1.

Introduction

Shear wave elasticity imaging (SWEI) has been widely used for mapping elastic properties of soft tissues¹ and remains a very active research area. It is well known that the viscoelastic properties of soft tissues can be indicative of overall tissue health, and their changes might indicate development of certain diseases and lesions.^2,3 Recent progress has enabled high-quality quantitative SWEI of deep tissues.⁴ Overall, SWEI is based on the excitation of one or several shear waves in tissues with consequent measurements of their propagation speed. Assuming the density of the soft tissue, the shear elastic modulus of tissue traversed by the shear wave can be reconstructed.⁵ This approach is employed in numerous applications in biomedical engineering and medicine, where assessment of biomechanical parameters of soft tissues at microlevel with high spatial resolution is required.^6–8

There are several excitation and imaging approaches capable of shear modulus assessment of soft tissues, for example, mechanical extensiometry,⁹ magnetic resonance elastography (MRE),¹⁰ and ultrasound elastography.³ However, some of these approaches, such as mechanical extensiometry, are destructive and, therefore, applicable for ex-vivo tissue assessment only. MRE is a nondestructive imaging technique but it mostly operates on large-scale parts of tissues and could not be used to assess tissue elasticity on a micrometer level and with high resolution. Although a micro-MRE imaging technique has been reported recently, the approach requires tissue preparation and could not be used in-vivo.¹¹ Even ultrasound elastography may not be suitable for this application because very small displacements caused by shear wave propagation are expected in relatively stiff tissues, such as heart muscle.

Recently, optical coherence elastography (OCE) was proposed for nondestructive, high-resolution SWEI of soft tissues at the micrometer scale.^12–15 This approach is based on generating a shear wave in a sample with consequent measurement of the shear wave propagation speed using phase-sensitive optical coherence tomography (psOCT) where the axial component of particle velocity is measured accurately using phase changes of the spectral interferograms. This makes it possible to detect tissue displacement well below 1 mm—detection of $\sim 10\text{-}\mathrm{nm}$ displacements has been reported using phase-sensitive measurements.¹⁶ Shear waves in tissues can be created using various methods, such as a laser pulse,¹⁷ mechanical actuator,¹⁸ air-pulse system,¹⁹ ultrasound,²⁰ and the Lorentz force.²¹ The speed of shear wave propagation can be calculated as a ratio of a distance between the excitation and observation points (i.e., travel distance) to the time of shear wave propogation.²² Alternatively, shear wave velocity measurements can be performed by changing the distance between the point of excitation and the point of observation.¹⁹ For example, using an air-pulse system to create a shear wave and psOCT system to measure displacement of phantom surface at different distances away from the point of excitation, noncontact assessment of Young’s modulus was in good agreement with the results of standard uniaxial testing.²³

Overall, feasibility of OCE to assess microelasticity of soft tissues has been successfully demonstrated suggesting that OCE may become a powerful and informative tool for preclinical and clinical applications. However, in-vivo measurements of biomechanical properties of internal tissues or measurements in areas where access is limited require a special probe. Such probe should be as miniature as possible, ideally comparable with endo- or intravascular devices, such as OCT catheters²⁴ or devices used in laparoscopic surgery. In addition, to avoid motion artifacts, this probe should allow for fast and preferably noncontact measurement in tissues of moving organs, such as heart or tissues subjected to motion due to the respiratory cycle. To overcome possible motion artifacts, all measurements of shear wave propagation must be acquired within a short segment of the cardiac cycle and respiratory cycle when the tissues are effectively stationary. Generally, cardiac rates are several times greater rather than respiratory rates. Therefore, in small animals, all measurements should be completed within a short ($<100\mathrm{ms}$) data acquisition window at the end of expiration in a ventilation cycle and triggered by $R$ peak in QRS complex within the cardiac cycle.²⁵

While OCT detection of shear waves can be performed using miniature probes,²⁶ limited access to internal tissues effectively narrows the choice of shear wave excitation methods. An ultrasound transducer can be used to generate an acoustic radiation force (ARF) and create shear waves in tissues.^27,28 Therefore, an array of transducers can be integrated with a small OCT probe such that the shear wave observation point (OCT beam) and shear wave excitation points (ultrasound beams) are located along a straight line. Multiple shear waves can be created at different points in tissue simultaneously or one right after another using different ultrasound transducers while psOCT can be used to detect the shear waves. If shear waves are created simultaneously by an array of ultrasound transducers, the measurement can be performed quickly thus avoiding motion artifacts leading to errors in shear modulus reconstruction. Indeed, assuming a very soft tissue with relatively low speed of shear wave of $0.5\mathrm{m}/\mathrm{s}$, propagation of shear waves over 5-mm distance can be measured within a 10-ms window.

