CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims the benefit of international application No. PCT/IB2014/061576 filed May 21 ^st 2014, the entire contents of which are incorporated herein by reference.

TECHNICAL FIELD OF THE INVENTION

The present invention relates to an ablation method and device or catheter, and in particular to a device and method for performing lowly invasive ablation of targeted tissues (e.g atherosclerotic plaques, renal nerves, cancer tumors) at high ablation energy peak power.

In particular, the ablation of the present invention concerns laser-induced optical breakdown (LIOB) ablation. Ablation being the removal or destruction of a body part or biological tissue using the energy of a laser pulse. The present invention in particular concerns a subsurface laser- induced optical breakdown (LIOB) catheter or device, a method for producing such a catheter or device as well as a method for carrying out subsurface laser-induced optical breakdown (LIOB).

BACKGROUND OF THE INVENTION

LIOB ablation has been used in a free space optical configuration (see "Sub-Surface, Micrometer-Scale Incisions Produced in Rodent Cortex using Tightly-Focused Femtosecond Laser Pulses", J. Nguyen et al., 2012, Lasers in Surgery and Medicine). However, such a configuration does not permit LIOB to be carried out in-vivo inside a body, for example, in arteries.

As opposed to the catheters for removing plaque currently known (see, for example, Patent US 6673064), the present invention concerns a device or catheter that is specifically designed to handle high energy (>0.1 μΓ), ultrashort (less than 1000 picosecond, preferably less than 100 picosecond and most preferably <1 picosecond) laser pulses. To carry out LIOB in-vivo, this high energy and ultrashort pulse must be delivered and focused outside the catheter away from an initial optical pulse delivery axis. Conventional optics that are used to redirect and focus such a pulse produce temporal delays that change the ultrashort pulse duration resulting in a broadened pulse (or change in pulse shape) and degraded ablation or no ablation at a target site.

Moreover, conventional optics used in a low profile catheter would produce low numerical aperture (NA) focusing which could produce damage at the tissue surface. This is particularly problematic in regions close to vital organs such as the heart or the brain where fragments of damaged tissue can subsequently enter the blood stream if inner tissue walls are damaged or removed during the ablation process. SUMMARY OF THE INVENTION

The present invention solves the above mentioned problems by providing a subsurface laser-induced optical breakdown LIOB catheter including a probe assembly for providing at least one optical pulse to a subsurface target biological tissue site and an optical waveguide for guiding the at least one optical pulse to the probe assembly. The probe assembly includes at least one optical element configured (i) to redirect the at least one optical pulse, upon reflection at the at least one optical element, from a waveguide propagation direction to the subsurface target biological tissue site, and (ii) to shape the at least one optical pulse upon reflection at the at least one optical element to maintain a pulse duration of the at least one optical pulse at the subsurface target biological tissue site.

The catheter of the present invention advantageously includes an optical element to redirect and focus a high energy and ultra-fast optical pulse produce that avoids introducing temporal delays that change the ultrashort pulse duration resulting in a broadened pulse. This thus prevents a degraded ablation and assure successful ablation at a target site. Furthermore, this optical element advantageously permits moderately high NA focusing to be obtained. The invention advantageously provides focusing with a moderately high numerical aperture (for example NA > 0.5 or more preferably NA > 0.8) permitting successful ablation to be achieved.

The present invention also relates to a method for carrying out subsurface laser-induced optical breakdown LIOB according to claim 30.

