Double-clad fiber (DCF) is herein used in conjunction with a double-clad fiber coupler (DCFC) to enable simultaneous and co-registered optical coherence tomography (OCT) and laser tissue coagulation. The DCF allows a single channel fiber-optic probe to be shared: i.e. the core propagating the OCT signal while the inner cladding delivers the coagulation laser light. We herein present a novel DCFC designed and built to combine both signals within a DCF (>90% of single-mode transmission; >65% multimode coupling). Potential OCT imaging degradation mechanisms are also investigated and solutions to mitigate them are presented. The combined DCFC-based system was used to induce coagulation of an ex vivo swine esophagus allowing a real-time assessment of thermal dynamic processes. We therefore demonstrate a DCFC-based system combining OCT imaging with laser coagulation through a single fiber, thus enabling both modalities to be performed simultaneously and in a co-registered manner. Such a system enables endoscopic image-guided laser marking of superficial epithelial tissues or laser thermal therapy of epithelial lesions in pathologies such as Barrett’s esophagus.
© 2015 Optical Society of America
Optical coherence tomography (OCT)  is currently investigated as an aid to guide biopsy for the diagnosis and surveillance of Barrett’s esophagus, a pre-malignant lesion that bears an increased risk of progressing to esophageal adenocarcinoma [2, 3]. Image-guided biopsy shows great potential to reduce false-negatives and increase the overall efficiency of the procedure as opposed to performing randomized biopsy sampling, the current surveillance strategy . Imaging of the esophagus is done through an inflated balloon-catheter, which maintains distal optics in the center of the lumen. An optical rotary junction is used to acquire circumferential images of the inner wall of the esophagus, along with a motorized translation stage providing a helical scanning of the laser beam over a 6-cm long segment in about 2 minutes [4, 5]. A second higher power laser source is used to perform laser tissue coagulation leaving on the epithelium marks, that are visible with conventional white light endoscopy . These marks are then used to guide sample collection under endoscopic visualization. Previous coagulation lasers were coupled directly to single-mode fiber (SMF) based OCT systems through wavelength division multiplexers, thus requiring minimal modification to the imaging system. Laser coagulation for tissue marking was demonstrated in vivo using a continuous-wave laser source with 410 mW output power at a wavelength of 1450 nm during 2 s . A pulsed Raman fiber laser was also used to demonstrate single-pulse marking using 900-µs pulses at a wavelength of 1436 nm with pulse energies greater than ~7 mJ , taking an advantage of the temporal confinement of the thermal energy delivery to achieve higher temperature in a shorter time in the optical zone . However, the first configuration required stopping the probe’s rotation to provide continuous irradiation over a few seconds, preventing real-time monitoring of the coagulation process. A pulsed laser source could therefore allow for a more efficient procedure, but the single-mode delivery previously demonstrated required a rather expensive and cumbersome laser source, not commercially available, which, additionally, lead to power densities in the fiber core very close to or higher than the damage threshold of many elements within the balloon-catheter and the rotary junction.
We herein investigate the use of a double-clad fiber (DCF) based system combining OCT – through the single-mode core – and laser tissue coagulation – through the multimode inner cladding – simultaneously and in a co-localized manner. Such a system would allow the use of a less expensive multimode marking laser, an increased mark visibility through a larger beam spot size as well as lower power density travelling through the catheter. Furthermore, the use of DCF presents a clear advantage for multimodality optical systems, compared to multi-fiber setups , as it allows for a more compact and robust probe tip as well as the intrinsic co-registration of different modalities. Such a system could additionally be extended to monitor dynamic processes during laser therapy allowing an increased control over the energy deposition and treatment depth [9–12].