The multitransducer ultrasound ARF/psOCT probe would be useful for many in-vivo applications, for example, to assess the microelasticity of cardiac muscle that is damaged by myocardial infarction (MI).²⁹ Indeed, tissues damaged by MI exhibit elevated stiffness compared to that of benign cardiac muscles,^23,30,31 and development of new therapeutic methods of heart self-renewal requires in-vivo monitoring of new cardiomyocyte delivery and proliferation in cardiac muscles.³²

The goal of the paper is to investigate the feasibility of miniature ultrasound transducers to create shear wave displacements measurable by psOCT, and then to design a multielement ultrasound transducer probe and prototype a combined ARF/psOCT probe capable of assessing microelasticity properties of cardiac tissues.

2.

Materials and Methods

2.1.

Design and Operation of Ultrasound Transducers

To test the capability of small ultrasound transducers to create shear waves that are detectable by a psOCT system, a miniature ultrasound transducer was designed and manufactured. The diagram and microphotograph of the transducer are shown in Fig. 1(a). This transducer is comprised of a lead zirconate titanate (PZT) crystal with two thin vacuum sputtered nickel electrodes (T120-A4E-602, Piezo Systems, Inc.). The thickness of the $1.5\mathrm{mm}\times 2.5\mathrm{mm}$ rectangular crystal was 0.020 in. ($\sim 0.5\mathrm{mm}$). The backing was made out of electrically conductive epoxy (E-Solder 3022, VonRoll Insulation, Inc.). Metal housing made out of brass tube (#8126, Precision Metals, Inc.) with outer and inner diameter of 2.38 and 1.7 mm, respectively, was used both as the transducer holder and as the shield for the 32-gauge center conductor located inside the housing.

Fig. 1

(a) Diagram and microphotograph of the miniature ultrasound transducer. The scale bar is 2 mm. (b) Alignment of the ultrasound transducer and an OCT intravascular catheter. Thick red and thin blue arrows indicate the ultrasound and OCT beams, respectively. The scale bar is 2 mm.

At the tip, the upper half of the metal housing was removed over a length of $\sim 5\mathrm{mm}$. The remaining lower half of the housing was slightly flattened such that its internal width in this area was 2.7 mm, where the rectangular PZT crystal with backing was glued using the conductive epoxy as shown in Fig. 1(a). The center wire was electrically and mechanically fixed to the top crystal electrode using the same epoxy. The housing and center wire were connected to an amplifier via coaxial cable. The center frequency of the transducer, measured using a vector network analyzer (E5061B, Keysight Technologies, Inc.) was 4.8 MHz.

2.4.

Integrated ARF/psOCT Probe

We have designed and built a multielement ultrasound transducer probe as shown in Fig. 2(a). Two $1.5\mathrm{mm}\times 1.5\mathrm{mm}$ ultrasound transducers were glued at the tip of the metal housing such that surfaces of the crystals were parallel. The distance between centers of the transducers was 1.65 mm. Each transducer was connected to an amplifier using a 32-gauge signal wire. The fully assembled ultrasound probe was coated by a thin layer of silicone to provide electrical isolation.

Fig. 2

(a) Microphotograph of the multitransducer ultrasound array with transducer size of $1.5\mathrm{mm}\times 1.5\mathrm{mm}$ square. The scale bar is 2 mm. (b) A diagram of the assembled multitransducer ARF/psOCT probe. (c) A diagram of phantom experiments. (d) Photograph of the ARF/psOCT probe positioned above the rabbit heart. Scale bar is 2 mm.

The diagram of the assembled ARF/psOCT probe is shown in Fig. 2(b). The same OCT imaging probe was integrated with the two-element ultrasound transducer probe into a single device. The tip of OCT probe was positioned as close to the ultrasound transducers as possible. Furthermore, the ultrasound transducers and the point where OCT beam entered the tissue were aligned along a straight line.

3.

Results and Discussion

A typical psOCT M-mode image of phantom is shown in Fig. 3(a). The shear wave displacement profile reconstructed from the M-mode image is shown in Fig. 3(b). Ultrasound excitation to generate shear wave is depicted by a shaded area confined by two vertical solid lines. Clearly, a $1.5\mathrm{mm}\times 2.5\mathrm{mm}$ ultrasound transducer can create a shear wave with the peak-to-peak displacement magnitude of 150 nm, which can be reliably detected by psOCT imaging.

