We report on the fabrication by a femtosecond laser of an optofluidic device for optical trapping and stretching of single cells. Versatility and three-dimensional capabilities of this fabrication technology provide straightforward and extremely accurate alignment between the optical and fluidic components. Optical trapping and stretching of single red blood cells are demonstrated, thus proving the effectiveness of the proposed device as a monolithic optical stretcher. Our results pave the way for a new class of optofluidic devices for single cell analysis, in which, taking advantage of the flexibility of femtosecond laser micromachining, it is possible to further integrate sensing and sorting functions.
© 2010 OSA
Biophotonic devices based on optical forces are powerful tools for single cell study and manipulation without physical contact [1–3]. The exploitation of optical forces for analysis at single cell level provides significant information, opening new scenarios for the comprehension of basic biological mechanisms and for the early detection of several diseases. In particular, the investigation of the viscoelastic properties of trapped cells and their response to the application of intense optical forces, able to cause a significant deformation of the cytoskeleton, is of great interest. It is widely recognized that alterations of the cytoskeleton deformability are present in many diseases and their measurement can be used as a reliable marker of the cell status [4,5].
Mechanical properties of single cells can be efficiently tested using a recently developed device based on optical fibers, named optical stretcher (OS) . OS basic idea relies on a double laser beam trap  obtained with two counter-propagating fiber beams [8–10]. Increasing the laser power, the radiation pressure exerted by the two beams over the trapped cell surface leads to an elongation of the cell along the beam axis, providing meaningful information on the cell health. Although the effectiveness of the OS has been widely demonstrated, the typical set-ups, based on assembling optical fibers with glass capillaries or PDMS microchannels, presents some criticality mainly due to the fine and stable alignment required between discrete optical and microfluidic components [6,8].
The lab-on-chip approach integrates microfluidic and optical functions onto a single chip. This integration is a key issue in order to turn biophotonic devices, based on optical forces, into operative tools to be exploited for biological studies. A previous work reported on the realization of a GaAs/AlGaAs chip for the fabrication of integrated traps exploiting a dual beam scheme . The chip, including both laser sources and microfluidic channel, has a quite complex fabrication procedure. Although efficient trapping was obtained, it should be noted that the use of semiconductor integrated lasers could reduce the chip flexibility due to the limited power available, the poor spatial quality of the optical beams and the insurgence of heating effects. In addition, the chip substrate is not transparent to visible light, thus preventing straightforward imaging of trapped cells obtainable through an optical transmission microscope.
Recently, femtosecond lasers have been demonstrated to be valuable tools for micromachining of transparent materials . Differently from standard fabrication technologies this innovative technique, if combined with chemical etching, is able to provide direct writing of both optical waveguides and microfluidic channels , ensuring extreme flexibility and accuracy, together with intrinsic three-dimensional capabilities. The use of femtosecond lasers for micromachining of optofluidic devices has already proved to be successful in several bio-photonic applications [14–18].
In this work we present a monolithic chip, fabricated by femtosecond laser micromachining, which allows performing mechanical analysis on single cells without physical contact and with high reproducibility. The integrated chip is based on a fused silica glass substrate, thus providing high transparency for cell imaging, and represents a significant improvement in terms of stability, robustness and optical damage threshold over existing optical cell stretchers. Optical trapping and manipulation of red blood cells (RBCs) in the optofluidic chip are obtained by means of two counter-propagating beams coming from two integrated optical waveguides orthogonal to the microfluidic channel (Fig. 1 ). By increasing the optical power a considerable stretching of the trapped cell can be observed, allowing one to probe the cell deformability. The delivery of the cell suspension to the trapping region is accomplished by an easy connection of the microchannel to an external fluidic circuit, which guarantees a controlled flow and a high-throughput analysis. A fiber laser source is butt-coupled to the waveguides in the chip, delivering the light required for the trapping and stretching of cells. Since glass absorption in the wavelength range adopted in the experiments (near infrared) is very low, the high powers needed for optical stretching can be easily coupled without heating appreciably the chip. Moreover, the high spatial quality of the trapping beams is guaranteed by the waveguide spatial mode distribution.
The device we propose is user-friendly and reliable as it doesn’t require any critical alignment between discrete optical and fluidic elements. In addition, it allows very stable and reproducible operation, which is a very important asset when quantitative analysis of the cell deformability is required.