BRIEF DESCRIPTION OF THE FIGURES:

The above object, features and other advantages of the present invention will be best understood from the following detailed description in conjunction with the accompanying drawings, in which:

Figure 1 illustrate a general view of a catheter according to the present invention;

Figure 2 presents the use of a catheter balloon for affixing the probe assembly within an arterial wall;

Figure 3 presents a generalized configuration for the probe assembly of the catheter;

Figure 4 illustrates an assembly in which a central shaft is located in the middle of the probe assembly of the catheter;

Figure 5 illustrates an asymmetric positioning of the central shaft within the probe assembly of the catheter;

Figure 6 illustrates an off-axis paraboloidal mirror for focusing a collimated beam;

Figure 7 shows an example of parabola section;

Figure 8 shows two paraboloid mirrors for focusing a beam emitted from a fiber;

Figure 9 illustrates an off-axis elliptical mirror for focusing a diverging beam emitted from an optical fiber;

Figure 10 shows an example of an elliptical section;

Figure 11 shows an off-axis elliptical mirror for focusing a diverging beam similar to embodiment 5.3;

Figure 12 presents an off-axis ellipsoidal mirror for focusing a diverging beam similar to embodiment 5.3;

Figure 13 presents a refractive index modulator used to vary the index of refraction of a segment located after the focusing optics for modulating the focused light depth; Figure 14 presents the integration of an imaging system with focusing optics;

Figure 15 shows an example of a bifocal parabolic sections;

Figure 16 shows the embodiment of Figure 12 where a bifocal mirror is configured to rotate relative to the collimated beams;

Figure 17 illustrates a further embodiment with the integration of an imaging system with focusing optics; and

Figure 18 shows the embodiment of Figure 14 where a bifocal mirror is configured to rotate relative to the diverging beams. DETAILED DESCRIPTION OF THE INVENTION

The present invention relates to an ablation device for performing lowly invasive ablation of targeted tissues, via a catheter. A high peak power, ultrashort laser pulse delivered through a catheter ablates tissue under the surface tissue through a multi-photon ionization process called laser induced optical breakdown. Several possible subsystem embodiments of the invention will be presented.

This probe is specifically designed to handle high energy (>0.1 uJ), ultrashort (<1 picosecond) laser pulses. The function of the probe is to focus a pulse delivered by a waveguide (fiber optic or photonic crystal fiber) and focus it away from the initial optical axis and outside of the probe device. The invention advantageously provides focusing with a moderately high numerical aperture (for example NA > 0.5 or more preferably NA > 0.8 ) permitting successful ablation to be achieved. This is difficult to achieve using conventional optics in a low profile catheter. In this invention we describe the use of curved mirrors to achieve moderately high NA focusing.

Figure 1 presents the general idea of the catheter probe. A guide wire is used for positioning the probe assembly. A guide wire shaft (1 in Figure 1) surrounds the guide wire for placement of the ablation catheter after the guide is placed in the artery near the target. A stabilization system (2 in Figure 1) affixes the position of the catheter relative the arterial wall (3 in Figure 1). An ultrashort optical pulse is delivered to the probe assembly via a fiber optic waveguide (4 in Figure 1). The probe assembly (5 in Figure 1) then redirects and shapes the pulse to focus it out the side of the assembly (6 in Figure 1). This directs the focus under a tissue layer at the desired target location (7 in Figure 1). Within the probe assembly 5 are also the components necessary to enable optical or acoustic imaging of the target area. The visualization information is relayed out of the probe assembly through the ablation pulse delivery waveguide, a second optical or acoustic waveguide, or an electrical signal along a wire (4 in Figure 1). The present invention as described herein concerns different embodiments for the invention subsystems or features: guide wire, stabilization subsystem, pulse delivery mechanism, probe assembly, focusing optics, and visualization subsystem.

1. Guide wire

A catheter guide wire is used for exact placement of the catheter probe. The guide wire is guided by a physician through the arteries to the target area. Once in place the catheter follows the guide wire for placement in the target artery. 2. Stabilization system

With the catheter probe in place it may be preferable to stabilize it in relation to the arterial walls for ablation. Subsystem embodiment 2.1

Figure 2 presents the use of a catheter balloon 9 for affixing the probe assembly within the arterial wall 3. The balloon central position is determined by a central shaft 13. When inflated the balloon presses against the arterial walls 3, thus creating a fixed position in the artery. In this subsystem the probe assembly 5 and waveguides 4 are free to rotate and translate along the central shaft 13 within the balloon 9, thus allowing ablation along the arterial wall or at a biological tissue site at any point within the balloon 9. The balloon material is at least partially transparent at the wavelength of the electromagnetic energy used to carry out ablation. The rotation and translation of the probe assembly 5 can be accomplished manually or through a motor.