The use of DCFs was previously demonstrated for endoscopy [13, 14] and multimodal imaging combining OCT through the core and using the inner cladding to collect fluorescence or spectroscopic signal in, for instance [15–20]. Furthermore, double-clad fiber couplers (DCFCs) have been demonstrated to efficiently extract inner cladding light out of the DCF [21, 22]. To optimize extraction ratio, our group introduced an asymmetric design, exploiting waveguides presenting uneven optical etendue. This design provides a multimode coupling ratio of >70% while preserving the core signal (≥95%) . Such DCFCs were however optimized for light extraction from the inner cladding and are, therefore, not suitable for the present application, which requires injection of light into the DCF’s inner cladding.
In this paper, we report the use of DCF and a dedicated DCFC enabling simultaneous and co-localized OCT and laser tissue coagulation. The fabrication and characterization of this novel DCFC optimized for simultaneous OCT imaging and multimode light injection are presented. We also quantitatively assess potential OCT image degradation mechanisms due to crosstalk between both channels of the DCF at junction sites and propose strategies for their minimization. Finally, we demonstrate the potential of this combined system for image-guided laser tissue coagulation by acquiring OCT images simultaneously to laser irradiation on ex vivo biological tissue.
2. Materials and methods
2.1 Double-clad fiber coupler
Figure 1 shows the DCFC designed to deliver coagulation laser energy to the OCT imaging site. This design is inspired from a theoretical description of DCFCs introduced by Madore et al. , which showed that power distribution between ports of the coupler is proportional to the ratio of fiber’s etendues. The etendue, , of a fiber is given byEq. (1), is achieved either by a very low NA, a small area or a combination of both. The fabrication technique was further amended by pre-tapering the multimode fiber, therefore allowing multimode expulsion for easier transfer into the DCF’s inner cladding, and by using an asymmetrical fusion geometry resulting in a doubly asymmetrical DCFC, as described in De Montigny et al. .
The proof-of-principle prototype of such a DCFC was fabricated using a commercial DCF (Nufern, SM-9/105/125-20A) for the imaging fiber as it matches the standard SMF used in our OCT system in core and outer cladding dimensions as well as core numerical aperture. A small inner cladding DCF (core, inner cladding and outer cladding diameters of 4.1:25.8:125 µm, respectively) was used as the injection fiber for its multimodal inner cladding, neglecting the core, which does not guide light at long wavelengths. These particular fibers were used as the refractive index of the outer cladding of the injection fiber (which acts as the multimode waveguide in the tapered section) matches that of the inner cladding of the imaging DCF, therefore enabling coupling between these two regions. A custom-built fusion-tapering setup using a micro-torch was used to fabricate the coupler . The injection fiber was first tapered down to an inverse tapering ratio (ITR) of 0.13 such that, in the constant section, most light escapes the multimode inner cladding and leaks into the outer cladding. To further maximize multimode transmission from the injection fiber to the inner cladding of the imaging DCF, the outer cladding of the imaging fiber was removed over a 30-mm section by hydrofluoric acid etching, as described in . Both fibers were then placed side-by-side, the transition section of the injection fiber facing the etched region of the imaging fiber (illustrated in Fig. 1(a) – dashed box) and maintained with a geometry favoring maximum contact during the fusion. The fusion step was stopped when losses in single-mode transmission reached a maximum value of 0.5 dB.
Single-mode transmission was characterized over the wavelength range of our OCT system (~1260–1375 nm) while multimode coupling was measured at 1436 nm, corresponding to the operating wavelength of our coagulation laser. Single-mode transmission was monitored in-line during the fabrication process by injecting a broadband source (Hewlett-Packard, Broadband Light Source, 83437A, 1200–1650 nm) into the DCF core at Port 1 and detected with an optical spectrum analyzer (ANDO Electric, AQ6317). To ensure that only single-mode light was collected, a SMF was spliced to Port 2, therefore filtering out any multimode light. Within the OCT spectral range, ≥90% of the core signal was preserved through the coupler (Fig. 1(c)). Transmission curve presented in Fig. 1(c) was smoothed to eliminate spectral features of the source. Multimode transmission of the coupler was measured using a continuous-wave modulated Raman fiber laser (RFL) built in-house, as described in , which provides a narrow bandwidth spectrum centered at 1436 nm corresponding to a strong absorption peak of water (main constituent of biological tissues), making it suitable for laser coagulation. The RFL light is injected at Port 3 using a lateral offset splice to excite the inner cladding modes of the injection fiber. For an average input power of 300 mW, approximately 66% of the laser power (i.e. ~200 mW average power) was transferred to Port 2 through the inner cladding.