2. Femtosecond laser fabrication of the monolithic dual beam optical trap
In this work, for the first time, a complete optofluidic chip is fabricated by a high repetition rate fs-laser, which enables a rather high processing speed. The laser writing fabrication of the device ensures a very accurate transversal alignment between the optical waveguides and a precise positioning with respect to the microchannel. In the following we will describe the different steps that were required for the fabrication of the complete monolithic device.
The geometrical specifications of the device are decided according to the typical dimensions reported in the literature for fiber-based OS [4–9]. The critical parameters are the microchannel diameter, the waveguide mode size and the optical distance between the waveguide end-faces. For the microchannel diameter a value of about 100 µm is chosen, since the capillaries used in the fiber-based OS have an internal dimension of the same order; the distance between the waveguide end-faces should range between 200 µm and 400 µm; the target mode size for the waveguides is set to 3.5 µm radius in order to match the fiber single mode at 1.07 µm wavelength. Indeed, in the trapping and stretching experiments a laser wavelength of λ ≈1 µm is chosen due to the following reasons: i) availability of compact fiber lasers with average power sufficient to achieve trapping and stretching of cells; ii) very low absorption of glass and cells; iii) possibility to filter out the laser light used for trapping, keeping the full spectrum of visible light for the cell imaging.
In the following we will describe the different steps that are required for the fabrication of the complete monolithic device.
2.1 Microfluidic channel fabrication
Femtosecond laser irradiation followed by chemical etching in fused silica substrates is a powerful technique to directly fabricate buried microchannels [19–21]. A recent method, based on irradiating the substrate with beam trajectories more complex than just straight lines allowed demonstrating a very fine control of the microchannel shape and size . This technique is exploited in this work to create large access holes on the side facets of the chip in order to achieve easy connection with external capillary tubes (Fig. 2 ). Fabrication of lateral access holes (instead of top surface holes) guarantees a completely clear path for imaging the trapped cells with a microscope in transmission mode. The access-hole diameter (350 µm in the current work) is designed to exactly match the outer diameter of the capillary tubes. This tailoring is achieved by irradiating multiple coaxial helixes with different radii and with a pitch of 2 µm (an example is represented in Fig. 2(a) – red line). The number of coaxial helixes depends on the desired size of the access hole; for a 350 µm diameter, 3 helixes are written with diameters of 80 µm, 160 µm, and 240 µm, respectively (Fig. 2(b)). The two access holes are connected by a straight line that, once etched, will provide a slowly tapered microchannel with a uniform central portion of 80 µm diameter where the optical trapping is achieved (Fig. 2(c)). The channel walls have a minimum radius of curvature of 40 µm and show the typical surface pattern obtained with this technology  providing an estimated roughness in the 300-500 nm range.
Irradiation is performed by focusing through a 50 × (0.6 NA) objective a frequency-doubled cavity-dumped Yb:KYW laser, providing 230-fs pulses at 600 kHz repetition rate with a pulse energy of 290 nJ at the second harmonic wavelength of 515 nm. The laser polarization is set perpendicular to the microchannel axis, which is placed at a depth of 400 µm with respect to the top surface. With the high-repetition-rate laser an irradiation speed of 1 mm/s is feasible; therefore, although complex structures are irradiated, the processing of the full chip represented in Fig. 2(b) takes about 30 minutes. The chip is then immersed in an ultrasonic bath with 20% of hydrofluoric acid (HF) in water for 4.5 hours to obtain the 3-mm-long buried microchannel (Fig. 2(c)).
2.2 Optical waveguide writing
Femtosecond laser waveguide writing is by now a well assessed technology that has been used to demonstrate several devices [12,23]. Focusing ultrashort laser pulses in the bulk of a transparent material can induce a smooth refractive index increase located at the focal volume. Suitable translation of the sample allows direct writing of an optical waveguide with three-dimensional capabilities. Waveguide writing in the fused silica sample is performed in the same conditions used for the irradiation step in the microchannel fabrication, i.e. focusing through a 50 × objective the frequency-doubled cavity-dumped Yb:KYW laser; however, this time the laser is operated at a repetition rate of 1 MHz since in this regime the fabricated waveguides exhibit lower propagation losses.