3. Pulse delivery mechanism

The delivery of high energy, ultrashort pulses to the catheter probe 5 for tissue ablation necessitates the use of special optical fibers. Conventional optical fibers have silica core and cladding with small mode diameters. High energy, ultrashort pulses would potentially induce laser induced optical breakdown within a conventional optical fiber because of the high intensity, thus destroying it.

A preferred pulse delivery mechanism is hollow core photonic crystal fibers, which do not guide light in a silica core, but in air. Thus, decreasing the likelihood of laser induced optical breakdown within the fiber. Furthermore, the hollow core fibers can guide light with mode diameters much larger than conventional fibers, thus decreasing the intensity in the fiber and further reducing the probability of laser induced optical breakdown.

Small mode diameter, hollow core photonic crystal fibers often have the characteristic of a higher numerical aperture compared to larger mode diameter fibers. Thus, light emitted from them can be more divergent. Whereas, the large mode diameter, hollow core photonic crystal fibers have a lower numerical aperture. Thus, light emitted from these have low divergence. As a result these fiber types require different optical systems for focusing the ablation beam at an appropriately high numerical aperture.

4. Probe assembly

The probe assembly consists of the focusing and imaging optics used to focus the beam out of the side of the assembly and to collect light or acoustic waves for imaging, as well as the housing which encapsulates these components. This assembly must be sealed to prevent liquid from entering the assembly. Figure 3 presents a generalized configuration for the probe assembly 5. An optical fiber 15 delivers pulsed light inside of the probe assembly 5. The focusing optics 17 reflect and focus the light to a point outside of the assembly 5 through the window 19 on the side of the probe assembly 5. The focusing optics may also collect light which passes from the tissue and through the window 19 and deliver it to an optical fiber for image creation. Additionally, any components for acoustic or optical observation of the tissue or ablation process are integrated into this subsystem. A key component of the probe assembly 5 is the location of the central shaft 13 which determines the location of the probe assembly 5 in relation to the balloon 9. Following are two subsystem embodiments for central shaft location within the probe assembly.

Subsystem embodiment 4.1

Figure 4 illustrates an assembly in which the central shaft 13 is located in the middle of the probe assembly 5. This configuration is ideal for scenarios in which the probe assembly 5 is similar in size to the artery. Thus, window 19 of the probe assembly 5 will be placed alongside the artery endothelium allowing for a predictable optical focus 21 depth.

Subsystem embodiment 4.2

Figure 5 illustrates the use of an asymmetric positioning of the central shaft 13 within the probe assembly 5. This shaft 13 is placed in the probe assembly 5 near the side opposite the focusing window 19. The asymmetry allows for the window 19 of the catheter probe 5 to be positioned close to or against an arterial wall (as shown in figure 1) to ensure positioning of the optical focus 21 under the endothelium. This would be most useful in wider arteries. Furthermore, the central shaft 13 acts as the axis of rotation for the probe assembly 5. Thus, the probe assembly window 19 will always be near the endothelium as the probe assembly 5 rotates within the balloon 9. 5. Focusing optics

The focusing optics subsystem 17 comprises the optical components which reflect and focus the pulsed light emitted from the optical fiber 15.

Subsystem embodiment 5.1:

Figure 6 presents the use of an off-axis paraboloidal mirror 23 (for example, an off-optical axis segment taken from one side of a full paraboloid mirror) for focusing a collimated beam. In this representation a lens 25 collimates the diverging light emitted from an optical fiber 15. The paraboloidal mirror 23 includes a reflecting surface or layer(s) that takes the shape of an inner section of a paraboloid mirror, not including the paraboloid vertex. This off-axis paraboloid focuses light away from the optical axis. Any light arriving on the mirror surface parallel to the paraboloid axis of symmetry will converge to the paraboloid focus, which is designed to be outside of the probe assembly 5 and, for example, at an ablation site in the biological tissue to be ablated. The paraboloid mirror 23 has the beneficial characteristic that the path length of all photons arriving at the mirror 23, traveling the path parallel to the paraboloid axis of symmetry, will reflect and converge at the paraboloid focus after traveling the same path length. This advantageously ensures that temporal delays typically induced in glass lenses are not introduced with the reflecting mirror surface and the ultrashort pulse duration remains unchanged.

