better elasticity imaging resolution and a higher contrast for elasticity imaging, which generates a force applied only to the surface of interest, which does not require any destruction of the surface of interest and which is not limited to superficial targets but can be used to measure the elasticity of surfaces deep within the body. Also required are devices for carrying out such methods.

Citation of any reference in this section of the application is not an admission that the reference is prior art to the invention. The above noted publications are hereby incorporated by reference in their entirety. Summary of the invention

The present inventors have surprisingly found that acoustically responsive particles displaced against a material of interest using acoustic waves can result in sufficient transient deformation of the material to allow measurement of the elasticity of that material, either through the measurement of the deformation itself or through the measurement of the force response. Unlike methods known in the art, the method of the present invention can provide a physiologically relevant method of using acoustically responsive particles as a stress source for measuring the elasticity of a material of interest, such as human tissue.

Accordingly, in a first aspect the present invention provides a method of measuring the elasticity of a material of interest, the method comprising: a) palpation of the material of interest using at least one acoustically responsive particle which has been administered into or adjacent to said material, wherein said palpation comprises the use of acoustic waves to push the at least one acoustically responsive particle against the material of interest to cause transient deformation of the material; and b) measuring the transient deformation of the material and/or the force response. This method may be described as acoustic particle palpation.

A second aspect of the present invention provides a device for palpating a material of interest with at least one acoustically responsive particle and measuring the resulting deformation of the material of interest and/or force response, the device comprising: a) a palpation component comprising one or more transducers configured to generate an acoustic wave suitable for pushing at least one acoustically responsive particle against the material of interest to cause transient deformation of the material of interest within a region of palpation; and b) an active monitoring component configured to actively image the material of interest and/or the acoustically responsive particle.

Detailed description of the invention

Elasticity is defined as the resistance of a material to deformation when it is exposed to a force. For a method of measuring elasticity using palpation to be successful, it needs to generate sufficient displacement of the material of interest to allow imaging, tracking and/or measurement of that displacement. It has been found that acoustically responsive particles exposed to acoustic waves experience a primary acoustic radiation force that causes them to move in the direction of wave propagation.

An alternative way of displacing an acoustically responsive particle is through the creation of a standing wave. Acoustically responsive particles will move towards or away from different nodes and anti-nodes. A standing wave can be created when an emitter or transducer emits a pressure wave in one direction and the wave is reflected by a reflector in the opposite direction. These opposite travelling pressure waves create standing waves that acoustically responsive particles can interact with.

Either of these techniques may be used to displace the acoustically responsive particles in the method of the present invention.

The acoustic waves may be emitted in the form of a focused wave. Where a focused beam is emitted to a volume of material, the deformation and force response of that volume, or of sub- regions within that volume, may be imaged and/or measured. Multiple different regions of the material may be palpated to construct an image.

Alternatively, the acoustic waves may be emitted in the form of a plane wave. Where a plane wave is emitted to a large region of material, the deformation and force response of sub-regions within that region may be imaged and/or measured.

The present invention uses acoustic waves to push one or more acoustically responsive particles against a material to the point of safe and transient deformation. The extent of the deformation and/or the force response can then be used to determine the elasticity of the material.

Preferably, the acoustically responsive particles are displaced by the acoustic waves at a higher magnitude than the surrounding material. Preferably, the acoustically responsive particles are displaced by the acoustic waves at a magnitude between 2 and 100 times higher than the surrounding material. This occurs where the acoustically responsive particles have an acoustic impedance mismatch with the surrounding material, the result of which is that the acoustically responsive particles experience a far greater acoustic radiation force than the surrounding material. One way in which this can be achieved is to use lower acoustic pressure and intensities than other acoustic wave-based elasticity imaging methods. The present inventors have determined that the amount of displacement can be greatly increased by using acoustic waves with a centre frequency within the resonance frequency range of the acoustically responsive particle(s). Preferably, the centre frequency of acoustic waves matches, or approximately matches, the resonance size of the acoustically responsive particle(s). Under some circumstances, it may be desirable to use acoustic waves with a frequency which does not precisely match the resonance frequency of the acoustically responsive particles. Particles exposed to ultrasound off-resonance frequency can still move due to reflection, scattering and absorption. If the size of the particle is greater than the resonance size, increasing the size will increase the amount of energy reflected, thereby propelling the particle away from the transducer source in a manner similar to if it was at resonance. Preferably, the centre frequency of the acoustic waves is within a few orders of magnitude of the resonance frequency of the acoustically responsive particle, preferably within one, two or three orders of magnitude.

The use of acoustic waves with centre frequencies within the resonance frequency range of the acoustically responsive particles enables the particles to deform the material of interest to a much greater extent than would normally be expected. The magnitude of the primary acoustic radiation force, and hence the response or displacement of an acoustically responsive particle, increases as the frequency of the acoustic waves gets closer to the resonance frequency. The response decreases as the frequency of the acoustic waves gets further from the resonance frequency. This provides a unique frequency range in which the acoustically responsive particles will maximally respond.

Acoustic waves may be used at a range of frequencies. If the population of acoustically responsive particles is monodisperse, a narrow range of frequencies is preferred as this will usually generate the highest acoustic radiation force. If the population of acoustically responsive particles is polydisperse, a broader range of frequencies may be desirable in order to excite different centre frequencies, either in sequence or in the same pulse, in order to take advantage of the full range of particle sizes present. Alternatively, it may be desirable to use a narrow range of frequencies to selectively displace acoustically responsive particles within a particular size range.

Preferably, the acoustic waves are ultrasound waves. Lower-frequency ultrasound enables very deep tissue palpation. The preferred centre frequency (f_c) of the ultrasound will depend at least in part on the depth of the imagining required. Preferred ultrasound centre frequencies range from 100kHz to 500MHz, more preferably 200kHz to 50MHz, more preferably 300kHz to 20M Hz, more preferably 500kHz to 15MHz, more preferably 500kHz to 10M Hz, more preferably 100kHz to 10M Hz, more preferably 2M Hz to 10M Hz, more preferably 100kHz to 5M Hz, for example lM Hz to 5M Hz. The ultrasound frequency(ies) used will depend on the application in mind. For clinical ultrasound scanners, the centre frequency is preferably in the range 0.1M Hz to 40MHz. For an ultrasound microscopy system, the centre frequency is preferably in the range 5M Hz to 500MHz. The ultrasound waves may have a centre frequency (f_c) of 5M Hz.

The ultrasound waves may be provided in the form of ultrasound pulses. Such pulses may constitute alternating periods in which ultrasound is emitted and not emitted ("on/off pulses"). Alternatively, such pulses may constitute alternating periods of high and low amplitude ultrasound, created by amplitude modulation of the signal to provide a pressure akin to sinusoidal pressure.

A number of factors affect the length of pulse (PL) to be used. The length of each pulse determines the magnitude of the force applied. Account must be taken of the stiffness of the material of interest. Account must also be taken of the length of time that the acoustically responsive particle may take to reach the material of interest. For example, if the surface of interest is an artery wall, account must be taken of the time necessary for the acoustically responsive particles to travel through the blood to the artery wall. For biologically relevant tissue, and clinical applications, palpation-induced deformation occurs very quickly, in the order of a few microseconds or milliseconds. Preferably, the length of each pulse is within the range 1 cycle to 100ms, more preferably 2 cycles to 50ms, more preferably 3 cycles to 20ms, more preferably 4 cycles to 10ms, more preferably 50 cycles to 2ms. Most preferably, ultrasound pulses 1000 cycles to 2ms in length are used.

Repeated bursts of ultrasound pulses may also be used.