Fig. 3

(a) Typical M-mode psOCT image of a phantom. The vertical and horizontal scale bars are $100\mu \mathrm{m}$ and 2 ms, respectively. (b) Shear wave detected in phantom. The magnitude of the displacement is smaller than the pixel size of $2.9\mu \mathrm{m}$.

Figure 4 demonstrates the results of the ARF/psOCT measurements in the 10 wt. % gelatin phantom. Solid blue-colored lines depict the shear waves induced by the ultrasound transducer [see Fig. 2(c)] positioned closer to the point of observation (i.e., OCT beam). Dashed green-colored lines correspond to the shear waves created by the ultrasound transducer located further away from OCT beam. The normalized shear displacements demonstrate the repeatability of the results. Ultrasound excitation is depicted by shaded area between two vertical lines.

Fig. 4

Shear waves induced in 10 wt. % gelatin phantom by ultrasound transducers closer (solid blue-colored lines) and further (dashed green-colored lines) from the OCT observation spot.

The times of shear wave arrival created by the different transducers were $6.24\pm 0.02\mathrm{ms}$ and $6.82\pm 0.03\mathrm{ms}$, respectively ($\text{mean value}\pm \text{standard deviation}$) resulting in $0.58\pm 0.04\mathrm{ms}$ time delay. The speed of the shear wave was calculated to be $2.8\pm 0.2\mathrm{m}/\mathrm{s}$. Assuming the density of all phantoms to be $1.1{\mathrm{g}/\mathrm{cm}}^3$, the shear modulus of the 10% gelatin phantom was estimated to be $8.6\pm 1.0\mathrm{kPa}$ using Eq. (1). In the phantoms with gelatin concentration of 12.5%, 15%, and 25%, the measured speeds of shear waves were calculated to be $4.0\pm 0.2\mathrm{m}/\mathrm{s}$, $4.7\pm 0.5\mathrm{m}/\mathrm{s}$, and $6.1\pm 0.7\mathrm{m}/\mathrm{s}$, respectively. Corresponding shear moduli of these phantoms were calculated to be $17.6\pm 1.4\mathrm{kPa}$, $24.3\pm 4.0\mathrm{kPa}$, and $40.9\pm 7.3\mathrm{kPa}$, respectively. These results demonstrate a monotonic increase of shear moduli with gelatin concentration. Our measurements qualitatively agree with the literature data,³⁷ while the differences can be possibly attributed to addition and greater concentration of optical and/or ultrasound scatterers.

The developed multitransducer ARF/psOCT probe was tested in ex-vivo studies. Representative shear waves for each ultrasound transducer are shown in Fig. 5. The duration of the ARF excitation is shown using shaded area between two vertical black lines. The solid blue line and the dashed green line in Fig. 5 correspond to the shear waves generated by the ultrasound transducer that is located closer to and further from the optical beam, respectively.

Fig. 5

Representative shear waves produced in rabbit heart muscle by different ultrasound transducers of the ARF/psOCT probe located closer to its tip (solid blue line) and further from the tip (dashed green line).

Times of arrival of the shear waves created by ultrasound transducers located closer to and further from point of observation were measured to be $3.45\pm 0.05\mathrm{ms}$ and $3.76\pm 0.08\mathrm{ms}$, respectively. Given the 1.65-mm distance between the transducers, the shear wave speed was estimated to be $5.3\pm 1.5\mathrm{m}/\mathrm{s}$. Assuming the density of cardiac muscle to be $1.06{\mathrm{g}/\mathrm{cm}}^3$, the cardiac muscle shear modulus was calculated to be $29.8\pm 12.6\mathrm{kPa}$ using Eq. (1).³⁸ Our results are in good agreement with the measurement in the murine heart.²³ At the same time, lower values of shear elastic modulus of human heart tissues are also reported.³⁹ This discrepancy could be explained by differences between rabbit and human tissues and the fact that we have used a heart from an atherosclerotic rabbit genetically prone to cardiac disease.

The calculations of shear moduli using Eq. (1) assume that shear waves propagate in an infinite medium. Obviously, this assumption is too strong for both phantoms and heart muscle thus speeds of shear waves and, therefore, shear moduli are slightly underestimated.⁴⁰ In the future, the ultrasound transducer can also be used to measure directly the thickness of the heart muscle, which would allow for correction of the measured speeds of shear waves and, therefore, shear moduli.