For example:

Figure 7 shows a parabola section defined using the parabolic equation:

1 2

y =— ^χ - p,

4p

where p = 400 μπι. The parabola is defined from x = 300 μιη to 1300 μηι. To create a paraboloid surface the section is rotated about a line which connects the parabola vertex and focus (the axis of symmetry of the parabola). The mirror is limited to the region on which the laser reflects off. Light delivered by an optical fiber is collimated by a lens. The collimated light travels in the x-direction, parallel to the line which connects the defined parabola vertex and focus and towards the concave surface of the paraboloid. After reflection off the mirror, the light converges and focuses at the paraboloid focus, which is 300 μηι below the defined mirror. The focusing NA in this scenario is 0.61.

Subsystem embodiment 5.2:

Figure 8 presents the use of two paraboloid mirrors for focusing the beam emitted from the fiber 15. The first paraboloid mirror Ml (for example, an on-optical axis segment taken from one side of a full paraboloid mirror) includes a reflecting surface or layer(s) that takes the shape of an inner paraboloid section including the paraboloid mirror vertex and with the paraboloid vertex aligned along the central axis of the optical fiber 15. The fiber distal end is placed at the focus of the Ml paraboloid, which will be near or placed at the reflecting surface and center of the second paraboloid mirror M2. Placement in the center of M2 requires a through bore or hole in the mirror M2 for the optical fiber 15 positioning. Any light emitted from the fiber (marked as Al on Figure 8) which reflects off mirror Ml (marked as A2 on Figure 8) will propagate parallel to the axis of symmetry of the Ml paraboloid mirror. This light, in turn, propagates parallel to the axis of symmetry of off- axis paraboloid mirror M2. Thus, upon reflecting off mirror M2 (marked as A3 on Figure 8), this light will converge to the mirror M2 paraboloid focus, thus creating a high intensity optical focus outside of the probe assembly 5 and, for example, at an ablation site in the biological tissue to be ablated. The second paraboloid mirror M2 includes, for example, an off-optical axis segment taken from one side of a full paraboloid mirror. The second paraboloid mirror M2 takes the shape of an inner paraboloid section not including the paraboloid mirror vertex.

The paraboloid mirrors Ml , M2 have the beneficial characteristic that the path length of all photons arriving at the mirrors, will reflect and converge at the paraboloid focus after traveling the same path length. This advantageously ensures that temporal delays typically induced in glass lenses are not introduced with the reflecting mirror surfaces and the ultrashort pulse duration remains unchanged.

Subsystem embodiment 5.3:

Figure 9 presents the use of an off-axis elliptical mirror 26 (for example, off-optical axis segment taken from one side of a full ellipsoidal mirror) for focusing a diverging beam emitted from an optical fiber 15. The ellipsoidal mirror 26 includes a reflecting surface or layer(s) that takes the shape of an inner ellipsoid section. Any light arriving on the mirror surface originating from a focus (Focus 1 at which the fiber emitting end is positioned) of the ellipsoid surface will converge to the second focus (Focus 2), thus creating a high intensity optical focus.

The ellipsoidal mirror 26 has the beneficial characteristic that the path length of all photons arriving at the mirror 26, traveling the path from Focus 1 , will reflect and converge at Focus 2 after traveling the same optical path length.

This advantageously ensures that no temporal delays are introduced and the pulse duration remains unchanged. This embodiment is particularly useful for optical fibers which have a (sufficiently) divergent beam.