Repeated bursts of ultrasound pulses may be used to sample multiple palpation sites. For example, repeated bursts of ultrasound pulses may be used to measure the elasticity of a plurality of different palpation sites, by allowing the acoustically responsive particle(s) to move to a new palpation site between bursts.

Repeated bursts of ultrasound pulses may also be used to allow a previously-used acoustically responsive particle to move away from a palpation site and another acoustically responsive particle to enter the palpation site, for example where there is a flow of such particles across the palpation site.

A suitable pulse repetition frequency (PRF) for use in the present invention depends upon a number of factors. Such factors include how the technology is implemented, the time particles take to flow to another location for use as a palpation source at that other location and the time required to allow new particles to flow into the location. Suitable PRFs include those within the range of 0.1 Hz to 100 Hz, for example l.OHz, 1.5Hz, 2.0Hz, 2.5Hz, 3.0Hz, 3.5Hz, 4.0Hz, 4.5Hz and 5.0Hz, for example 2.5Hz. Very high PFRs may also be used, for example when the same particle is being used for multiple palpations. Suitable PFRs therefore also include those up to 100kHz, for example those within the range of 0.5kHz to 100kHz, for example 10kHz, 20kHz, 30kHz, 40kHz, 50kHz, 60kHz, 70kHz, 80kHz, 90kHz and 100kHz.

The number of pulses (N_p) applied during sonication can be low or extremely high. For example, N_p can range from 1 to 1,000,000, 1 to 4000, 1 to 3000, 1 to 2000, 1 to 1000, 1 to 500 or 1 to 100. Preferably, N_p ranges from 1 to 20, more preferably 1 to 10, more preferably 5 to 10, for example 6.

The peak-negative pressure (p_n) and mechanical index define the magnitude of the force applied. Mechanical index refers to the p_n/sqrt( _c). The pressure applied may be modified depending upon the exact method employed, including the type of acoustically responsive particle used, and the material of interest. Ultrasound waves with a low peak-negative pressure (p_n) and mechanical index (Ml) are preferred. Preferably, the mechanical index is suitable for water and blood. Where the acoustically responsive particle is a microbubble, the mechanical index is preferably relatively low, for example below 2.0, in order to prevent the microbubble from collapsing violently and damaging the material of interest. Where a solid acoustically responsive particle is used, for example a glass bead, a much higher mechanical index may be used.

One or more acoustically responsive particles may be used in the method of the present invention. On their own, these particles do not cause any significant deformation of the material of interest. It is only when these particles are exposed to acoustic waves of a suitable frequency do they cause significant, measurable deformation of the material of interest.

The acoustically responsive particles of the present invention are preferably administered into or adjacent to the material of interest. In other words, the acoustically responsive particles are pre- formed prior to administration into or adjacent to the material of interest and are not formed within the material of interest. This is an advantage over methods in which the acoustically responsive particles used for palpation are generated within the material of interest using destructive techniques, for example those methods in which bubbles are generated within the material of interest using a laser. Such methods are not suitable for use in most clinical situations in which the destruction of tissue is preferably avoided.

The concentration and distribution of the particles used determines the extent to which the force applied to the surface of interest is localised i.e. the size of the palpation site. The size of the palpation site may be in the same order of magnitude as the acoustic wave beam width, but a lower size may be achieved if the acoustically responsive particles are separated to a greater degree.

Where a single acoustically responsive particle is used, the particle will apply highly localised stress that is orders of magnitude smaller than the width of the beam of the acoustic wave, for example the ultrasound beam. As such, it is possible to measure tissue elasticity at sub-wavelength resolution. While more than one acoustically responsive particle may be present in or near the material, this method will be considered "single acoustic particle palpation" if the region in which each particle causes deformation of the material, the "region of palpation" (ROP), does not overlap with the ROP for any other particle. This is likely to occur with particle concentrations and force magnitudes low enough that each particles' force does not influence the tissue deformation caused by neighbouring particles. An example of such a scenario would be the use of low concentrations of acoustically responsive particles in a capillary bed of soft tissue.

Alternatively, multiple acoustically responsive particles, or clusters of acoustically responsive particles, may be used. As used herein, a "cluster" of acoustically responsive particles is a group of such particles that are very close to each other. Preferably, the distance between the particles is less than the diameter of the particles. This usually occurs due to the attractive forces between the particles, either due to intrinsic properties of the particles or attractive forces generated by the acoustic wave acting on the particles.

Where multiple acoustically responsive particles or clusters of acoustically responsive particles are used, the acoustic particle concentration and force magnitude may be high enough such that each particle's force influences the tissue deformation of neighbouring particles. Under such

circumstances, the ROP for each particle overlaps with those of one or more other particles. There is no linear relationship between one palpating particle and two palpating particles if they are very near one another. Multiple and clusters of particles that share the same ROP will collectively generate a single, larger ROP if it is separate from other multiple or clusters of particles. The use of multiple or clusters of acoustically responsive particles can also be used to measure tissue elasticity at sub-wavelength resolution, depending upon the size of the particles use and their

number/concentration.

A "cloud" of acoustically responsive particles may also be used. As used herein, a "cloud" of acoustically responsive particles is a large population of particles, in particular a population of such a high concentration of particles that the attractive forces between the particles begin to influence the behaviour of the population.

Where a cloud of acoustically responsive particles are used, a large number of particles will share the same ROP. In that case, the palpation region will be defined mostly by the size of the ultrasound beam. For example, if there is an ultrasound focal size, then the ROP will be approximately the size of that focus (an order of magnitude between 0.1 and 10 times the focal beam diameter). If the ultrasound is a plane wave, then the ROP will be roughly the width of the plane wave.

Thus, adjusting the type of particles used, and the concentration and distribution of particles used, allows palpation at different scales and resolutions. The smaller the particle used, the better the resolution. The use of multiple acoustically responsive particles and clouds allows for a greater magnitude of palpation, which increases the contrast and sensitivity of the elasticity imaging techniques.

A wide range of concentrations of acoustically responsive particles may be used, from very low to high. For example, the acoustically responsive particles may be used in concentrations of approximately 1 x 10⁵ to approximately lxlO⁸ acoustically responsive particles (ARPs)/ml, preferably approximately 1 x 10^s to approximately lxlO⁸ ARPs/ml, more preferably approximately lx 10⁶to lx 10⁷ ARPs/ml, most preferably approximately 3x10^s ARPs/ml. The deformation magnitude has been found to increase with the concentration of the acoustically responsive particles.

The acoustically responsive particles used must be large enough to enable the application of sufficient force to the material of interest, to produce a measurable deformation when displaced by the acoustic waves. However, the use of acoustic waves with a centre frequency within the resonance frequency range of the particles means that even small particles can produce a large primary acoustic radiation force. The particles are preferably also small enough to enable circulation throughout the material or tissue of a subject, including circulation within capillaries, other small vasculature and the interstitial space. Human capillaries are between 4 and ΙΟμιη in diameter and thus clinical applications of the present invention preferably make use of particles smaller than ΙΟμιη in diameter. As discussed above, it is also preferred that the size of the particles is matched to the frequency of the acoustic waves used to displace them, such that the frequency of the acoustic waves falls within the resonance frequency range of the particles.

Preferably, the diameter of the acoustically responsive particle(s) is within the range Ο.ΟΙμιη to ΙΟΟμιη, more preferably Ο.ΐμιη to ΙΟΟμιη, more preferably Ο.ΐμιη to 50 μιη, more preferably 0.5μιη to 50μιη, more preferably 0.5μιη to 20μιη, more preferably Ιμιη to 20 μιη or 0.5μιη to 15μιη, more preferably Ιμιη to 15 μιη, more preferably Ιμιη to ΙΟμιη in diameter, for example 0.5μιη to 10 μιη, 0.5μιη to 8μιη, or 1 μιη to 5μιη. The diameter of the acoustically responsive particles may be Ιμιη or less, for example Ο.ΟΙμιη to Ιμιτι.