2.2 Experimental setup
The OCT setup used for this work was previously described in [26, 27]. Briefly, it comprises a polygon-based wavelength-swept laser  centered at 1320 nm with a full sweep range of 115 nm. This results in a full width at half maximum (FWHM) of the intensity point spread function of 13 µm in air assuming a Hanning-window for the spectral shape. OCT images were acquired at a rate of 54k axial scans per second providing frames of 4,096 A-lines at a rate of 13 per second, which is suitable for in vivo esophageal imaging as demonstrated in . The system uses an acousto-optic frequency shifter to remove depth-degeneracy, therefore allowing for a full imaging range of ~6 mm, as well as dual-balanced and polarization-diverse detection.
Figure 2 shows a schematic diagram of the experimental setup used to combine OCT imaging with laser tissue coagulation. DCF was integrated to the combined system by splicing the SMF output of the OCT system to the imaging fiber (at S1). OCT illumination is guided to the sample through the core of the DCF (red path in Fig. 2). Coagulation light from the RFL is delivered to the system through the injection fiber at Port 3 using a lateral offset splice (at S4) as described above. The DCFC, as described in the previous section, was used to transfer light from the RFL within the inner cladding of the imaging fiber. An additional length of DCF and a pigtailed angle-polished connector were spliced (S2 and S3, respectively) at Port 2. Port 4 was connected to a beam dump to avoid undesirable back-reflections.
For ex vivo and in vivo imaging, galvanometer-mounted mirrors were used for scanning in the lateral directions. The beams were focused on the sample using a scanning lens (Thorlabs, LSM03) providing an OCT beam spot diameter of 33 µm. The coagulation laser beam spot diameter at the sample (i.e. after the galvanometer-mounted mirrors and scanning lens) was estimated using the 10/90 knife-edge method yielding a width of ~220 µm. A dispersion compensator (Thorlabs, LSM03DC) was used in the reference arm. For coagulation experiments, a manual shutter was placed immediately after the collimator in order to control the irradiation time. To acquire sensitivity measurements, a calibration arm including an achromatic lens was used to ensure optimal light re-collection and precise alignment.
2.3 Image quality
Using DCFs and DCFCs for coherent detection may introduce image artifacts from the crosstalk between inner cladding modes and the fundamental mode (i.e. core mode) occurring at different coupling interfaces. Artifacts occur when inner cladding modes, excited at a primary crosstalk site, propagate within the inner cladding and couple back into the core at a secondary coupling site. Such light travelling through the inner cladding will accumulate a phase shift compared to light travelling through the core. In our system, potential crosstalk sites include fusion splices between the OCT system’s SMF and Port 1 of the DCFC (Fig. 2 – S1), splices between segments of DCF (Fig. 2 – S2-3), the DCF-air interface at the distal tip of the imaging arm and the fused section of the DCFC. At fusion sites, a misalignment between cores may result in the excitation of some inner cladding modes.