For example:

An elliptical mirror is designed using the ellipse equations:

x = a cos Θ + b sin Θ

y = a sin Θ + b sin Θ

where = 1200 μηι and c = 1030 μηι. Θ ranges from 4.79 rad to 5.94 rad to provide a segment of the ellipse (shown in Figure 10). The ellipsoidal surface is defined by rotating the ellipse around the ellipse major axis. The mirror would be limited to the area over which the beam reflects off of the mirror. A hollow core, photonic crystal fiber designed for single mode propagation of 1030 nm light, with mode diameter of 7.5 μηι, has a beam half angle divergence of 5 degrees. The fiber core is placed at a focus of the ellipse and tilted to 17 degrees relative to the major axis of the ellipsoid, such that the emerging light is directed toward the mirror.

The reflected light is then refocused to the second ellipsoid focus at a position orthogonal to the original beam direction and 250 μηι below the mirror. The focusing NA in this scenario is 0.61. Subsystem embodiment 5.4:

Figure 11 presents the use of an off-axis elliptical mirror 27 (off-optical axis segment taken from one side of a full ellipsoidal mirror) for focusing a diverging beam just as in subsystem embodiment 5.3. However, this embodiment is useful for an optical fiber 15 which does not have a sufficiently divergent beam (that is a weakly divergent beam at the fiber exit). In this embodiment a diverging lens 29 (a lens with negative focal length) is used to create a more strongly diverging beam. The ellipsoidal mirror 27 is placed such that the virtual focus from the diverging lens is placed at Focus 1 of the ellipsoidal mirror surface to converge all light at Focus 2.

The ellipsoidal mirror 27 has the beneficial characteristic that the path length of all photons arriving at the mirror, traveling the path from Focus 1 , will reflect and converge at Focus 2 after traveling the same optical path length. This advantageously ensures that no temporal delays are introduced and the pulse duration remains unchanged.

Subsystem embodiment 5.5:

Figure 12 presents the use of an off-axis ellipsoidal mirror M2 (for example, an off-optical axis segment taken from one side of a full ellipsoidal mirror) for focusing a diverging beam just as in subsystem embodiment 5.3. However, this embodiment is useful for optical fibers which do not provide a sufficiently divergent beam or weakly divergent beam exiting the fiber end (marked as light emitted from the fiber 15, labeled as Bl in figure 12).

The fiber 15 is placed in the center of mirror M2 requiring a hole in the mirror M2 for the optical fiber positioning. In this embodiment a convex mirror Ml including a reflecting surface or layer(s) shaped as the outside surface of a paraboloid, or approximated with a spherical surface, is placed to create a more strongly diverging beam (labeled as B2). This vertex and focus of this paraboloid are aligned along the optical axis of the emitted light of the fiber 15. The divergent light in turn reflects off of the ellipsoidal mirror M2, which is placed such that the virtual focus from the mirror Ml is placed at Focus 1 of the ellipsoid surface to converge all light at Focus 2 (labeled as B3).

The ellipsoidal mirror M2 has the beneficial characteristic that the path length of all photons arriving at the mirror, traveling the path from Focus 1, will reflect and converge at Focus 2 after traveling the same optical path length. This advantageously ensures that no temporal delays are introduced and the pulse duration remains unchanged.

Subsystem embodiment 5.6 Figure 13 presents the use of a refractive index modulator 33 to vary the index of refraction of a segment after the focusing optics 17 as the light focuses. This index modification displaces the focal spot location allowing for axial control over the focus location (control of a focus location at a given depth in the biological tissue). This can be realized via a liquid crystal or electrowetting device as the refractive index modulator or refractive index changing device 33. The index modulator 33 can be implemented with any one of the above mentioned subsystems 5.1 to 5.5.

6. Imaging system The present invention further includes an imaging system to characterize the arterial plaque for identification of ablation target, as well as for monitoring the ablation process. These include optical and acoustic imaging modalities. The imaging light and image fed back, for example through a waveguide, is visualized using the associated image reconstruction technique, for example optical coherence tomography (OCT). 6.1 Optical Coherence Tomography

OCT is an established imaging modality for intravenous imaging. Combining OCT with sub-surface tissue ablation allows for identification and characterization of plaque. The bubbles created through the ablation process can be imaged via OCT.