Where the diameter of the acoustically responsive particle(s) is within the range Ιμιη to ΙΟμιη, the frequency of ultrasound used to displace the particles is preferably in the range 0.25MHz to 10MHz.

The population of acoustically responsive particles used may be monodisperse or polydisperse.

Many different types of particle can be used in the method of the present invention, provided that the particles are acoustically responsive i.e. that they are displaced in space upon the application of acoustic waves such as ultrasound. Preferably, the particles have an acoustic impedance mismatch with the material of interest.

Scattering agents such as microbubbles, glass beads and microspheres, including gas-filled microspheres that comprise a rigid shell that does not oscillate, may be used. Microbubbles are particularly preferred. When exposed to acoustic waves such as ultrasound, microbubbles undergo volumetric acoustic oscillations due to their compressibility and scatter the incident wave.

Microbubbles advantageously remain compartmentally contained within most vessels, cavities and other spaces within the human body due to their size. Microbubbles are also stable when exposed to an acoustic wave. The microbubbles preferably comprise a core, more preferably a gas core. Suitable gases include air and stabilised gases. Suitable gases are those that won't rapidly diffuse out of the microbubble when stored in a vial or administered into a fluid body, such as blood. They are also preferably biologically inert. Examples include perfluorocarbons, for example perfluorobutane and

octaf!uoropropane, and sulphur hexafluoride.

The microbubbles for use in the present invention preferably comprise a solid shell to increase the stability of the bubbles in liquid. The stability required depends on the system into which the microbubbles are administered. Where the microbubbles are administered into a subject, for example a human subject, the microbubbles are preferably stable for between 1 and 10 minutes. The shell may comprise one or more lipids, such as l,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC), l,2-dihexadecanoyl-sn-glycero-3-phosphate (DPPA) and/or l,2-distearoyl-sn-glycero-3- phosphoethanolamine-N-[methoxy(polyethylene glycol )-2000] (DPPE-PEG2000), one or more polymers such as albumin and/or one or more proteins such as serum albumin. Lipid-shelled bubbles are preferred.

Suitable microbubbles include those that are currently used in medicine, for example those used as contrast agents in ultrasound imaging. Such microbubbles are routinely administered in clinics via intravenous administration to enhance the echogenicity of vasculature. They have been proposed for use in other cavities, such as the lymphatic system, through subcutaneous injections and via catheters and drains. Many such microbubbles have received clinical approval but have never before been used for tissue elasticity imaging.

Particularly preferred microbubbles include Sonovue^® (sulphur hexafluoride microbubbles), Definity^® (lipid-coated microspheres filled with octafluoropropane gas), Optison™ (Perflutren Protein-Type A Microspheres) and Albunex (air-filled albumin microspheres prepared from sonicated 5% human serum albumin), all of which have received clinical approval in numerous countries for use as contrast agents.

The examples below demonstrate the efficacy of the present invention using lipid-shelled microbubbles with a stabilised gas core and a diameter of 1.32±0.76μιη. A 12μιη displacement was obtained in the presence of microbubbles at a low concentration (approximately 3x10^s

microbubbles/ml) and low peak-negative pressure (approximately 600kPa). The force generated was larger than A F-only techniques and was applied only to the material of interest. In the method of the present invention, the elasticity of a material of interest is determined by imaging, tracking and/or measuring the deformation of the material and/or force response resulting from the displacement of the acoustically responsive particles by the acoustic waves. Suitable methods of measuring the deformation of the material and the force response, and then calculating elasticity from such measurements, are known in the art⁸.

The method of the present invention allows both qualitative and quantitative elasticity values to be provided. Where the acoustic properties and geometry of the acoustically responsive particle is known, the force generated when the acoustic waves are applied can be estimated. A quantitative measure of deformation on-axis can then be derived from this force. This is in contrast to methods making use of acoustic waves alone, where it is only possible to measure quantitative elasticity off- axis.

The position and location of the material of interest or the acoustically responsive particles may be monitored by successive imaging frame capture before, during and after acoustic wave imaging both on-site and off-site to the region of palpation. For example, conventional pulse-echo, plane wave, and pulse-inversion ultrasound imaging may be used to monitor the movement of the acoustically responsive particles. Super-resolution ultrasound imaging algorithms can be used to locate and track individual acoustically responsive particles down to a few micron precision.

The deformation of the material of interest may be monitored spatially and temporally by tracking the movement of the material of interest itself. For example, conventional pulse-echo or plane wave ultrasound may be used to monitor the movement of the material of interest. The deformation of the material may also be imaged using high-speed optical microscopy. One or more images of the surface of interest may be captured before, during and/or after sonication, for example at a plurality of time-points following the start of sonication. The displacement of the surface can then be measured by tracking algorithms, such as cross-correlation.

The deformation of the material may also be measured indirectly or outside the region of palpation. For example, elasticity may be determined by measuring the velocity of shear waves propagated from the material at the palpation site. Many elasticity measurement techniques characterise the shear waves that propagate from the palpation site^{6 13}. Among the diverse properties characterised, the shear wave speed is the most prolif ically measured, because it provides a quantitative estimate of elasticity. Imaging methods relying on compressional waves such as ultrasound can therefore be used to record propagation of shear waves, which propagate with speeds that are several orders of magnitude slower than those of compressional waves. Other methods of indirectly measuring the deformation of the material of interest are known in the art⁸.

Alternatively or in addition, the deformation of the material of interest may be monitored spatially and temporally by tracking the movement of the acoustically responsive particles.

The force response may be monitored spatially and temporally.

Where the acoustically responsive particles are microbubbles, the method of the present invention preferably comprises a contrast-enhanced ultrasound imaging technique.

High-speed optical microscopy, ultrasound imaging, magnetic resonance imaging and other methods may be used to measure the deformation of the material, the force response and the velocity of shear waves propagated from the palpation site and to track the acoustic particle movement.

The elasticity modulus may be calculated from displacement data in a number of ways which are known to the person skilled in the art. The following exemplary methods may be used:

1) On-Axis - Quantitative: If a strain field of the region of interest can be obtained (identifying displacement of some points inside the region) and if the applied force (stress) is known, the elasticity (Young's or shear) modulus can be calculated via stress/strain.

2) On-Axis - Qualitative: If a strain field of the region of interest can be obtained (identifying displacement of some points inside the region) but the applied force (stress) is not known, qualitative elasticity can be measured. This is a map of relative stiffness of one region from other regions if the same force is applied.

3) Off-Axis - Quantitative: If the displacement of region of interest after each excitation can be identified, shear waves propagating inside material can be tracked and shear wave speed calculated. If shear wave speed is known, shear modulus can be calculated via the expression shear wave speed = sqrt (shear modulus/density), where Young's modulus is approximately three times of shear modulus.

The duration of the applied acoustic radiation force excitation may be quasi-static, transient (impulsive) or harmonic⁸. The excitation (pushing) pulses may be applied quasi-statically to achieve a steady-state response, transiently in an impulse-like fashion, or harmonically to excite the tissue at specific frequencies.

The location of the tracking beam used to monitor the deformation response may be within the region of excitation (on-axis) or outside the region of excitation (off-axis).

Suitable tracking algorithms include shear wave dispersion and supersonic shear imaging.

The method of the present invention is particularly suitable for use where the material of interest is a surface of interest. Preferably, the surface of interest is a fluid-structure interface. Under such circumstances, the acoustically responsive particles may be administered into fluid adjacent to the fluid-structure interface.