This crosstalk-induced delay is proportional to the optical path difference, defined as the product of the physical distance travelled, , with the refractive index difference, , between the effective refractive index of a particular inner cladding mode, , and that of the fundamental mode, . The delayed light may interfere with the reference arm causing a ghost image of the sample, i.e. an attenuated replica of the sample shifted with respect to the main image. The delay, , between the main image and a ghost produced by a certain inner cladding mode is given by :
3.1 Sensitivity analysis
In order to quantify respective artifact contributions from each site, we imaged two samples. Sample 1 is a mirror providing specular reflection and, thus, limited excitation of inner cladding modes at the distal end of the system assuming a proper alignment. Sample 2 is a frosted microscopy slide emulating a scattering sample. Index-matching gel was applied to the clear face of the microscopy slide to avoid contribution from the second glass-air interface. A fibered attenuator was used in the imaging arm to avoid saturation and to control the sample arm power to levels representative of biological tissues. Figure 3(a)-3(f) shows OCT M-mode (i.e. depth profiles as a function of time) images (1,024 A-lines) of Samples 1 (a-c) and 2 (d-f) acquired with an imaging arm composed of standard SMF (a & d), of DCF without attenuation (b & e) and of DCF with attenuation (c & f). For the latter case, attenuations of 8 and 12 dB were used for Sample 1 and 2, respectively, therefore mimicking signal levels representative of biological tissue imaging. The arrow points at signal from a sample positioned at a particular depth within the imaging range while curly brackets identify the region used for noise level measurement. SMF images are artifact free and their curly bracket region defines the benchmark lowest noise floor. When using the DCF, fine lines appear above the signal line. These lines are constant across A-lines and correspond to the excitation of particular cladding modes each associated with a phase delay. As inner cladding modes have a lower effective index of refraction than the fundamental mode, the axial position of the artifact-ridden region is shallower than that of the main signal. The artifact region contains fewer ghost lines for the mirror than for the scattering sample, suggesting that reflected light excites fewer modes, which is consistent with the limited angular span of the specular reflection.
For each sample type, signal and artifact levels are defined as the maximum intensity in the region identified by the arrow and curly brackets, respectively, and averaged over the 1,024 A-lines of the M-mode image. Signal and artifact levels were measured for different axial positions of Sample 1 (Fig. 3(g)) and Sample 2 (Fig. 3(h)) across the axial field of view. SMF and DCF signal collection are similar, however, the artifacts cause the noise floor to go up by up to 8 dB for the reflective sample (Fig. 3(g) blue curve) and by 9 dB for the scattering sample (Fig. 3(h) blue curve) for the regions affected by the artifacts (yellow curly brackets area). When signal levels are attenuated, artifacts disappear below the noise floor for the reflective sample (Fig. 3(g) overlapping red (artifacts) and black (noise) curves) and reach a maximal value of ~3.5 dB above the noise floor for the scattering sample (Fig. 3(h) magenta (artifacts) and black (noise) curves).
To further track the origin of artifacts, we measured the delay between the signal and ghost lines from Fig. 3(b) and 3(e), as ranging from 1.6 to 2.3 mm. This delay can be converted into fiber length, , between crosstalk sites using Eq. (2), by comparing the effective index of refraction of the fundamental mode () to that of higher order modes having an effective refractive index ranging from to . The estimate for responsible for artifacts thus ranges from 10 cm for artifact contribution from higher order modes to 105 cm for artifact contribution from lower order modes, having an effective index closer to that of the fundamental mode. The contribution of each crosstalk site was experimentally assessed observing fluctuations of ghost images when each of these DCF segments was moved, as modes propagating through the inner cladding are highly sensitive to the fiber’s motion. Doing so, we identified DCF segments on each side of the DCFC (S1-DCFC and DCFC-S2), which are the shortest segments in our setup, to be responsible for most of the artifacts. Considering that the excited inner cladding modes are most likely to present an in the upper range, the length of these segments, 70 and 85 cm respectively, correspond well to our prediction.
Figure 4 shows in vivo images of human skin (adjacent to the finger nail bed) acquired with the DCFC-based OCT system. Images were taken with the sample at two different axial locations within the field of view by moving the position of the reference arm to ensure the same focusing conditions. The curly bracket highlights ghosts produced from the tissue-air interface when the sample occupies the lower half of the image (Fig. 4(b)). Signal coming from subsurface structures is too low to produce ghosts above the noise floor. Figure 4(c)-4(d) shows that artifacts are indeed mitigated when index-matching gel is applied on the skin. Artifacts do not alter imaging of sub-surface structures, as they do not overlap with the image of the sample.
