2.1.

Simulation and Validation of Ball Lens Endoscope FLIM Illumination Performance

Prior to fabrication, we simulated the performance of the endoscope via the optical ray-tracing program OpticStudio (Zemax) to optimize the ball lens design. A polished ball lens at the end of fiber endoscope is used to focus and reflect light, as shown in Fig. 1. Simulation is important to minimize manufacturing time by establishing a preferred shape for achieving target optical properties with the partial ellipsoid lens. In the simulation, the $1.3\text{-}\mu \mathrm{m}$ OCT light was considered to be completely contained within the single mode core of the DCF. The light exiting the core was simulated with Gaussian beam propagation after the ball lens to predict beam details for both illumination and collection, which is more accurate for side-viewing, noncircularly symmetric, and nonparaxial ball lenses than the ABCD matrix method.^23,24 The ball lens shape was optimized according to a merit function considering optical performance and tolerance analysis. Once manufactured, it was straightforward to directly measure the output of the ball lens using a beam profiler and $1.3\text{-}\mu \mathrm{m}$ laser diode illumination for comparison.

Fig. 1

(a) Schematic drawing of the catheter. Heated and shaped CLF fusion spliced with an MMF and then with a DCF, and the fiber was glued with torque coil and silica tube. Torque coil’s i.d. is $270\mu \mathrm{m}$ and o.d. is $450\mu \mathrm{m}$. Silica tube’s i.d. is $530\mu \mathrm{m}$ and o.d. is $620\mu \mathrm{m}$. Scale bar: $100\mu \mathrm{m}$. DCF, double-clad fiber; MMF, multimode fiber; CLF, coreless fiber. (b) Side view of the fabricated ball lens before polishing. Heated and shaped CLF fusion spliced to an MMF [DCF was not shown due to small field of view (FOV) of the machine]. Scale bar: $50\mu \mathrm{m}$. (c) Side view of the ball lens after polishing by customized fiber polishers. The polishing angle is 41 deg.

FLIM excitation was considered to transmit through both the core and the inner cladding of the DCF. We simulated the beam profile including spatial mixing within the multimode fiber (MMF). Mixing of spatial modes within the MMF leads to rapid variations in the intensity profile exiting the fiber, which poses a problem both for adequate simulation and finally comparison with measured ball lens performance. To overcome this issue, we sought to identify an intensity profile that could be modeled in OpticStudio as well as reliably generated and measured using standard optical techniques. We chose the super-Gaussian profile that approximates a top-hat profile with Gaussian tails on either side, which could fit the beam profile in both near field and far field. The intensity profile is $I(r)=I_0\exp [-2{(r/w_m)}^n]$, where $r$ is the radial dimension, $w_m$ is the beam waist, and $n$ is the order of the super-Gaussian. A model intensity profile was generated by measuring the intensity profile from the core and inner claddings of a segment of DCF using a beam profiler. The angular distribution was fit to a super-Gaussian using a series of five far-field measurements at intervals of $508\mu \mathrm{m}$. These measurements were done nine times to test for reproducibility, each time removing and replacing the fiber. The fit of the super-Gaussian proved to be robust to fiber bending and readily reproducible with a standard deviation of 1% or less for the nine measurements at the five intervals ($n=45$). We used this to approximate the intensity profile impinging on the coreless fiber (CLF), which was modeled based on the source-radial using nonsequential mode with OpticStudio.

The final piece needed for simulation is the ball-lens design itself. The ball-lens was drawn in Solidworks with the desired ellipsoidal shape and imported into OpticStudio. Using the approaches described above, the performance for both OCT and FLIM could be modeled. Likewise, since we directly used the super-Gaussian intensity, we measured from the DCF to simulate the ball-lens, thus we could reliably test performance of the ball-lens after fabrication.

2.2.