The method of the present invention has both clinical and non-clinical applications.

One example of a non-clinical application is the measurement of the elasticity of tissue in a tissue scaffold. The method of the present invention can also be used to measure the elasticity of material or a surface of interest in three-dimensional tissue engineering samples, a hydrogel, a microfluidics network, any nano soft machine such as flexible, elastic robots, nanotubes, any environment where the material should be evaluated in-situ such as the materials in M EMS or any other soft material comprising pores, holes, channels or vessels.

The method of the present invention also has clinical applications. Changes in elasticity and stiffness occur in nearly every disease and disorder ranging from cardiovascular diseases to cancer and neurodegeneration and it occurs across the scale from the molecular to tissue level. The method of the present invention can therefore be used to diagnose a disease or disorder in a subject, wherein said disease or disorder is characterised by a change in elasticity of the material of interest. Such diseases and disorders include cardiovascular disease, atherosclerosis¹⁶, fibrosis¹⁷, heart failure¹⁸, cancer¹⁹ and neurodegeneration. When used for a clinical application, it is preferred that the acoustically responsive particle is biocompatible.

The at least one acoustically responsive particle may be administered, pre-formed, into the vasculature, lymphatic system, interstitial fluid or a body cavity of the subject, for example into a blood vessel lumen. The particles are preferably administered parenterally, preferably intravenously. Administration may be via subcutaneous injection, a catheter or drain.

The acoustic waves may then be generated by a source external to the subject and transmitted in the direction of the at least one acoustically responsive particle in the vasculature or body cavity of the subject. The method of the present invention therefore allows the measurement of the elasticity of both superficial surfaces of materials and materials deep within the body. Examples of materials to which this method may be applied include blood vessels, lymph vessels, brain tissue (e.g.

palpation from the cerebrospinal fluid), and the microvasculature. As the vessels decrease in size, the wall thickness decreases and they eventually take on the mechanical properties of the soft tissue organ. Thus, palpation by acoustically responsive particles will effectively measure soft tissue elasticity.

If the acoustically responsive particle is approximately the same size as its surrounding vessel environment, for example the microvasculature or microfluidics chamber, its porous environment, for example the interstitial space, or any other environment, for example a tissue scaffold, then this may no longer be a surface palpation, but simply a localised regional palpation. However, if the concentration is high enough, then all these small sites of palpation will increase their region of palpation, which can overlap to create one large region of palpation.

A second aspect of the present invention provides a device for palpating a material of interest with at least one acoustically responsive particle and measuring the resulting deformation of the material of interest and/or force response, the device comprising: a) a palpation component comprising one or more transducers configured to generate an acoustic wave suitable for pushing at least one acoustically responsive particle against the material of interest to cause transient deformation of the material of interest within a region of palpation; and b) an active monitoring component configured to actively image the material of interest and/or the acoustically responsive particle. Preferably the at least one acoustically responsive particle has been administered into or adjacent to the material of interest. The device of the second aspect is suitable for measuring the elasticity of a material of interest and is preferably suitable for carrying out the method according to the first aspect. All of features of the first aspect apply mutatis mutandis to the second aspect of the present invention and vice versa.

One embodiment of the device of the second aspect of the invention is illustrated in Figure 5. The acoustic hardware is a series of wires, circuits, electronics, etc. that controls the excitation of each transducer. The purpose of this hardware is to control the generation of the ideal pressure field experienced by the acoustically responsive particle while accounting for the distance and medium that it needs to propagate through to get to where the particle is. There should be as many wires/cables/connections between the transducer and the hardware. The hardware may be similar to that of an ultrasound imaging scanner. In most instances, a computer will tell the acoustic hardware what to do.

Transducers, or pistons, generate a pressure wave. In audible sound and ultrasound, the pressure wave is usually generated by a transducer driven by an oscillating voltage. The electrical signal causes the transducer to expand and contract and generates an acoustic wave that propagates through the material.

The acoustic wave generated by the palpation component may comprise either a focused wave or a plane wave. A focused wave may be generated where the one or more transducers comprise a curved transducer surface or wherein the one or more transducers comprise an acoustic lens. A plane wave may be generated where the one or more transducers have a flat or a slightly curved transducer surface.

When considering the need to generate a primary acoustic radiation force, one is concerned with the dimensions of the cross-sectional area of the acoustic beam/field. For a plane wave, this is several orders of magnitude greater than the wavelength of the acoustic wave. For a focused wave, it is around the same order of magnitude as the wavelength.

Preferably, the acoustic wave is an ultrasound wave.

The palpation device may comprise a single transducer. A single transducer has a fixed beam field, although the transducer may be physically moveable using a positioning system. Preferably, the palpation device comprises multiple transducers, for example hundreds or thousands of transducers. This is the preferred configuration, because it has far more capabilities than a single transducer. Multiple transducers can not only generate a single focus or a plane wave, but also multiple foci, incrementally moved foci, and plane waves of different angles and shapes. This means that palpation can be performed in many directions and different locations.

When using multiple transducers, each of the transducers can emit an acoustic wave at a different time to enable a far greater flexibility in the kinds of beam patterns that can be generated.

The transducers can be located in a plurality of configurations to generate the required pressure field to push the acoustic particle. The device may comprise a one-dimensional (ID) array of transducers, for example a one-dimensional circular array of transducers. The device may comprise two-dimensional (2D) array of transducers, for example a two-dimensional circular array of transducers or a two-dimensional spherical segment array of transducers. The transducers can be located on a ID or 2D surface of a probe (similar to an ultrasound imaging probe), the inside surface of a cylindrical apparatus, or the inside surface of a spherical or half-spherical segment (essentially, the inner surface of a bowl or large sphere). Examples of these transducer arrangements are illustrated in Figure 6.

Preferably, the palpation component comprises a hand-held probe with a ID or 2D surface.

The device may additionally comprise a passive monitoring component comprising one or more transducers configured to receive and record the acoustic wave generated by the palpation component, wherein said passive monitoring component is configured to passively monitor the dynamics of the at least one acoustically responsive particle.

The one or more transducers of the passive monitoring component may or may not be the same one or more transducers of the palpation component.

The passive monitoring component enables collection of data on the position, distribution, velocity, acceleration, type and/or magnitude of cavitation activity of the at least one acoustically responsive particle. This can be interpreted through the temporal and/or spectral analysis of the received signals. Passive recording can be made before, during, and/or after the emission of acoustic waves. The entire acoustic emissions data may be recorded and processed by using temporal and spectral methods. The data collected by the passive monitoring device may enable determination of (1) whether a force is being applied or not, (2) the magnitude of the force being applied, and/or (3) safety limits.

The active monitoring component may comprise one or more transducers, and these transducers may or may not be the same one or more transducers as the palpation component and the passive monitoring component.

The active monitoring component may be configured to generate imaging pulses and receive the resulting signal. The term "active" refers to the emission and receiving of acoustic waves for the specific purpose of imaging. These acoustic waves may be in the form of imaging pulses and the active monitoring component may generate acoustic waves with a very different acoustic waveform than the palpation component.

The active monitoring component may be configured to actively image the transient deformation of the material of interest, the force response and/or the acoustically responsive particle dynamics. This active imaging may be performed using a method selected from the group consisting of ultrasound imaging, optical imaging and magnetic resonance imaging.

The position of the material of interest or acoustically responsive particle may be imaged by the active monitoring component by successive imaging frame capture before, during and after generation of the acoustic waves by the palpation component, both on-site and off-site to the region of palpation.

