**3.4. Optical Coherence Tomography (OCT)**

that a RMS wavefront error as low as *λ*/10 (@527.5 nm) can be obtained, which is a signifi‐ cant result for a sensorless AO system. The correction was not particularly effective only in those cases in which the unidirectionality of the mirror deformation did not allow aberration compensation, as in case of astigmatism. However, with a different choice of AO system, the

**Figure 11.** Wavefront plots (obtained with the wavefront analysis, WFA) before and after the correction applied by the deformable mirror for three considered cases. Top: the main aberration is defocus; Middle: the main aberrations are astigmatism and coma; Bottom: the main aberrations are defocus and astigmatism. The blue dashed lines over

plotted to the Z axis represent the total wavefront excursion.

system is very effective in identifying and correcting aberrations.

56 Adaptive Optics Progress

Optical coherence tomography, OCT, is an imaging modality allowing acquisition of micro‐ meter-resolution three-dimensional images from the inside of optical scattering media (e.g. biological tissue). OCT is analogous to ultrasound imaging, except that it makes use of light instead of sound. It relies on detecting interferometric signal created by the light back scat‐ tered from the sample and from a reference arm in a Michelson or Mach-Zehnder interfer‐ ometer. OCT has many applications in biology and medicine and can be treated as a sort of optical biopsy without requirement of tissue processing for microscopic examination.

One of the interesting features of OCT is that, unlike in most optical imaging techniques, the axial and lateral resolutions are decoupled, thus allowing for an improved axial resolution, which is independent of transverse resolution. The axial resolution Δz is determined by the roundtrip coherence length of the light source and can be calculated from the central wave‐ length (λ0) and the bandwidth (∆λ) of the light source as [37]:

$$
\Delta z = \frac{2\ln 2}{\pi} \frac{\lambda\_0 z^2}{\Delta \lambda}.
$$

The lateral resolution (∆x) in OCT is defined similarly to the confocal scanning laser oph‐ thalmoscopy (cSLO), since OCT is based on a confocal imaging scheme. In many imaging systems, however, the confocal aperture exceeds the size of the Airy disc, which degrades the resolution to the value known from microscopy, i.e. [38]:

$$
\Delta \mathfrak{x} = 1.22 \lambda \frac{f}{D}.
$$

Therefore, as for standard microscopy, AO enhanced devices might be necessary to achieve diffraction limited transverse resolution. As a result, only a combination of OCT with AO has the potential to achieve high and isotropic volumetric resolution. The use of broadband light sources that are necessary for OCT and the complexity of both the AO and the OCT technique, make the combination very challenging [39]. In general, any AO-OCT instrument can be divided into two subsystems: an adaptive optics subsystem, with wavefront sensing and wavefront correction, and an interferometric OCT subsystem. In every implementation of AO-OCT all the elements of the AO subsystem are located in the sample arm of the OCT interferometer. Indeed, there is no need to have AO correction in the reference arm because aberrations introduced within this part of the system will not influence the transverse reso‐ lution of the image. In most of the AO-OCT systems, a Shack–Hartmann wavefront sensor is used to measure aberrations and, then, to control adaptive optics correction.

To test the performance of our sensorless AO-OCT system, we evaluated the image quality of a sample, consisting of a USAF resolution test chart with an adhesive tape glued to its front side, after insertion of a trial lens with 0.5 Diopter astigmatism in front of the imaging objective. We were able to achieve improved resolution by using the following merit func‐

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59

(*x*, *y*)*dxdy*,

where I(x,y) is the intensity in the OCT en-face image plane. This approach is simillar to PSF optimization. In fiber based OCT systems single mode fiber introduces OCT beam to the sample and also act as detector for back scattered light. Therefore we have a point source that is imaged by the optical system and the confocal pinhole that allows direct mesurment of light intensity trougput by the system. As expected, the algorithm performed the optimi‐

Fig. 14 shows some examples of the en-face projection views extracted from OCT volumes: there are the initial view acquired from the sample, and three improved views after correc‐ tion of additional aberrations, namely, defocus and two astigmatisms. Clearly, at each cor‐ rection step the images of the test target get sharper. Additionally, the features of the adhesive tape attached to the back of the Air Force test target become more visible as well.