To enable OCT in the catheter device, light for imaging must be delivered to the probe 5. This can be done through the ablation fiber 15, or through a separate fiber included in the catheter. This light must then be focused into the tissue. Preferably this light should be focused at a lower numerical aperture than the ablation light. This will allow for a deeper imaging depth. The backscattered light which contains the necessary OCT image information can be collected by either the OCT optics or the ablation focusing optics and returned for analysis through either fiber 15 and/or the additional fiber.

The OCT system can utilize near infrared light from a high bandwidth source to enable deep penetration and high resolution axially [1]. The design numerical aperture of the OCT mirrors would preferable be less than 0.5.

Subsystem embodiment 6.1.1 Figure 14 presents the integration of an OCT imaging system with the focusing optics. The OCT light is delivered to the probe assembly 5 via a single mode fiber optic 35 (SMF). This light emerges diverging from the fiber facet and is collimated by a lens 37 (2 in Figure 14). This light is then reflected off of a bifocal, paraboloid mirror 39. For example, the elements for focusing the inputted ablation pulse are the same as those described previously in the embodiment of Figure 6 but the off-axis paraboloid focusing mirror additionally includes a reflector section to direct and focus the OCT light delivered to the probe assembly 5. The backscattered light from the zone to be imaged is collected by mirror 39 and is propagated back for imaging and visualization via either or both of fibers 15, 35.

The ablation beam comes from optical fiber 15 placed next to the OCT SMF 35. This ablation beam is collimated by a lens 41 and also reflects off of the bifocal, paraboloid mirror 39. The collimated ablation and collimated OCT beams do not spatially overlap on the mirror 39. The bifocal focusing mirror 39 is designed (comprises a specific curvature) to focus the OCT light, which reflects off of a separate OCT portion of the bifocal reflective mirror 39 with a different curvature to the curvature for forming a focused ablation spot, and provides a lower numerical aperture and deeper focal spot than the ablation beam. Thus, allowing for a large OCT imaging depth. The OCT optical path could overlap the ablation focal spot for monitoring of the ablation process.

For example:

Figure 15 shows two parabola sections each defined using the parabolic equation: y = ^-(x- x ₀ ) x ² - p,

4p

where for the lower portion (for the ablation beam) xo = 0 and p = 400 μηι across the range from x = 300 μηι to 1300 μηι. The upper portion ( for the OCT beam) is defined with xo = -200 μηι and p = 312.25 μηι across the range x = 1300 μηι to 1800 μπι. To create a paraboloid surface each section is rotated about a line which connects the respective parabola vertex and focus (the axis of symmetry of the parabola). Each mirror section is limited to the region on which the respective laser reflects off of it.

Two beams of light are delivered via separate optical fibers and each is collimated by separate lenses and travel in the same direction. The collimated light from each beam travels in the x-direction, parallel to the line which connects the defined parabola vertex and focus and towards the concave surface of the bifocal paraboloid. After reflection off the mirror, each beam converges and focuses at the respective paraboloid focus. For the ablation component, this is 300 μηι below the defined mirror with a focusing NA of 0.61. The OCT section focuses 500 μηι below the defined mirror with a focusing NA of 0.15. Subsystem embodiment 6.1.2 Figure 16 presents a variation on previous subsystem embodiment 6.1.1. In this case the subsystem further includes rotation means or a rotator. The bifocal mirror 39 rotates independent of or relative to the collimated beams, either through the rotating means carrying out rotation of the mirror or rotating the fibers 15, 35 and lenses 37, 41. In this embodiment bifocal mirror 39 is configured or designed such that the OCT portion of the mirror 39 always focuses to the same focal spot location (relative to the mirror 39) during rotation despite a shift of the direction of beam incidence. The focused spot depth of the ablation light pulse is also not displaced by the relative rotation of the bifocal mirror 39 to the fibers 15, 35 and the lenses 37, 41. Subsystem embodiment 6.1.3

Another variation of subsystem embodiment 6.1.1 or 6.1.2 also utilizes the bifocal mirror 39 rotating independent of the collimated beams, either through rotation of the mirror or rotating fibers 15, 35 and lenses 37, 41. However, the bifocal mirror in this embodiment is designed such that the OCT portion of the mirror focuses to different focal spot locations relative to the mirror 43 under rotation to provide additional information through the rotation.