As discussed above in relation to the first aspect of the present invention, advantage can be taken of the resonance behaviour of an acoustically responsive particle to maximise the palpation force. Thus, the acoustic waves generated by the palpation device preferably have a centre frequency within the resonance frequency range of the acoustically responsive particle(s).

Preferably, the device of the second aspect allows control over the magnitude of the force applied by passively and actively monitoring the force applied during palpation. All features of the first aspect of the present invention apply mutatis mutandis to the second aspect of the present invention and vice versa.

The invention is described below with reference to the following examples and figures in which:

Figure 1 is a diagram of the experimental setup used in Example 1. A phantom box containing a wall-less tunnel (diameter: 800 μιη) im m ersed i n a water ta nk was so nicated by a 5 M H z focused u ltrasou nd tra nsducer. U ltrasou nd pu l ses forced m icro bu bbl es aga inst the tissue wa l l to ca use a transient deformation that was monitored by high-speed optical microscopy (right rectangular area).

Figure 2 illustrates the feasibility of acoustic particle palpation. A high-speed camera imaged a small area within the ultrasound region of exposure (see Figure 1). Ultrasound travelled left to right and focused onto a volume that overlapped with a wall-less tunnel phantom (2.5% gelatin) containing microbubbles. Images were acquired (a) 0ms, (b) 0.83ms, (c) 2.50ms, (d) 8.3ms, and (e) 40.8 ms after the start of the sonication (f_c: 5 M Hz, p_n: 625 kPa, PL: 40 ms). (f) The right wal l deformation was tracked with microbubbles present (squares) and without microbubbles present (circles) .

Figure 3 shows displacement at the palpation site, (a) Wall deformation of a wall-less tunnel phantom (5% gelatin) were tracked with microbubbles (squares: top line of graph) and without microbubbles (circles: bottom line of graph) during exposure to ultrasound (f_c: 5 MHz, p_n: 625 kPa, PL: 20 ms, P F: 2.5 Hz, N_p: 6). (b) The maximum displacement was measured and force values were estimated for different acoustic pressures.

Figure 4 shows shear wave propagation away from the palpation site. A wall-less

tunnel in a 2.5% gelatin phantom contained microbubbles and was exposed to ultrasound (f_c:

5 MHz, p_n: 625 kPa, PL: 40 ms). (a) Tissue displacement along the wall occurred within the focal volume and then spread away from the palpation site, (b-f) Difference images depict shear waves generated first on the proximal wall and then along the distal wall at t = 0ms (b), 0.83ms (c), 2.50ms (d), 3.33ms (e) and 4.2ms (f).

Figure 5 illustrates one embodiment of the device of the present invention. Figure 6 illustrates different transducer arrangements for the palpation component, passive monitoring/imaging component and active monitoring/imaging component.

Figure 7 shows displacement of the wall during the time that an excitation would take. Parameters: 1M Hz frequency, 5% gelatin phantom.

Figure 8 shows displacement of the wall over time for different amount of acoustic pressures. Parameters: 3.5 M Hz frequency, 5% gelatin phantom.

Figure 9 shows displacement of the wall during an excitation. Parameters: 5MHz transducer, 5% gelatin phantom.

Figure 10 shows channel wall displacement for different M is. Transducers with frequencies of 1 MHz, 3.5MHz and 5MHz were used. Parameters: 5% gelatin phantom, microbubble concentration of 2X.

Figure 11 shows channel wall displacement for different Mis. Transducers with frequencies of 1 MHz, 3.5 MHz and 5M Hz were used. Parameters: 5% gelatin phantom, microbubble concentration of 20X.

Figure 12 shows displacement of the wall for five different microbubble concentration. Parameters: 5M Hz transducer, 5% gelatin.

Figure 13 shows channel's wall displacement for different microbubble concentrations. The left hand side figures were obtained from 10% gelatin phantom and the right hand side figures show the results from 5% gelatin. To be able to compare the different frequencies, each figure is plotted for a certain M l.

Figure 14 shows the effect of phantom stiffness on the displacement of the wall. Parameters for left figure: 1 MHz center frequency, 7.5% gelatin, 10X microbubble solution. Parameters for right figure: 5MHz center frequency, 7.5% gelatin, 10X microbubble solution. Examples

The m icrobubble manufacturing process, experimental hardware and data analysis used in the following examples are described in the Supplementary Material, set out below.

Example 1: Feasibilty of acoustic particle palpation

Method

A gelatin-based tissue-mimicking material containing a wall-less tunnel (diameter: 800 μιη, 2.5% gelatin) was immersed in a water tank. Lipid-shel led m icrobubbles with a stabilised gas core ( d ia m ete r: 1.32±0.76 μιη , co n ce nt ratio n a p p roxi m ate ly 7xl0⁷ M Bs/ml)) were a d m i n iste red i nto th e t u n ne l so th at th ey were compartmentally separate from the surrounding material (see Figure 1).

Focused ultrasound pulses were emitted from a single-element transducer [centre frequency (f_c): 5 M Hz, peak-negative pressure (p_n): 625 kPa, pulse length (PL): 40 ms], which was driven by a function generator through a power amplifier. The centre frequency was selected to match the resonance size of the microbubbles in order to maximise the generated force.

Deformations of the tunnel wall were measured with high-speed optical microscopy (frame rate: 1.2 kHz), with and without m icrobubbles.

Results

Prior to sonication, the microbubbles were distributed uniformly throughout the tunnel (see Figure 2(a)). Application of ultrasound stimulated the acoustic particles to move through the fluid, accumulate on the distal tissue surface, and deform the surface (see Figures 2(b)-(c)). Removal of the acoustic field allowed the tissue to return to its normal geometry (see Figure 2(e)). Without the presence of microbubbles in the tunnel, no (or low) tissue deformation was observed (see Figures 2(f)-(j))- The progressive change in optical contrast distribution is due to the movement of microbubbles. Ultrasound exposure produced a large displacement (100.6±4.6 μιη) in the presence of microbubbles (t: 0-6 ms) at the beginning of the pu lse length (see Figu re 2(k)) . As the m icrobu bbles moved away from the focus, the displacement decreased ( t: 6-16 ms). As the microbubble concentration at the focus decreased further, the displacement was reduced further (t: 16-40 ms) until it returned to normal when ultrasound was turned off ( t: 40-60 ms). Example 2: Physiological relevance of acoustic particle palpation

Method

In order to evaluate the relevance of the above-described technique to biological tissue, a stiffer phantom material was used (5% gelatin) with a Young's modulus (approximately 1.5 kPa) similar to liver and a lower microbubble concentration near the clinically recommended dose (approximately 3x10^s MBs/ml) that was made to flow through the tunnel using a syringe pump (flow rate: 1 ml/minute). The deformation of a wall-less tunnel exposed to ultrasound [p_n: 625 kPa, PL: 20 ms, pulse repetition frequency (P F): 2.5 Hz, number of pulses (N_p): 6] was determined for microbubbles and water only (see Figure 3).

Results

Without the presence of microbubbles in the tunnel, low tissue displacement (<1.5 μιη) was observed. However, when microbubbles were administered, a higher net displacement (e.g., 11.8 ± 3.3 μιη at t = 5.83 ms) was observed in the direction of wave propagation (see Figure 3(a)). The deformation magnitude increased with peak-negative pressure (see Figure 3(b)). The force magnitude was estimated based on the elastic properties of the material and deformation value was estimated by assuming the elastic medium to be isotropic, homogeneous, incompressible, and inviscous (see Supplementary Material, below). An estimated force of 10 μΝ was obtained for a pressure of 800 kPa when microbubbles are used, which has been previously shown sufficient to create a deformation that is detectable using ultrasound imaging methods⁸.

