Laser-Induced Cavitation Bubble

Laser-induced microbubbles – bubbles formed in tissues or materials with a high-focused laser pulse – are playing an increasingly important role in various areas of medicine, biology, tissue engineering etc. In fact, these microbubbles are the main instrument in laser microsurgery [1] and can be used to assess the biomechanical properties of soft tissues at the micro-level [2].

To create a bubble the laser beam is focused inside of a tissue. The wavelength of laser pulse is chosen to avoid linear absorption in tissues since the bubble is formed according to non-linear mechanisms of light absorption [3,4]. The formation and dynamic behavior of microbubbles should be a repeatable process but modest variations of optical properties of tissues and fluctuations of laser beam parameters result in formation of microbubbles that differ in size [1]. In addition, the location of these microbubbles varies. The variations in size and location of the microbubbles should be as small as possible to perform successful surgery, to reduce the volume of damaged tissue, and to meet the stringent requirements of laser safety. Ideally, the laser-induced microbubbles should have micron and sub-micron sizes and are formed typically using femtosecond laser pulses.

Optical methods [5 -7], such as bright-field and phase-contrast time-resolved imaging, stroboscopic photo-picturing, light scattering techniques, etc., cannot often be used to characterize microbubbles at sufficient depth because light scattering in tissues is high. However, the propagation of an ultrasound wave does not depend on light scattering. Moreover, microbubbles are excellent reflectors of ultrasound waves – because of this property, microbubbles are used as contrast agents in ultrasound imaging [8]. Finally, a strong shock wave is generated during the laser-tissue interaction [9]. All of the above suggests that ultrasound can be used to characterize and to monitor laser-induced microbubbles in opaque media.

The first set of experiments was performed using microbubbles formed in gelatin and in water. The study was designed to measure sizes and temporal dynamics of microbubbles using the ultrasound-based approach developed by our group. When a bubble is formed, a shock wave is generated and propagates spherically away from the site of laser-tissue interaction. During this propagation, the shock wave is converted into an acoustic wave. This process is named “passive acoustic emission”. At the same time, an external ultrasound pulse can be sent to tissue to initiate “active ultrasound probing” of the microbubble using pulse-echo ultrasound where a pressure pulse is reflected from the microbubble.

To measure the size and location of the microbubble, the magnitudes of both passive acoustic emission and active ultrasound probing cannot be used since both waves attenuate unpredictably in tissue. Instead, the time of arrival technique is used to estimate the parameters of the bubble. Consider the sources of both active probing signals and passive acoustic emissions. When the laser pulse interacts with tissue, a microbubble does not yet exist [10] – the source of the passive acoustic emission corresponds to the origin of microbubble, i.e., a geometrical center of a forming microbubble. To probe the forming microbubble, the ultrasound pulse is sent to a tissue at a later time – this pulse is reflected from the tissue-bubble boundary. By controlling the time delay between laser pulse and ultrasound pulse, the microbubble can be probed at different stages of its evolution. Therefore, combination of passive and active ultrasound measurements allow for the detection of bubble formation and for the monitoring of bubble dynamics.

Fig.1. Experimental setup and geometry of measurements

Figure 1 demonstrates the experimental setup and further describes the proposed approach. The microbubble is shown schematically, not to scale, at a certain stage of its evolution. The passive acoustic emission propagates the distance (D) directly to the transducer thus defining the location of origin. The external ultrasound pulse, reflected from a surface of the microbubble, has to pass the distance (d) between the transducer and the upper wall of the microbubble. Both distances (D and d) are calculated using the time-of-flight approach. It is clear that the difference between the measured distances (D-d) corresponds to the radius of the microbubble.

In our experimental studies, we used both nanosecond and femtosecond laser pulses. In both cases we estimated the size of microbubbles at different stages of its evolution by adjusting the delay between laser pulse (bubble formation) and ultrasound pulse (evolving bubble). The results are shown in Fig. 2. The left panel corresponds to the bubbles formed by long and high-energy pulses (5 ns and 3.75 mJ) – the diameter of microbubble is more than a millimeter. In the case of short and low-energy laser pulses (100 fs and 200 nJ), the bubbles are smaller (right panel of the Fig. 3). The experimental results compare well with the theoretical model of microbubble dynamics based on the Rayleigh model (dashed lines).