**Figure 13.** Graph of the Merit Function of AO-OCT images for different values of aberrations generated by the modal

deformable mirror. Note that higher values correspond to better AO-corrections.

*S* =*∫I* <sup>2</sup>

tion S [41] on the OCT en-face projection images

zation by adjusting only defocus and astigmatism (see Fig. 13).

Bonora and Zawadzki recently demonstrated that sensorless correction can be implemented in optical coherence tomography by using a specially developed resistive deformable mir‐ ror. This novel modal deformable mirror, MDM, was successfully employed in the UC Da‐ vis AO-OCT system to image static samples, test targets and tissue phantoms. Fig. 12 shows a schematic representation of the sensorless AO-OCT system used in the experiments.

**Figure 12.** Schematic representation of the system for sensorless adaptive optics - optical coherence tomography. Note that there is no wavefront sensor in the sample arm. The far-field camera (FF) is used to check if the AO correc‐ tion generates improved focal spots. DM : deformable mirror; V: vertical mirror galvanometer; H: horizontal mirror gal‐ vanometer. In the reference arm: NDF is a neutral density filter. The detection channel comprises a grating (DG) and a linear CCD detector (LSC). The quality of the image acquired with the OCT detection channel is used to search for DM shapes that correct aberrations in the imaged sample. The imaging system used to acquire the data was developed in the Vision Science and Advanced Retinal Imaging Laboratory (VSRI). Details of the OCT system components can be found in [40]. Here, we briefly describe the main characteristics of the system. In the current configuration, the light source for OCT was a superluminescent diode (Broadlighter) operating at 836 nm and with a 112 nm spectral band‐ width (Superlum LTD), allowing to achieve a 3.5 μm axial resolution. The beam diameter at the last imaging objective was 6.7 mm, allowing for up to 10 μm lateral resolution when a 50 mm focal length imaging objective was used. The AO correction was optimized by using the intensity of the AO-OCT en-face projection views during the volumetric da‐ ta acquisition. In the current system configuration, we have used about 9 mm diameter of the modal deformable mir‐ ror. The light reflected from the sample is combined with the light from the reference mirror, and then sent to a spectrometer. There, a CCD line detector acquires the OCT spectrum.

To test the performance of our sensorless AO-OCT system, we evaluated the image quality of a sample, consisting of a USAF resolution test chart with an adhesive tape glued to its front side, after insertion of a trial lens with 0.5 Diopter astigmatism in front of the imaging objective. We were able to achieve improved resolution by using the following merit func‐ tion S [41] on the OCT en-face projection images

interferometer. Indeed, there is no need to have AO correction in the reference arm because aberrations introduced within this part of the system will not influence the transverse reso‐ lution of the image. In most of the AO-OCT systems, a Shack–Hartmann wavefront sensor is

Bonora and Zawadzki recently demonstrated that sensorless correction can be implemented in optical coherence tomography by using a specially developed resistive deformable mir‐ ror. This novel modal deformable mirror, MDM, was successfully employed in the UC Da‐ vis AO-OCT system to image static samples, test targets and tissue phantoms. Fig. 12 shows