Example 3: Measurement of shear waves

Method

A gelatin-based tissue-mimicking material containing a wall-less tunnel (diameter: 800 μιη, 2.5% gelatin) was immersed in a water tank. Lipid-shelled microbubbles with a stabilised gas core (diameter: 1.32±0.76 μιη, concentration approximately 7xl0⁷ MBs/ml) were administered into the tunnel so that they were compartmentally separate from the surrounding material (see Figure 1). Focused ultrasound pulses were em itted from a single-element transducer [centre frequency (f_c): 5 M Hz, pulse length (PL): 40 ms], which was driven by a function generator through a power amplifier.

In order to characterise the shea r wave propagation, the displacem ent was measu red a long the tu nnel wa l l for 2.5% gelatin phantom when exposed to the u ltrasound in the presence of microbubbles (See Figure 4(a)).

Results

Immediately after ultrasound was applied, shear waves were observed to propagate away from the focal volume (see Figure 4(a)), which was also observed using difference images (see Figure 4(b)). The shear wave velocity and elasticity modul us were calculated to be v_s = 0.39 ± 0.03 m/s and E = 0.46 ± 0.06 kPa for a 2.5% gelatin phantom and v_s = 0.71 ± 0.07 m/s and E = 1.54 ± 0.32 kPa for a 5% gelatin phantom. These values are in good agreement to values in the literature (i.e., E = 0.1-0.5 and 0.8-2.9 kPa for 2.5 and 5% gelatin, respectively) ^14,15. No clear shear wave generation was observed in the control experiments for either the proximal or distal walls, thus indicating that palpation by A F alone using the same ultrasound parameters is insufficient.

Example 4: Identification of ultrasound parameters and physiological factors influencing acoustic particle palpation

The ultrasound parameters that can generate a physiologically relevant elasticity imaging system were investigated and the results of this investigation are discussed below.

The main deliverable was to evaluate the impact of four different parameters on the outcome of the acoustic particle palpation technique. The four parameters are: acoustic pressure, centre frequency of the ultrasound, microbubble concentration, and stiffness of the tissue material. These parameters were chosen due to their significance on the outcome of this new palpation method. Ideally, the acoustic pressure should be minimal to decrease the tissue heating during palpating. However, the pressure must be enough to lead to an acceptable deformation of the tissue so the displacement can be distinguished from noise and can be tracked by available tools. Tissue mimicking phantoms were prepared in a range of stiffness to investigate the ability of the proposed method to deform tissues with different stiffness. The microbubbles apply a higher force in deformation in comparison with applying acoustic radiation force to the tissue directly. However, low concentration of the microbubbles is desirable. Finally, the effect of different ultrasound frequencies is studied to explore the possibility of using lower frequencies to increase the depth of diagnosis. Additionally, if the bubbles are sonicated with their resonance frequency, they will apply a higher force due to their higher oscillations. Therefore, using a transducer with frequency close to the resonance frequency of the bubbles might result in getting the same results but with lower acoustic pressure and lower microbubbles' concentration.

Tissue mimicking phantoms made from gelatin were used to study the acoustic particle palpation. A wall-less 800 μιη channel was created in the phantom to administer a flow of M Bs inside. The effect of radiation force frequency was studied by using different centre frequency transducers of 1M Hz, 3.5M Hz and 5MHz. Pulse durations of 20 msec, 10 msec and 10 msec were determined for each transducer respectively which were monitored after some initial experiments to be high enough to see the displacements. Five different microbubble diluted solutions were tried to observe the effect of the microbubble concentration on the displacement of the wall. On top of that, each transducer was derived in a range of acoustic pressures to see the outcome of various pressures on. The channel displacement was tracked optically by a high frame rate camera with the channel's image resolution of about 10 μιτι/pixel. Cross-correlation method was used to measure the displacements. The displacement value is calculated by comparing the intensity values of the pixels in frames where the excitation has occurred and the frames before the excitation.

The experiments were conducted numerous times. However, the results that are presented are from one video each but have averaged for ten consecutive excitations. The best videos were chosen for each experiment to extract the important data. However, due to time limitation, the experiment with 3.5 MHz transducer were conducted only once for two phantoms with 5% and 10% gelatin. Therefore, there was no options for the videos in using 3.5 M Hz transducer to select between them. It is also important to mention that all the values for the pressures are related to the peak to peak acoustic pressure unless otherwise stated.

Acoustic Pressure

Different voltages were applied to each ultrasound transducer. The transducers have different sensitivities, which leads to a different acoustic pressure. Based on the values of the voltage, the value of acoustic pressure was calculated. In Figure 7, displacement over time is depicted for different acoustic pressures. The centre frequency, microbubbles concentrations and phantom stiffness was kept constant. The values for the acoustic pressure were 0.71 and 1.06 MPa. Initially, the experiment was conducted with water inside the channel. Although the pressure was 0.71MPa, no displacement of the channel's wall was observed. However, administrating microbubbles with concentration of 5X leaded to an observable displacement of the channel. According to the Figure 7, increasing the acoustic pressure, increases the displacement and the maximum displacement of 13.91±2.97 μιη at 1.06MPa was obtained. The same experiment was repeated but with different transducers.

Figure 8 shows the displacement of the wall versus time for different acoustic pressures ranging from 0.65 to 2.79 MPa under condition of centre frequency of 3.5M Hz, constant microbubbles concentration and phantom stiffness. No displacement was observed for control experiment. It was expected that the general increase in pressure would increase the acoustic pressures however, the maximum displacement of 27.60±10.61 μιη was obtained by applying 1.87 MPa pressure and for higher values of acoustic pressure, the maximum displacement was lower than that of

1.87MPa. The shape of the displacement during an excitation was derived from being symmetrical for high pressures (above 1.87 MPa)

The effect of 5MHz frequency was also investigated and the results for acoustic pressures of 1.23 and 3.62 MPa are shown in Figure 9.

The expected general trend of increase in the amount of displacement during excitation by increasing the amount of acoustic pressure was also seen for 5MHz transducer. No deformation for the control experiment was observed. The maximum displacement of 16.17±10.31 μιη was acquired in the experiment by applying 3.62 MPa acoustic pressure although the peak went higher by increasing the pressure.

Centre Frequency

Three different ultrasound transducers were used each with a different centre frequency. The transducers had frequencies of lMHz, 3.5MHz and 5MHz. These different transducers were chosen to be close to, less than and higher than the resonant frequency of the microbubbles. The resonant frequency of the microbubbles are approximately 3M Hz. This value was calculated based on the mean radius of the used microbubbles. Therefore the 3.5MHz transducer was nearly matching the resonance frequency, and the other two transducers were above and lower than this frequency.

The displacement of the wall is shown for three frequencies based on their Mechanical index (Ml) in Figure 10 and Figure 11. In both experiments the gelatin concentration was 5%. The microbubble concentration for Figure 10 and Figure 11 are 2X and 20X correspondingly. Each frequency is depicted using different shapes where the squares shows the results for lMHz, the circles for 3.5MHz and the triangles correspond to the results for 5MHz transducer. The results are shown for different Mis not for acoustic pressures since the range of acoustic pressures in which each transducer was driven is different from the others. Generally according to the figures, in a phantom for a certain Ml below 0.42, the displacement which was obtained by using 3.5M Hz is the highest among the others. Then the amplitude of displacements that 1MHZ had generated in the wall are placed although 5MHz generated the lowest displacements. However, the displacements obtained for 3.5MHz were dramatically decreased by increasing the Ml above 0.58.