Fig.2. Measured (time-of-flight measurements) and modeled (Rayleigh model of cavity collapse) dynamic behavior of microbubbles produced by the pulsed nanosecond laser (right) and femtosecond laser (left) pulses.

The role of microbubbles in surgery can be extended beyond the surgical tool. For example, an exciting extension of application of the laser-induced microbubbles in surgery is the measurement of elastic properties of tissues at a microscopic level. Indeed, knowledge of the elastic properties of tissue is very important in remote palpations [11], corneal and cataract laser surgery [12], elasticity imaging, etc.[13].

Fig. 3. Displacement of the sphere in gelatin with the same shear modulus initiated by different applied forces. Radiation force was applied during the time indicated by two vertical lines.

To test this approach, experiments using a rigid sphere were conducted. The rigid sphere was chosen to accurately model the translational displacement of a microbubble. Generally, a force acting on the sphere results in sphere displacement proportional to the Young modulus of the surrounding tissue. However, since ultrasound waves are attenuated in soft tissues, the precise amplitude of the acoustic force is not known and, therefore, the amplitude of the displacement may not accurately represent the elastic properties of the surrounding medium. This effect is demonstrated in Fig. 3 where the rigid sphere was displaced using forces produced by acoustic pulses of various amplitudes. Clearly, the displacement magnitude increases as the force acting on the sphere rises.

To assess the shear modulus and to avoid the problem with attenuation of applied force, we have developed an approach to assess tissue elasticity based on temporal measurements of the displacements. Indeed, the analysis of displacements, presented in Fig. 3, indicates that modest changes in magnitude of the acoustic force scales linearly the displacement magnitude but does not affect the temporal behavior of the displacement. However, once the pulsed acoustic radiation force is applied, the time required for the sphere to reach the maximum displacement is the same regardless of the amplitude of the applied force. Therefore, the result in Fig. 3 suggests that temporal behavior of the microbubble’s displacement in response to pulsed acoustic radiation force can be used to assess tissue elasticity.

Fig. 4. Displacement of the rigid sphere in gelatin with different shear moduli. The duration and the amplitude of the applied acoustic radiation force was the same. Radiation force was applied during the time indicated by two vertical lines.

The relationship between tissue elasticity and temporal behavior of a modeled microbubble (i.e., rigid sphere) is demonstrated in Fig. 4. Clearly, the amplitude of the sphere’s displacement increases in softer medium – this is expected since the same radiation force was applied. However, it is also noticeable here that the time when the sphere reaches its maximum displacement also changes – it takes longer for the sphere to be displaced in a softer medium.

Figure 5 demonstrates the relationship between the time when the sphere reaches its maximum displacement and the shear modulus of the surrounding tissue. Clearly, the time reduces for stiffer materials – the dependence between temporal behavior of the sphere and material property is consistent with experimental observations and theoretical analysis [14].

Fig. 5. The relationship between the time when the sphere reaches the maximum displacement and the shear modulus of the surrounding material. Points represent experimental measurements and the solid line outlines the theoretically derived dependence.

Another important parameter which should be measured is the viscosity of tissue. We have already developed a theoretical foundation and have conducted preliminary experiments demonstrating that the bubble-based approach can be successfully employed to measure shear viscosity.

Overall, we are developing a technology to assess viscoelastic properties of tissue based on ultrasound and optical measurements of laser-induced gas bubble dynamics. This technology can be used to assist several biomedical applications where laser-induced gas bubbles are used for diagnosis, therapy, and monitoring. The applications of the developed technology may include cellular and molecular imaging, targeted drug delivery, gene therapy, nanosurgery, and many other fields in clinical practice and biomedical research.

REFERENCES

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14. S. R. Aglyamov, A. B. Karpiouk, Y. A. Ilinski, E. A. Zabolotskaya and E. S. Y., "Acoustic radiation force initiated motion of a solid sphere embedded in a viscoelastic medium: theoretical analysis and experimental verification," Journal of the Acoustical Society of America 122, 1927-1936 (2007).

Laser-Induced Cavitation Bubble

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REFERENCES