Design of Ball Lens Endoscope

A fiber preparation and splice workstation (FFS2000, Vytran) was used for endoscope fabrication. The 1.55-m OCT/FLIM endoscope was constructed starting with a DCF, SM-9/105/125-20A by Nufern. The DCF was fusion spliced to an MMF (FG105UCA, Thorlabs). The MMF was cleaved to a precise length and then fusion spliced to a CLF (HCF-125-47, Coractive). Following splicing, the CLF was cleaved to $1250\text{-}\mu \mathrm{m}$ length, heated and shaped into an ellipsoid ball with a total CLF length of $632\mu \mathrm{m}$ and predetermined semiprinciple axes (140.6 and $123.1\mu \mathrm{m}$) in the tungsten filament furnace of the workstation, shown in Fig. 1(b). A schematic drawing is shown in Fig. 1(a). The ellipsoidal shape was designed to compensate for the astigmatism caused by the toroidal shape of the catheter sheath and the silica tube used to cover and protect the fiber endoscope. The distal ball lens was polished with an angle of 41 deg using a commercial fiber polisher (69-3000-160, Buehler) with custom modifications to enable ball lens polishing at precise angles for side viewing endoscopes. The 41-deg angle was chosen so that total internal reflection would occur at the polished glass–air interface, hence obviating the need for reflective coatings on the polished surface. In addition to the angle, precise control of the polishing depth from an unpolished ball is necessary to achieve good performance. Images taken before and after polishing are shown in Figs. 1(b) and 1(c).

The fiber endoscope was then enclosed in a customized torque coil (ASAHI INTECC) with high-speed torque transmission and high flexibility. The outer diameter (o.d.) of the torque coil is $450\mu \mathrm{m}$ and the ID is $270\mu \mathrm{m}$. The proximal end was connectorized with a custom FC/APC connector and strain relief boot, mating to a custom broadband fiber optic rotary joint.¹⁶ The distal end was covered by a laser-cut silica tube (TSP530620, Polymicro) with an outer diameter of $620\mu \mathrm{m}$ and an inner diameter of $530\mu \mathrm{m}$, shown in Fig. 2. The silica tube serves both to protect the ball lens and preserve the air–glass interface at both the flat polished and curved surfaces of the endoscope even when submerged. The i.d. of the silica tube was chosen so that the torque coil could fit inside, and therefore, make a strong bond between the two. Using a smaller diameter for the silica tube would unnecessarily increase the astigmatism introduced by the tube. The fiber endoscope was inserted into a custom designed flexible housing with an i.d. of 0.812 mm and o.d. of 1.066 mm.

Fig. 2

(a) Ball lens enclosed within silica tube. Both ends of the silica tube were sealed by epoxy. (b) Image of a catheter endoscope, the proximal end was connectorized by FC/APC connector.

2.3.

OCT/FLIM Instrumentation

A schematic diagram of the dual-modality imaging system is shown in Fig. 3. The diagram can be divided into the OCT module, FLIM module, and common path. The common path refers to the portion of the optical system that is common to both OCT and FLIM, i.e., the respective beams are copropagating.

Fig. 3

Schematic diagram of the dual-modality imaging system, including OCT module, FLIM module, and common path. L1 to L3, L7 to L9, free-space collimation and coupling lenses; L4, reflective collimator; L5, L6, fiber-connected collimation and coupling lenses; M, mirror; DM1 to DM5, dichroic mirrors; F1 to F3, bandpass filters; PL, polarizer; BD, balanced detectors; Cir, circulators; P1 to P3, circulator ports 1 to 3; APD1 to APD3, avalanche photodiodes; RT, rotary joint. Orange line indicates MMF fiber transmission. Yellow line indicates SMF transmission. Other thick colored lines indicate free-space transmission. Black thin line indicates electrical line.

2.4.

Tissue Preparation and Imaging

Human coronary artery segments were obtained from autopsy cases within 5 days of the time of death according to a protocol approved by the Texas A&M University Institutional Review Board. The arterial segment was imaged with helical beam scanning by rotating the endoscope inside the lumen of the artery while pulling the endoscope back. Immediately after imaging, each segment was ink marked for correlation with histopathology, fixed in 10% formalin, and sent for histopathology analysis. Each imaged artery segment was consecutively sectioned every 1 mm. The sections were stained with Movat pentachrome and CD68. Movat stains collagen, elastin, muscle, mucin, and fibrin in tissue sections and is commonly used for coronary arteries. CD68 stains macrophages.