Subsystem embodiment 6.1.4

Figure 17 presents the integration of an OCT imaging system with the focusing optics. The OCT light is delivered to the probe assembly 5 via a single mode fiber optic 35 (SMF). This light emerges diverging from the fiber facet and could be modified to be either less divergent or collimated by a lens 37. This light is then reflected off of a bifocal mirror 43 including a reflecting section designed or configured to focus the OCT beam. The ablation beam comes from a fiber 15 placed next to the OCT SMF 35. This ablation beam diverges from a virtual focus placed at a focus of the ellipsoidal surface of the bifocal mirror 43 after passing through a diverging lens 45. This beam reflects off an ellipsoidal shaped part of the bifocal mirror 43. For example, the elements for focusing the inputted ablation pulse are the same as those described previously in the embodiment of Figure 11 where the off- axis paraboloid mirror additionally includes the reflector section to direct OCT light delivered to the probe assembly 5. The backscattered light from the zone to be imaged is collected by mirror 43 and is propagated back for imaging and visualization via either or both of fibers 15, 35. The ablation and OCT collimated beams do not spatially overlap on the mirror 43. The bifocal focusing mirror 43 includes a reflector section configured or designed to focus the OCT light, which light interacts with a separate portion of the bifocal reflective mirror 43 with a lower numerical aperture and deeper focal spot than the ablation beam. Thus, allowing for a large OCT imaging depth. The OCT optical path could alternatively spatially overlap the ablation focal spot for monitoring of the ablation process.

Subsystem embodiment 6.1.5

Figure 18 presents a variation on previous subsystem embodiment 6.1.4. In this case the subsystem further includes rotating means or a rotator. The bifocal mirror 43 rotates independent of the collimated beams, either through a rotating of the mirror 43 or rotating the fibers 15, 35 and lenses 37, 45. The rotation occurs around the center of the optical axis of the ablation beam. The bifocal mirror 43 in this situation is configured or designed such that the OCT portion of the mirror always focuses to the same focal spot location (relative to the mirror 43) during rotation despite a shift of the direction of bean incidence. The focused spot of the ablation light is also not displaced by the rotation of the mirror 43 relative to the fibers 15, 35 and lenses 37, 45. Subsystem embodiment 6.1.6

A variation of subsystem embodiment 6.1.5 also utilizes the bifocal mirror 43 rotating independent of the collimated beams, either through a rotating of mirror 43 or rotating fibers 15, 35 and lenses 37, 45. However, the bifocal mirror in this situation includes a reflective section designed or configured such that the OCT portion of the mirror focuses to different focal spot locations (relative to the mirror 43) under rotation to increase the spatial information through the rotation. 6.2 Two Photon Fluorescence Imaging

Two photon fluorescence imaging utilizes two-photon excitation of fluorophores (TPF) for imaging. This technique requires that high intensity, ultrashort laser pulses be tightly focused. The high-energy laser pulses are delivered through a fiber optic. These will be focused and scanned on the sample [2].

The optical configurations combining ablation and OCT beams could be implemented with TPF imaging by utilizing the previously described OCT optics and components for TPF imaging.

To enhance the TPF imaging the plaque or tissue may be stained with targeted dyes.

TPF imaging would preferably be implemented with a high repetition rate, ultrafast pulsed laser operating at near infrared wavelengths.

6.3 Fluorescence Imaging

Fluorescence imaging can also be used to provide imaging contrast in the artery. Fluorescence imaging requires an excitation laser to interact with the tissue sample of interest. The excitation beam can be delivered to the probe assembly 5 via the ablation light delivery waveguide 15 or through an additional waveguide. The optical excitation beam works with a high numerical aperture focus, or with a low numerical aperture focus. This can also be enhanced by staining the tissue or plaque [3].