**Figure 12.** Schematic representation of the system for sensorless adaptive optics - optical coherence tomography. Note that there is no wavefront sensor in the sample arm. The far-field camera (FF) is used to check if the AO correc‐ tion generates improved focal spots. DM : deformable mirror; V: vertical mirror galvanometer; H: horizontal mirror gal‐ vanometer. In the reference arm: NDF is a neutral density filter. The detection channel comprises a grating (DG) and a linear CCD detector (LSC). The quality of the image acquired with the OCT detection channel is used to search for DM shapes that correct aberrations in the imaged sample. The imaging system used to acquire the data was developed in the Vision Science and Advanced Retinal Imaging Laboratory (VSRI). Details of the OCT system components can be found in [40]. Here, we briefly describe the main characteristics of the system. In the current configuration, the light source for OCT was a superluminescent diode (Broadlighter) operating at 836 nm and with a 112 nm spectral band‐ width (Superlum LTD), allowing to achieve a 3.5 μm axial resolution. The beam diameter at the last imaging objective was 6.7 mm, allowing for up to 10 μm lateral resolution when a 50 mm focal length imaging objective was used. The AO correction was optimized by using the intensity of the AO-OCT en-face projection views during the volumetric da‐ ta acquisition. In the current system configuration, we have used about 9 mm diameter of the modal deformable mir‐ ror. The light reflected from the sample is combined with the light from the reference mirror, and then sent to a

spectrometer. There, a CCD line detector acquires the OCT spectrum.

a schematic representation of the sensorless AO-OCT system used in the experiments.

used to measure aberrations and, then, to control adaptive optics correction.

58 Adaptive Optics Progress

$$S = \left[ I^{-2}(x, y) dx dy, y \right]$$

where I(x,y) is the intensity in the OCT en-face image plane. This approach is simillar to PSF optimization. In fiber based OCT systems single mode fiber introduces OCT beam to the sample and also act as detector for back scattered light. Therefore we have a point source that is imaged by the optical system and the confocal pinhole that allows direct mesurment of light intensity trougput by the system. As expected, the algorithm performed the optimi‐ zation by adjusting only defocus and astigmatism (see Fig. 13).

Fig. 14 shows some examples of the en-face projection views extracted from OCT volumes: there are the initial view acquired from the sample, and three improved views after correc‐ tion of additional aberrations, namely, defocus and two astigmatisms. Clearly, at each cor‐ rection step the images of the test target get sharper. Additionally, the features of the adhesive tape attached to the back of the Air Force test target become more visible as well.

**Figure 13.** Graph of the Merit Function of AO-OCT images for different values of aberrations generated by the modal deformable mirror. Note that higher values correspond to better AO-corrections.

**Figure 14.** En-face projection views of the AO-OCT images of the test target for the best corrected values of the Zer‐ nike coefficients; (a) before correction, (b) after defocus correction, (c) after defocus and Ast 0° correction, (d) after defocus, Ast 0° and Ast 45° correction.

These recent results demonstrate that wavefront sensorless control is a viable option for imaging biological structures for which AO cannot establish a reliable wavefront that could be corrected by a wavefront corrector. Future refinements of this technique, beyond the sim‐ ple implementation presented in this chapter, should allow its extension to in-vivo applica‐ tions. An example of sensorless adaptive optics scanning laser ophthalmoscopy (AO-SLO) for imaging in-vivo human retina has been recently presented [42].

> **Figure 15.** Experimental setup for the generation of harmonics from a femtosecond tunable high-energy mid-IR opti‐ cal parametric amplifier, OPA. Dotted line: optical path before the insertion of the MDM. Red line: optical path realiz‐

> > Test 4 - 5th Harmonics 290nm, 500V

Defocus Ast 0 Ast45 Spherical Ab Coma 0 Coma 90

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Aberration (-1,1)

**Figure 16.** Optimization of the voltage generated by the photomultiplier over a 50 Ω load for the 5th harmonic at 290

ed for the experiment with the deformable mirror.

0.1

nm, obtained by the use of krypton gas.

0.2

0.3

0.4

0.5

0.6

Photomultiplier Voltage (V)

0.7

0.8

0.9 1

### **3.5. Laser process optimization**

Similarly to the optimization process presented in section 2.1.1 [24], we report here about the optimization of a laser process by the use of a sensorless AO [43]. In the former case, the generation of harmonics from an ultrafast laser was improved by the use of a genetic algo‐ rithm. In the latter case, an algorithm derived from the image-based procedure was em‐ ployed in conjunction with the use of a MDM deformable mirror similar to the one described in section 3.1. The advantages in terms of experimental complexity and conver‐ gence time are discussed in the given reference.