Microbubbles concentration

The concentration of microbubbles was changed to observe the effect of different concentrations and also investigate the possibility of palpating the tissue for an acceptable deformation with low bubble concentration. Assuming the clinical dose as 2.04x10^s MBs/ml, microbubbles were diluted with certain amount of distilled degassed water to prepare different concentrations. Five different solution of IX, 2X, 5X, 10X and 20X were used from the lowest to the highest.

Figure 12 was given by applying various acoustic pressures with 5M Hz transducer to the tissue mimicking phantom with 5% gelatin concentration. The transducer was drived at 2.42 MPa, 3.62 MPa and 4.22 MPa in turn. Accordingly, using higher microbubble concentrations resulted to increase in displacement however; the relationship between them is not linear.

In Figure 13, displacement over microbubbles concentration is depicted for different centre frequencies. The right side figures are related to the phantom with 1% gelatin concentration and the left hand ones were tested on 10% gelatin phantom. For the first row of images, the Ml is 0.53 which corresponds to acoustic pressures of 1.06 MPa, 1.87 M Pa and 2.42 MPa for 1M Hz, 3.5MHz and 5MHz respectively. The Ml is 0.42 in the second row and 0.33 in the third row. Generally, it can be seen that for a given microbubble concentration and constant Ml, 3.5 MHz transducer produces higher displacement among the others. However, the effect of this transduces has decreased for the M is above 0.58.

Stiffness of tissue material

Tissue mimicking phantoms made from gelatin were used to test the new palpation method. Phantoms were prepared in three different concentrations: 5%, 7.5% and 10%. It was expected that the higher in concentration the gelatin, the stiffer it would be.

In Figure 14, the effect of phantom stiffness on the amount of wall deformation is shown. Centre frequency, microbubbles concentrations and phantom stiffness was kept constant for each experiment. The values for the acoustic pressure ranged from 0.6 to 1.06 for the lMHz transducer and 1.23 to 2.42 MHz for 5 MHz transducer. It can be seen that the stiffer the phantom, the less the deformation under equal conditions. The results are in accordance with the expectations since high stiffness material would resist to deformation more than the looser one.

Supplementary Information

1. Microbubble manufacturing

Lipid-shelled microbubbles were manufactured following previously described protocols²¹. The microbubble shell consisted of three phospholipids originally in powder form (Avanti Polar Lipids Inc, AL, USA): l,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC), l,2-dihexadecanoyl-sn-glycero-3- phosphate (DPPA), and l,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(polyethylene glycol)-2000] (DPPE-PEG2000). They were mixed in proportions of 82%, 8%, and 10% respectively. The lipid powders were diluted with saline and glycerol. Perfluorobutane (FluoroMed L.P., Texas, USA) was introduced to the solution and then the mixture was mechanically amalgamated for 45 s. The resultant microbubble (MB) concentration was measured to be 3.25xl0⁹ MBs ml^"1 and the mean diameter was 1.32 ± 0.76μιη. The high concentration used in our study refers to ~7xl0⁷ M Bs ml^"1, which is approximately 30 times greater than the clinically recommended dose of 0.02 ml kg^"1 or 2x10^s MBs ml^"122. This clinical dose was calculated for an average body mass of 70 kg with a blood volume of 5 L²¹. The physiologically relevant concentration in this paper was ~3xl0⁶ MBs ml^"1. 2. Experimental hardware and data analysis

A 5 MHz spherical-segment, single-element transducer (diameter: 25.4 mm, focal length: 51 mm; Olympus Industrial, Essex, UK) was used to emit focused ultrasound pulses. The full width at half maximum (FWHM) focal diameter and length of the ultrasound beam was 2 and 20 mm, respectively (Fig. 1). Pressure calibrations, focal point location, and FWHM beam measurements were performed in a degassed water tank using a 200 μιη diameter polyvinylidene fluoride (PVDF) needle hydrophone (Precision Acoustics Ltd, Dorchester, UK). The wall-less tunnel in the phantom was placed 51 mm away from the transducer's surface and overlapped with the transducer's focal point. A function generator (33500B Series, Agilent technologies, Santa Clara, CA, USA) produced pulses that were transmitted through a 50 dB power amplifier (Precision Acoustics Ltd, Dorchester, UK) and to the 5 MHz transducer. Material deformation was captured with a high-speed camera (frame rate: 1.2 kHz, field of view: 416x144 pixels, model: Nikon 1 V3, Nikon Inc., USA). The camera was used with a lens (focal length: 10-lOOmm, zoom ratio: 10:1, model: 1 NIKKO 10-lOOmm f/4.0-5.6 VR, Nikon Inc., USA) and two magnifying glasses (zoom ratio: 6:1). The phantom was made by pouring 1.2%, 2.5%, or 5% gelatin (Fisher Scientific UK Ltd, Loughborough, UK) into a box containing an 800- μιη diameter tunnel. The phantom was left to solidify and then the box was submerged and positioned vertically in a tank containing deionized and degassed water. Microbubbles were injected into the tunnel using a syringe pump (Instech, Plymouth Meeting, PA, USA), which was set to a constant flow rate of 1 ml min^"1, corresponding to a fluid velocity of 33 mm s^"1.

The tunnel wall deformation was measured by image processing of the captured video data. The raw data was interpolated by a factor of 10 before 1-D cross-correlation at the distal material wall and along the axial direction was conducted as a function of time. The spatial resolution was 1.5-2.0 μιτι/pixel. The standard deviations were calculated using the response of the six pulses where the number of cycles was 100,000 (pulse duration: 20 ms) and the number of pulses emitted was 6 for the 5% gelatin phantom experiments (the number of cycles was 200,000 -pulse duration: 40 ms- for 2.5% gelation phantom experiments). Wall displacement measurements were repeated laterally along the entire length of the distal wall. Shear wave velocities ( v_s ) were calculated by measuring the distance travelled by the shear wave at each frame. A threshold was set to measure when the wave arrived at the corresponding point. The Young's modulus ( E ) of the material was calculated using v_s = JE /3p where p is the density of material assumed to be incompressible (i.e., Poisson's ratio: 0.5)²³. In order to visualise the spatio-temporal distribution of the shear waves generated at the proximal and distal vessel walls, we subtracted each image frame by the previous frame and then removed data from within the vessel.

3. Force estimation

In order to evaluate the relevance of this technique to different materials, we estimated the force generated by acoustic particle palpation at different acoustic pressures. The exerted force was estimated by using the Young's modulus calculated by the shear wave speed (Fig. 4) and the displacement of the material captured by our camera. In a previous study and for a different application, the force applied on a soft material was estimated theoretically by calculating the displacement of a gas sphere in an isotropic and incompressible elastic medium with linear approximations where tangential stress at the surface of the sphere is assumed to be zero²⁴. In this study, the displacement-force relationship was approximated for a fictive sphere at the interface of a bulk material and a liquid. In this case, a part of the sphere is applied onto the bulk material and the force is integrated over the immersed region of the sphere defined by the circle segment and its associated angle 2Θ = 2 - cos ¹ (l - δ I R) where R is the radius of the sphere and δ is the displacement of the bubble inside the bulk material. The polar axis of the spherical coordinates system coincides with the direction of the displacement. Assuming the pressure changes with P_r ⁼ Po cos(#) where p₀ is the pressure for Θ = 0 , its value is defined by the radiation force given by:

Evaluating the integral above gives p₀ as:

Using the equilibrium equation,³ the force expression is as follows:

where E is the Young's modulus of the material. A single fictive sphere with a radius R equal to the lateral radius of the focal volume of the transducer was assumed for our displaced microbubble cloud.