3.2 Simultaneous optical coherence tomography imaging and laser coagulation
A preliminary coagulation experiment was performed on a sample of swine esophagus ex vivo to assess the potential of the multimodal system to perform real-time monitoring of tissue laser coagulation. The coagulation process can be monitored through the increase of the scattering properties of the tissue which results from the thermal denaturation of proteins . The RFL was set to provide a duty cycle of 1% and a repetition rate of 20 Hz, resulting in a pulse width of 500 µs. Two different average power settings, 113 mW and 172 mW, were used, providing pulse energies of 5.7 mJ and 8.6 mJ on the sample, respectively. Figure 5 presents images taken before (a & e) and after (c & g) irradiation as well as M-mode images (b & f) acquired in real-time for each power setting. For both power settings, a region of increased scattering at the surface of the tissue can be observed after irradiation. This increased scattering also caused a shadow obscuring the view of deeper layers. M-mode images Fig. 5(b) and 5(f) show the real-time progression of the thermal injury during the irradiation. In Fig. 5(b), for the lower power setting, the coagulation threshold (yellow arrow) is reached ~1.4 s after the beginning of the irradiation (corresponding to 14 pulses), which results in a sudden change in scattering properties. As heat propagates through the tissue over time, tissue coagulation extends to deeper layers. For the highest power setting, the coagulation threshold (yellow arrow) is reached after only 3 pulses (red arrows) as seen in Fig. 5(f). Individual pulses are seen on OCT images as what may be the result of the motion of scatterers due to a thermal excitation. Figures 5(d) and 5(h) show a bright field microscopy image (Olympus, UC500) of the coagulation spots resulting from laser irradiation using the DCFC-based system. The laser-induced marks measure ~315 µm and ~930 µm for the low and high pulse energies, respectively.
Figure 5(i)-5(j) shows the reflected signal intensity as a function of depth (i.e. light attenuation) for the lower (Fig. 5(i) – 5.7 mJ/pulse) and higher (Fig. 5(j) – 8.6 mJ/pulse) power settings at different time points before and after coagulation. Linear fits, solid lines, were performed after averaging 265 consecutives A-lines within each region of interest. In each case, the attenuation slope is more pronounced post-coagulation. Coagulation also results in superficial signal increase due to scattering.
4. Discussion and conclusion
OCT, used along with laser tissue coagulation, shows promising potential to guide biopsy in the context of surveillance of Barrett’s esophagus. However, no system allowing for real-time monitoring of the coagulation laser delivery has been developed yet. Herein, we investigated the use of a DCFC-based system allowing OCT (through the core) and coagulation laser delivery (through the inner cladding) to be performed simultaneously through a single fiber. Such a system allows the use of a multimode laser diode, which is a relatively low-cost commercial option.
We developed a dedicated DCFC enabling injection of the coagulation laser light into the inner cladding of a DCF. This coupler provides single-mode transmission of more than ~90% over the wavelength range of our OCT system as well as 66% multimode coupling at 1436 nm. This novel scheme allows the coupling of high power laser light necessary to induce coagulation, without free space optics, thus providing safe and alignment-free procedures compatible with clinical practice. The first DCFC prototype was developed using fibers on hand and could in theory provide multimode coupling of more than 97%. To minimize multimode losses, we believe that fibers combination, ITR, as well as pre-processing and fabrication techniques could be further improved.