In the sensorless case, the laser source was a tunable high energy mid-IR (1.2µm-1.6µm) op‐ tical parametric amplifier with 10 Hz repetition rate [44]. The harmonics of the laser were generated by the interaction of the laser pulses with a krypton gas jet. In this system, the infrared pulses and the slow repetition rate made inconvenient, respectively, the use of a wavefront sensor and of an optimization algorithm needing hundreds of iterations.

The experimental setup used for this application is illustrated in Fig. 15. To demon‐ strate the easiness of integrating the sensorless AO device within the experiment, the optical path before the DM is shown with a dotted line. The additional elements are simply a plane mirror and a resistive MDM, which have been introduced without any complex operations. The system optimization consisted in the increase of the harmonic signal detected by the photomultiplier at the output of the monochromator. The ob‐ tained result is illustrated in Fig. 16, where it is possible to see that the photon flux on the photomultiplier is doubled with respect to the one obtained after the correction of the defocus.

**Figure 14.** En-face projection views of the AO-OCT images of the test target for the best corrected values of the Zer‐ nike coefficients; (a) before correction, (b) after defocus correction, (c) after defocus and Ast 0° correction, (d) after

These recent results demonstrate that wavefront sensorless control is a viable option for imaging biological structures for which AO cannot establish a reliable wavefront that could be corrected by a wavefront corrector. Future refinements of this technique, beyond the sim‐ ple implementation presented in this chapter, should allow its extension to in-vivo applica‐ tions. An example of sensorless adaptive optics scanning laser ophthalmoscopy (AO-SLO)

Similarly to the optimization process presented in section 2.1.1 [24], we report here about the optimization of a laser process by the use of a sensorless AO [43]. In the former case, the generation of harmonics from an ultrafast laser was improved by the use of a genetic algo‐ rithm. In the latter case, an algorithm derived from the image-based procedure was em‐ ployed in conjunction with the use of a MDM deformable mirror similar to the one described in section 3.1. The advantages in terms of experimental complexity and conver‐

In the sensorless case, the laser source was a tunable high energy mid-IR (1.2µm-1.6µm) op‐ tical parametric amplifier with 10 Hz repetition rate [44]. The harmonics of the laser were generated by the interaction of the laser pulses with a krypton gas jet. In this system, the infrared pulses and the slow repetition rate made inconvenient, respectively, the use of a

The experimental setup used for this application is illustrated in Fig. 15. To demon‐ strate the easiness of integrating the sensorless AO device within the experiment, the optical path before the DM is shown with a dotted line. The additional elements are simply a plane mirror and a resistive MDM, which have been introduced without any complex operations. The system optimization consisted in the increase of the harmonic signal detected by the photomultiplier at the output of the monochromator. The ob‐ tained result is illustrated in Fig. 16, where it is possible to see that the photon flux on the photomultiplier is doubled with respect to the one obtained after the correction of

wavefront sensor and of an optimization algorithm needing hundreds of iterations.

for imaging in-vivo human retina has been recently presented [42].

defocus, Ast 0° and Ast 45° correction.

60 Adaptive Optics Progress

**3.5. Laser process optimization**

the defocus.

gence time are discussed in the given reference.

**Figure 15.** Experimental setup for the generation of harmonics from a femtosecond tunable high-energy mid-IR opti‐ cal parametric amplifier, OPA. Dotted line: optical path before the insertion of the MDM. Red line: optical path realiz‐ ed for the experiment with the deformable mirror.

**Figure 16.** Optimization of the voltage generated by the photomultiplier over a 50 Ω load for the 5th harmonic at 290 nm, obtained by the use of krypton gas.