3.1.

Solarization and Absorption Losses in the Fiber Endoscope

Transmission loss due to absorption and solarization are important when coupling high power UV light into optical fibers as we do for FLIM excitation. As has been reported previously, high intensity pulses suffer greater losses due to two-photon absorption.^17,27 Two-photon absorption has been shown to be the dominant process for the formation of radiation-induced defects in silica material.¹⁸ Similarly, solarization and thermal cumulative effects are known to become significant for high power UV transmission through the core.^17,19,28 For our system, the intensity of light can be up to $0.1\mathrm{MW}/{\mathrm{cm}}^2$ if transmitting through the core. Therefore, both processes should be considered carefully since they can degrade the fiber transmission and affect the optical performance.

The fiber transmission, considering one- and two-photon absorption, can be described as

Fig. 4

One-photon absorption and two-photon absorption at 375 nm. Curves are fitting of Eq. (1). Transmission of the inner cladding is measured based on 40-cm DCF, and transmission of the core is measured based on 59.5-cm DCF.

Ultraviolet radiation degrades the fiber transmission (solarization), especially in high-OH fiber. The solarization rate of the DCF is important for our application since we want to be confident that the optical properties of the system are not changing over the course of an imaging session. In order to measure the solarization rate of the core, we used the following technique. An SMF (SMF28e, Corning) patch cable was presolarized by illumination with 50 mW of 375 nm light so that the transmission was stable and solarization recovery did not happen after 1 h. This patch cable was connected to a 1.55-m DCF patch cable so that the light from the core of the SMF was reliably transmitted to the core of the DCF. The output power from the DCF was then measured as a function of time to produce the curves in Fig. 5. As expected, the solarization rate increases when the input power to the DCF is increased. We tried a similar experiment on the inner cladding; however, we were not able to measure any change in transmission over the time course of the experiment, hence no significant solarization. We also measured the absorption coefficients using the same cut-off method noted above after the solarization experiment. The one-photon coefficient for the core more than doubled to $7.26\times {10}^{-3}{\mathrm{cm}}^{-1}$, while the one-photon coefficient remained approximately the same for the inner cladding, $1.61\times {10}^{-3}{\mathrm{cm}}^{-1}$. The small increase in the one-photon absorption can be explained by the fact that the measurements on the inner cladding include light that propagates through the core. The two-photon coefficients cannot be measured because the 50-mW incident power cannot induce a strong enough intensity for two-photon coefficients measurements when the one-photon absorption is so high.

Fig. 5

Solarization rate of DCF core. The DCF core was solarized when the power was higher than 0.7 mW. Solarization rate increases with power.

It would be advantageous to illuminate the sample through the core in order to achieve $\sim 4\times$ higher spatial resolution than illumination through the inner cladding. However, given that an imaging session will likely require the illumination of the catheter for 10s of seconds, if we excite through the core, we will lose about 30% to 40% of the excitation light due to solarization. Since collection of the fluorescence is largely dependent on the inner cladding transmission, it should not be significantly impacted. Hence given a bright enough light source, it should be possible to illuminate through the core to gain higher spatial resolution. More generally, if longer wavelength illumination can be used, it would also mitigate the solarization issue.

3.2.

FLIM Resolution and Mechanical Testing

Given that we were testing a new system with imaging times that would lead to strong solarization and that DCF have been fabricated to guide UV either entirely through the core or the inner cladding,²⁹ we chose to primarily illuminate through the inner cladding. We aligned the optical system to adjust the output power ratio between core (7%) and the inner cladding (93%) for a trade-off between spatial resolution and output power.