To further test the validity of this model, we implemented the Hertz theory which describes the contact mechanics between an elastic sphere and an elastic half-space, which stood in our case for the microbubble cloud and the gelatin/channel interface respectively. In a simplified Hertz model, the displacement induced by the elastic sphere is correlated with the indentation force as:⁵

where R is the radius of the elastic sphere, d is the induced displacement and E* is the reduced Young's modulus, computed as:

1 / E ⁼ {^ ^{~ V}sphere ) ^ ^sphere ⁺ {^ ^{~ V}half-space ) ^ ^half-space (^) where v is the Poisson's ratio, here equal to 0.5 for incompressible materials. Assuming that the Young's modulus of the elastic sphere (here the bubble cloud) is far larger than the one of half-space (here the gelatin/tunnel interface), i.e. E_sphem » E^^^ , the reduced Young's modulus was approximated as ?* = l.3E_gelatm . Both approaches gave indentation force values on the order of μΝ, thus confirming the validity of the estimated forces range.

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Claims

1. A method of measuring the elasticity of a material of interest, the method comprising: a) palpation of the material of interest using at least one acoustically responsive particle which has been administered into or adjacent to said material, wherein said palpation comprises the use of acoustic waves to push the at least one acoustically responsive particle against the material of interest to cause transient deformation of the material; and

b) measuring the transient deformation of the material and/or the force response.

2. The method of claim 1 wherein the acoustically responsive particles are displaced by the acoustic waves at a higher magnitude than the material of interest.

3. The method of claim 1 or claim 2 wherein the centre frequency of the acoustic waves is within the resonance frequency range of the at least one acoustically responsive particle.

4. The method of claim 3 wherein the centre frequency of acoustic waves matches the

resonance size of the at least one acoustically responsive particle.

5. The method of any preceding claim wherein said acoustic waves are ultrasound waves.

6. The method of claim 5 wherein said ultrasound waves are ultrasound pulses.

7. The method of any preceding claim wherein the at least one acoustically responsive particle is selected from the group consisting of a single acoustically responsive particle or a cluster of acoustically responsive particles.

8. The method of any of claims 1 to 6 wherein the at least one acoustically responsive particle is selected from the group consisting of multiple acoustically responsive particles or a cloud of acoustically responsive particles.

9. The method of any preceding claim wherein said at least one acoustically responsive particle is 1 to ΙΟμιη in diameter.

10. The method of claim 9 wherein said acoustic waves comprise ultrasound of frequency 0.25 to 10M Hz.

11. The method of any preceding claim wherein the at least one acoustically responsive particle is a microbubble.

12. The method of claim 11 wherein said microbubble comprises a lipid or polymer shell.

13. The method of claim 11 or claim 12 wherein said microbubble comprises a core and wherein said core comprises air or a stabilised gas.

14. The method of any of claims 1 to 10 wherein the at least one acoustically responsive particle is a scattering agent.

15. The method of any preceding claim wherein the measurement of the transient deformation of said material of interest comprises measuring the velocity of shear waves propagated from the surface.

16. The method of any preceding claim wherein the measurement of the transient deformation of the material of interest comprises the use of one or more of optical microscopy, ultrasound imaging and magnetic resonance imaging.

17. The method of any preceding claim, wherein the material of interest comprises a surface of interest and wherein said method comprises palpation of the surface of interest and measurement of the transient deformation of said surface.

18. The method of claim 18 wherein the surface of interest is a fluid-structure interface.

19. The method of claim 18 wherein the acoustically responsive particles are administered, preformed, into fluid adjacent to said fluid-structure interface.

20. The method of any preceding claim wherein the material of interest is selected from the group consisting of tissue in a tissue scaffold, engineered tissue, a hydrogel, a microfluidics network and any soft material comprising pores, holes, channels or vessels.

21. The method of any preceding claim wherein the measurement of the elasticity of the material of interest is used to diagnose a disease or disorder in a subject, wherein said disease or disorder is characterised by a change in elasticity of the material of interest.

22. The method of claim 21 wherein the disease or disorder is selected from the group

consisting of cardiovascular disease, atherosclerosis, fibrosis, heart failure, cancer and neurodegeneration.

23. The method of claim 21 or 22 wherein the at least one acoustically responsive particle is injected into the vasculature or a body cavity of said subject.

24. The method of claim 23 wherein the acoustic waves are generated by a source external to the subject and are emitted in the direction of the at least one acoustically responsive particle in the vasculature or body cavity of said subject.

25. The method of claim 23 or claim 24 wherein the material of interest is a vascular or

microvascular wall.

26. A device for palpating a material of interest with at least one acoustically responsive particle and measuring the resulting deformation of the material of interest and/or force response, the device comprising:

a) a palpation component comprising one or more transducers configured to generate an acoustic wave suitable for pushing at least one acoustically responsive particle against the material of interest to cause transient deformation of the material of interest within a region of palpation; and

b) an active monitoring component configured to actively image the material of interest and/or the acoustically responsive particle.

27. The device of claim 26, additionally comprising a passive monitoring component comprising one or more transducers configured to receive and record the acoustic wave generated by the palpation component, wherein said passive monitoring component is configured to passively monitor the dynamics of the at least one acoustically responsive particle, wherein the one or more transducers of the passive monitoring component may or may not be the same one or more transducers of the palpation component.

28. The device of claim 26 or claim 27 wherein the active monitoring component comprises one or more transducers, wherein the one or more transducers of the active monitoring component may or may not be the same one or more transducers of the palpation component and the passive monitoring component.

29. The device of any of claims 26 to 28 wherein the acoustic wave is a focused wave.

30. The device of claim 29 wherein the one or more transducers comprise a curved transducer surface or wherein the one or more transducers comprise an acoustic lens.

31. The device of any of claims 26 to 28 wherein the acoustic wave is a plane wave.

32. The device of claim 31 wherein the one or more transducers have a flat or a slightly curved transducer surface.

33. The device of any of claims 26 to 32 wherein the acoustic wave is an ultrasound wave.

34. The device of any of claims 26 to 33 wherein the palpation device comprises a single

transducer.

35. The device of claim 34 wherein the single transducer can be moved using a positioning system.

36. The device of any of claims 26 to 33 wherein the palpation device comprises a plurality of transducers.

37. The device of claim 36 wherein each of the plurality of transducers emit an acoustic wave at a different time.

38. The device of claim 36 or 37 wherein the palpation device comprises a one-dimensional array of transducers.

39. The device of claim 38 wherein the palpation device comprises a one-dimensional circular array of tranducers.

40. The device of claim 36 or 37 wherein the palpation device comprises a two-dimensional array of transducers.

41. The device of claim 40 wherein the palpation device comprises a two-dimensional circular array of transducers.

42. The device of claim 41 wherein the palpation device comprises a two-dimensional spherical segment array of transducers.

43. The device of any of claims 26 to 42 wherein the palpation component comprises a handheld probe comprising the one or more transducers.

44. The device of any of claims 27 to 43 wherein the passive monitoring component enables collection of data on the position, distribution, velocity, acceleration, type and/or magnitude of cavitation activity of the at least one acoustically responsive particle.

45. The device of any of claims 26 to 44 wherein the active monitoring component is configured to generate imaging pulses and receive the resulting signal.

46. The device of any of claims 26 to 45 wherein the active monitoring component is configured to actively image the transient deformation of the material of interest, the force response and/or the acoustically responsive particle dynamics.

47. The device of claim 46 wherein the active imaging of the transient deformation of the

material of interest, the force response and/or the acoustically responsive particle dynamics is performed using a method selected from the group consisting of ultrasound imaging, optical imaging and magnetic resonance imaging.

48. The device of any of claims 26 to 47 wherein the position of the material of interest or acoustically responsive particle is imaged by successive imaging frame capture before, during and after generation of the acoustic waves by the palpation component, both on-site and off-site to the region of palpation.