The optical configurations combining ablation and OCT beams previously described can be implemented with fluorescence imaging by utilizing the described OCT optics and components for fluorescence imaging.

6.4 Photoacoustic Imaging Photoacoustic imaging is an imaging modality in which light interactions with materials can be measured through acoustic wave detection. Specifically, various biological tissues have unique absorption spectral characteristics. Photoacoustic imaging uses these differences to provide contrast in imaging various tissues. As light propagates through tissue, it is absorbed in tissue causing a temperature increase. The increase in temperature results in an expansion of the tissue, which creates an acoustic wave. An ultrasound transducer allows detection and measurement of the acoustic wave to quantify the absorption in the tissue and, in turn, tissue characterization [4].

An optical fiber provides a laser pulse at a wavelength tuned to plaque absorption for photoacoustic imaging, for example 1200 nm. This light is either tightly focused and scanned through the sample, or is focused with a low NA for large depth of field. An ultrasound transducer must also be placed in the probe assembly 5 for acoustic wave detection. In this case the detected acoustic wave would be converted to an electrical signal and passed out of the probe assembly and the catheter through an electrical wire. The photoacoustic optical focus and the acoustic detector sensitive region must overlap to ensure that the acoustic transducer can measure the acoustic waves.

The photoacoustic signal can also be used to quantify the ablation. During laser induced optical breakdown a cavitation bubble is created along with an acoustic wave, which can be detected with the acoustic transducer. 7. Rotating pullback

The catheter is built with a rotating pullback mechanism or functionality to allow for axial and rotational movement of the catheter [5]. When implemented with the present invention as for example illustrated in subsystem embodiment 2.1 the axial and rotational movement of the probe assembly 5 takes place within and independent of the balloon 9.

8. Flushing In a catheter which does not utilize stabilization, for example as described in subsystem embodiment 2.1 , the blood may be freely flowing around the catheter. The blood cells pose a potential problem to the ablation pulse and to imaging through scattering the light by the blood cells. In this case the blood may be flushed out with a non-scattering liquid to perform imaging and ablation.

Having described now the preferred embodiments of this invention, it will be apparent to one of skill in the art that other embodiments incorporating its concept may be used. This invention should not be limited to the disclosed embodiments, but rather should be limited only by the scope of the appended claims.

In particular, elements of each described embodiment can be combined with the elements included in other embodiments.

References

[1] B. H. Lee, E. J. Min, and Y. H. Kim, "Fiber-based optical coherence tomography for biomedical imaging, sensing, and precision measurements," Opt. Fiber Technol, vol. 19, no. 6, Part B, pp. 729-740, Dec. 2013.

[2] M. T. Myaing, D. J. MacDonald, and X. Li, "Fiber-optic scanning two-photon

fluorescence endoscope," Opt. Lett., vol. 31, no. 8, pp. 1076-1078, Apr. 2006. [3] S. Liang, A. Saidi, J. Jing, G. Liu, J. Li, J. Zhang, C. Sun, J. Narula, and Z. Chen, "Intravascular atherosclerotic imaging with combined fluorescence and optical coherence tomography probe based on a double-clad fiber combiner," J. Biomed. Opt, vol. 17, no. 7, pp. 0705011-0705013, 2012.

[4] K. Jansen, A. F. W. van der Steen, G. Springeling, H. M. M. van Beusekom, J. W.

Oosterhuis, and G. van Soest, "Intravascular photoacoustic imaging of human coronary atherosclerosis," 2011, vol. 7899, pp. 789904-789904-7.

[5] G. J. Ughi, T. Adriaenssens, W. Desmet, and J. D?hooge, "Fully automatic three- dimensional visualization of intravascular optical coherence tomography images: methods and feasibility in vivo," Biomed. Opt. Express, vol. 3, no. 12, pp. 3291-

3303, Dec. 2012.