We optimized the mode mixing in the DCF to generate a super-Gaussian profile. We found this to be stable even when the endoscope fiber is rotating rapidly. The stability of the super-Gaussian output gave us confidence that we could expect a similar intensity profile at the interface of the CLF and the DCF in the endoscope. The intensity profiles from the core and the inner cladding were modeled based on two source-radial sources. The collection efficiency was also modeled for resolution estimation. The three-dimensional (3-D) layout from OpticStudio including rays from the core (green) and inner cladding (blue) is shown in Fig. 6(a). The plane (orange) represents the receiver where the beam profile was simulated. The resolution at 1 mm from the center of the fiber is shown in Fig. 6(b). The XY section FWHM resolution and YZ section FWHM resolution were 99 and $88\mu \mathrm{m}$, respectively, while 70% [* in Fig. 6(b)] power was transmitting out of the core and 30% power was transmitting out of the inner cladding. Under our planned experimental conditions, i.e., 7% power transmitting out of the core, the XY section resolution and YZ section resolution were 111 and $99\mu \mathrm{m}$, respectively [* in Fig. 6(b)]. Nominally, the resolution can be tuned by adjusting the relative power transmitting through the core and the inner cladding.

Fig. 6

OpticStudio simulation FLIM endoscope performance. (a) 3-D layout of OpticStudio simulation. The gray object is the fabricated ball lens imported from SolidWorks. Blue rays are random layout rays launching from the inner cladding, and green rays are random layout rays launching from the core. The orange plane represents the receiver that was translated along the beam bath to extract the resolution plotted in (b). (b) Simulation of resolution of XY section and YZ section at 1 mm from the center of the fiber with different power transmitting out of the core. Asterisks indicate simulated resolution with two scenarios mentioned in the text.

In order to confirm the resolution of the endoscope under realistic imaging conditions, a resolution target (53-715, Edmund Optics) was wrapped around a silica tube with 1-mm i.d. and 2-mm o.d. The white paper has high fluorescence while the black ink has low fluorescence. The fluorescence intensity maps were used to define FLIM spatial resolution. The tube was filled with phosphate-buffered saline (PBS) and the endoscope inserted. Several pullback images were acquired with representative intensity maps shown in Fig. 7. In Fig. 7(a), the endoscope was rotated at 2358 rpm and pulled back at a rate of $0.4\mathrm{mm}/\mathrm{s}$. Under these conditions, the image sampling is $9.5\mu \mathrm{m}$ along the radial dimension and $10.2\mu \mathrm{m}$ along the longitudinal dimension at the surface of the resolution target. The fluorescence is attenuated in areas with ink hence, the bars show up dark in this gray scale image. We can clearly differentiate bars with a width of $111\mu \mathrm{m}$ (group 2 element 2) in the rotation direction and $99\mu \mathrm{m}$ (group 2 element 3) in the longitudinal direction. In Fig. 7(b), the endoscope was rotated at 2358 rpm and pulled back at a rate of 0.35 mm/s. Under these conditions, the image sampling is $9.5\mu \mathrm{m}$ along the radial dimension and $8.9\mu \mathrm{m}$ along the longitudinal dimension at the surface of the resolution target. We can clearly differentiate bars with a width of $99\mu \mathrm{m}$ (group 2 element 3) in the rotation direction and $88\mu \mathrm{m}$ (group 2 element 4) in the longitudinal direction. Experimentally, we expect deviations from the simulation based on the fact that we are imaging through a silica tube which has a higher refractive index than saline, and therefore, should lead to a smaller beam waist. Likewise, the endoscope is free to deviate from the center of the silica tube during rotation.

Fig. 7

Pullback fluorescence intensity images of resolution targets. (a) The bars show up dark in this gray scale image, and we can clearly differentiate bars with the width of $99\mu \mathrm{m}$ radially and $111\mu \mathrm{m}$ longitudinally while 7% power transmitting out of the core. (b) We can clearly differentiate bars with the width of $88\mu \mathrm{m}$ radially and $99\mu \mathrm{m}$ longitudinally while 70% power transmitting out of the core.