A quantitative sensitivity analysis of the DCFC-based OCT system was performed to assess the effect of crosstalk-induced artifacts on the imaging quality. These artifacts are the result of light that travels in the inner cladding, therefore accumulating a delay compared to the fundamental mode, and interferes with the reference signal. The effect of multimodal crosstalk on OCT imaging was previously assessed for fiber bundles [29, 31] and large-core fibers , but has not been thoroughly assessed in the case of DCF-based imaging. Our analysis showed that the artifacts could be as high as 9 dB above the typical OCT background noise floor when imaging a scattering sample and great care should therefore be taken for their minimization. For a reflectivity comparable to biological tissue, the SNR penalty of using DCF is reduced to ~3 dB as the amplitude of the artifacts is proportional to the tissue reflectivity. Moreover, we have shown that, when imaging a biological tissue, most artifacts arise from the air-tissue interface and that using index-matching gel to reduce reflectivity of this interface is sufficient to completely suppress ghost images under our experimental conditions. In addition, for optimal sample positioning, typically in the first half of the imaging range, we observed that no artifacts are visible and the system SNR is not altered.
As mentioned in [17, 20], another way to manage artifacts is to add enough fiber length to create a delay that would be higher than the depth of the imaging range. For a 6-mm imaging range, this requires adding meters of fibers between each coupling sites. By measuring the signal-ghost delay, we showed that the DCFC was responsible for most artifacts (either as a primary or secondary crosstalk site) such that DCF segments surrounding the coupler are critical. Therefore, it is believed that using a DCFC with much longer ports could push the artifacts out of the imaging range. Even though this condition was not met in the current setup, the delay between the signal and ghost images was sufficient to avoid any overlapping.
Using the DCFC-based system, simultaneous and co-localized OCT imaging and laser coagulation was performed. Our results show that the DCFC-based system allows real-time monitoring of dynamic processes associated with laser coagulation through changes in scattering properties. In addition, laser pulses were seen on OCT images as speckle motion, which may be due to a thermal shock wave. We, as well as others, have demonstrated real-time OCT imaging of thermal effects of laser irradiation [10–12]. Our DCFC-based system however presents the advantage of being intrinsically co-registered and compatible with an endoscopic implementation. Co-registration of the monitoring modality with laser irradiation is of paramount importance in the context of laser marking and therapy as it ensures the safety and the simplicity of the procedure.
Using the real-time feedback provided by OCT, it was possible to determine that the cumulative effect of 3 pulses of ~8.6 mJ each was sufficient to reach the coagulation threshold of esophageal swine tissues. It was shown previously that using single-mode delivery (beam spot size of ~37 µm), a single pulse of more than 7 mJ was sufficient to create a visible coagulation mark . As expected, using the DCF inner cladding to deliver the coagulation laser requires higher energy to yield equivalent results, as the beam spot size (~220 µm) is much larger and, therefore, less spatially confined. The larger beam spot represents the advantage that the mark produced is necessarily larger than that of a single-mode delivery and, therefore, better visible under white-light endoscopy in the context of biopsy guidance. We believe that using a higher peak power along with an adequate combination of pulse width (while maintaining temporal confinement) and repetition rate will allow a single pulse to create a visible mark enabling on-the-fly laser marking. Further optimization of the DCFC multimode coupling would likewise increase the pulse energy reaching the sample.
The RFL was used for this proof-of-principle, as it was readily available. Based on the results presented here, it should be possible to replace the RFL with a multimode diode laser, significantly reducing cost and complexity while also enabling simple control and synchronization of pulse parameters.
In conclusion, we have herein demonstrated a DCFC-based system using a novel DCFC to simultaneously combine OCT imaging with laser coagulation for the first time, to the best of our knowledge, through a single fiber. The combined system is intrinsically co-registered and induces minimal imaging quality degradation. In combination with suitable laser source, this system has the potential to enable on-the-fly laser marking with concurrent OCT guidance and monitoring of the marking process.
The authors thank Mikael Leduc for precious laboratory help as well as Etienne De Montigny and Mathias Strupler for fruitful discussions. This work was supported in part by Ideas to Innovation Grant from the Natural Sciences and Engineering Research Council of Canada (NSERC) and by the National Institutes of Health, grant P41 EB015903. KB and WJM were supported by a NSERC graduate scholarship; KB is additionally supported by the Fonds de recherche du Québec – Nature et Technologies (FRQ-NT) International Training Program.
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