Comparison of experimental results shown in Figs. 3(b) and 5 demonstrates that the magnitude of shear waves in the stiffer heart is greater than that in the softer phantom. One possible explanation is that the interaction between ultrasound waves and gelatin phantom with silica particles was less efficient compared to that with cardiac muscle. Also, the quality of ultrasound transducers in the ARF/psOCT probe was higher compared to the transducer used in the initial tests. In any case, reconstruction of shear modulus relies on temporal characteristics of shear waves rather than on their magnitude, and the only requirement is that shear displacements must be detectable in tissues by psOCT.

Combination of an ultrasound transducer and a miniature OCT probe has been shown previously for two-modality intravascular imaging.⁴¹ However, the ARF/psOCT probe reported here has been designed quite differently. First of all, in intravascular ultrasound imaging catheters, both the center frequency and bandwidth of an ultrasound transducer need to be high for better spatial resolution. On the contrary, ARF excitation is more efficient at lower frequencies and narrow bandwidths, which require thicker PZT crystals. The thicker crystal increases the size of the probe, and, therefore, there is a trade-off between the thicknesses of the crystal and the size of the ARF/psOCT probe. Second, the multitransducer ARF/psOCT probe was designed for quantitative measurements of shear elasticity within a wide range of shear moduli. Indeed, the experimental results presented in Figs. 4 and 5 demonstrate that the shear modulus of both soft ($<10\mathrm{kP}$) and hard ($>30\mathrm{kPa}$) tissues can be measured in $<3\mathrm{ms}$. The multitransducer ARF/psOCT probe enables measurements to be performed very quickly so that tissues are effectively stationary. Third, intravascular OCT imaging catheters are designed such that both imaging depth and spatial resolution are maximized. In contrast, the OCT detector in ARF/psOCT probe must be optimized for detection of very small motions while the lateral resolution is less important.

Unlike air pulse or laser-based excitation of a shear waves, ARF excitation requires an acoustic coupling between the ultrasound transducer and soft tissues. Such coupling can be created by a drop of physiological 0.9% saline that is delivered to the transducers inside a protection sheath, similar to catheters for intravascular imaging. Note that the water filled 0.2- to 0.3-mm gap between the ultrasound transducer and heart muscle will not affect the measurements because the typical shear wave displacements are much smaller than the size of the gap.

The current prototype of the multitransducer ARF/psOCT probe needs further miniaturization. First, the size of ultrasound transducers should be further decreased with simultaneous improvement of ARF generation efficiency. This could be achieved using lower frequency PZT crystals and different backing materials. Second, instead of commercial OCT catheters, a custom OCT detector needs to be designed. Third, the rigid part of the current prototype needs to be redesigned such that the tip of the probe can be remotely manipulated using, for instance, a robotic guidewire to operate the probe remotely.⁴²

While the two-transducer ARF/psOCT prototype was successfully tested ex-vivo, in-vivo evaluation using an animal model is still required. Future work also includes development and manufacturing a cost-effective two-channel power amplifier and its synchronization with electrocardiography equipment to confirm that shear modulus of heart muscle can be reliably measured in-vivo.

In a clinical environment, the proposed multitransducer ARF/psOCT probe could be potentially adapted in endoscopic or invasive procedures, such as intracardiac ultrasound imaging. Shear elasticity of human heart tissues is comparable with that of rabbit hearts or less.^39,43 Given lower human cardiac and respiratory rates,⁴⁴ the shear elasticity of heart tissue and atherosclerotic vessel tissue⁴⁵ could be assessed using an endoscope-like or catheter-like ARF/psOCT probe.

4.

Conclusion

A design of a multitransducer ARF/psOCT probe to assess biomechanical properties has been introduced and successfully tested both in phantom studies and ex-vivo. Shear waves were created, and propagation speeds were measured both in phantoms and in rabbit heart. It was found that the ARF/psOCT probe and method are suitable for assessment of tissues with shear moduli varying in a wide range that covers almost all soft tissues. Shear wave propagating with speed ranging from 2.8 to $6.2\mathrm{m}/\mathrm{s}$, which correspond to shear moduli in the range of 8.6 to 41 kPa, can be reliably measured. The ARF/psOCT probe was used to assess elasticity of a rabbit heart muscle. Our measurements of shear wave speed and shear modulus for the rabbit heart muscle agree well with the experimental results reported by other groups for murine and human heart muscle tissue. Overall, the combined multitransducer ARF/psOCT probe can be used for assessment of microelasticity properties of cardiac tissues.