3.3.

Fluorescence Lifetime Imaging

In order to validate the temporal and spectral calibration of the catheter imaging system, we acquired images of solutions of well-characterized fluorophores. The distal end of the catheter ($\sim 1\mathrm{cm}$) was immersed in a solution of 0.1-mM POPOP, 1-mM NADH, and 1-mM FAD, respectively. The emission was spectrally resolved into three channels centered at 405, 440, and 525 nm. A spectral intensity calibration to compensate for wavelength-dependent losses was performed on the three spectral bands based on the measured POPOP spectrum, known transmission curves of filters and dichroic mirrors, and the POPOP spectrum¹³ from the literature. The calculated lifetimes were $1.35\pm 0.04\mathrm{ns}$ for POPOP (note, also used as standard), $0.56\pm 0.23\mathrm{ns}$ for NADH, and $2.14\pm 0.40\mathrm{ns}$ for FAD. The literature values for the average lifetime of POPOP, NADH, and FAD are $1.38\pm 0.01$,³⁰ 0.44,³¹ and 2.07 ns,³² respectively. The results agree with literature values within one standard deviation of our measurement except POPOP, which differed by slightly more than one standard deviation. We consider this is to be good agreement.

3.4.

Ex Vivo Human Coronary Artery and Histology Analysis

Finally, we imaged a segment of human coronary artery obtained from autopsy. The catheter endoscope was inserted into the flexible catheter which was then placed in the artery lumen. A pullback at 1.87 mm/s was performed at a rotational velocity of 3000 rpm. A representative OCT image and overlapped fluorescence lifetime map are shown in Figs. 8 and 9. After completing imaging, the artery segment was fixed in formalin and sent for histopathology processing.

Fig. 8

(a) OCT/FLIM cross-sectional image of ex vivo of human coronary artery obtained from autopsy. The OCT is rendered in gray scale. The four color rings on the inside of the lumen are from inner to outer normalized intensity maps from channels 1, 2, 3 and lifetime map (channel 3 at 40 MHz). (b) Histopathology result (Movat staining). (c) Histopathology result (CD68 staining). Red arrow provides a point of reference for all four images.

Fig. 9

(a) OCT/FLIM cross-sectional image of ex vivo of human coronary artery obtained from autopsy. The OCT is rendered in gray scale. The four color rings on the inside of the lumen are from inner to outer normalized intensity maps from channels 1, 2, 3 and lifetime map (channel 3 at 40 MHz). (b) Histopathology result (Movat staining). (c) Histopathology result (CD68 staining). Red arrow points at a fibroatheroma with a thick cap and a small necrotic core (shown in zoomed view). Blue arrow points at the thickened intima.

The histopathology (Fig. 8) indicates a fibrotic plaque rich in collagen and smooth muscle cells around the entire lumen with no observable macrophage infiltration. The intravascular OCT (IVOCT) figure shows a layered architecture, with highly backscattering intima and adventitia and low backscattering media. The intima area is region of thickened collagen, which has lifetime of ~5 ns.

The histopathology of the second artery sample (Fig. 9) shows a fibroatheroma with a thick cap rich in collagen and smooth muscle cell covering a small necrotic (NC) [red arrow in Fig. 9(b)] showing extensive macrophage infiltration [red arrow in Fig. 9(c)]. The region next to the fibroatheroma shows a thickened intima rich in collagen [blue arrow in Fig. 9(b)] and minimal macrophage infiltration [blue arrow in Fig. 9(c)]. The OCT figure overlapped with FLIM lifetime map shows two distinct regions resembling: (a) a region (blue arrow) of thickened collagen (homogeneous IVOCT signal low signal drop-off and fluorescence lifetime $\sim 5\mathrm{ns}$) and (b) a region (red arrow) of combination of collagen and lipid (a fast IVOCT signal drop-off and fluorescence lifetime $>6\mathrm{ns}$¹⁴) reflecting the presence of the lipid-rich necrotic core and macrophage.