Waveguide writing parameters are optimized in order to have the best guiding properties at the operating wavelength of 1 µm. A pulse energy of 100 nJ and a translation speed of 0.5 mm/s allows obtaining single mode waveguides at the operating wavelength with a mode intensity radius at 1/e2 equal to ~4 µm and an ellipticity factor of 1.1. Figure 3(a) shows the experimental near field intensity profile of a guided mode, acquired at a polished end-facet of the optical waveguide. Measured propagation losses at the operating wavelength are equal to 0.9 dB/cm.
2.3 Integration of optical waveguides and microchannel
In the previous Sections we presented the fabrication of the basic components of the optofluidic device and showed that they have specifications in very good agreement with the targeted ones discussed at the beginning of Section 2. In this paragraph the integration of these components will be discussed.
The use of the same laser source and focusing conditions for fabricating both the microchannels and the optical waveguides is a key point that allows performing the whole irradiation process in a single step (Fig. 2(a) reports a sketch of the overall irradiation path, with the waveguides in green). The only variation between the microchannel irradiation and the waveguide writing is the laser repetition rate, which is simply switched electronically by applying a different parameter to the cavity dumper, without affecting the beam direction. In this way a very accurate alignment between the microchannel and the optical waveguides can be achieved with a spatial resolution equal to that of the translation stages, in our case (FiberGlide 3D, Aerotech) better than 100 nm. Two sets of waveguides are fabricated on the two sides of the microchannel, with a separation between the waveguide end-faces of 200 µm and 250 µm (Fig. 3(b)). Each set is composed of 3 waveguides, laterally spaced by 80 µm, that are fabricated at various depths with respect to the axis of the microchannel, i.e. + 5, 0 and −5 µm. In this way different depth positions of the trap are experimentally tested. Moreover, this approach could be exploited to fabricate several parallel traps able to intercept cells flowing at different heights, thus improving the measurement throughput.
In order to appreciate the quality of the trapping beams transmitted through the channel walls, an optofluidic chip is cut along the channel axis and finely polished until one half of the channel is exposed. A water drop and a thin glass slide are then placed on the open channel in order to directly image the beam profile, coming from a waveguide, in the middle of the channel (Fig. 3 (c)), where the cells are typically trapped. This measurement is performed by imaging on a camera (C2400, Hamamatsu) the beam profile corresponding to the plane where the chip cut edges are at focus. It can be noted that, notwithstanding the roughness of the channel walls, the quality of the beam in the channel is very good. The larger size of the beam profile in Fig. 3 (c), with respect to the waveguide mode in Fig. 3 (a), is due to the beam divergence at the waveguide output, which is placed at 60 µm from the channel wall; the decreased ellipticity is a dioptric effect due to the curvature of the channel wall.
The overall fabrication process can thus be summarized in the following steps: i) the femtosecond laser is set to a repetition rate of 600 kHz and the structures for the microchannels are irradiated (typically several structures are fabricated on the same glass substrate); ii) The laser repetition rate is switched to 1 MHz without losing the alignment and the sets of waveguides are written in each device (work is in progress to explore the possibility to perform also the irradiation for the microchannel at 1 MHz, in order to avoid the step of repetition rate switching); iii) the substrate is cut and different chips with 3 mm × 8 mm size are obtained; iv) etching of the microchannels is performed by immersion in the HF solution. Since the irradiation of both microchannels and waveguides is performed before chemical etching, the writing of the waveguides is interrupted 500 µm before the edge of the chip, in order to avoid any etching of the regions corresponding to the waveguides. After the etching the two end-faces are polished in order to expose the waveguide input ends and perform efficient fiber coupling.
2.4 Connection to external fluidic and optical circuits
Once a chip is fabricated, it is connected to external fluidic and optical circuits. Using a set-up composed by an optical microscope and accurate translation stages (NanoMax, Melles Griot) external capillaries are inserted in the access holes (Fig. 4(a) ). Once the capillary is firmly inserted it is glued by a drop of UV-curable resin. The external circuit is essentially made of two butterfly needles glued to the capillaries; the tubes at the other hand of the butterfly needles act as reservoirs. Cell suspension is transported through the trapping region by a controlled microfluidic flow, that is obtained by varying the relative heights of the two reservoirs and can be finely adjusted with a micromanipulator.
Optical fibers are aligned to the waveguides input-facets by means of two translation stages. Butt-coupling is presently used in order to have a flexible set-up, able to test all the waveguides in the chip; however, in a final device the fiber will be permanently pigtailed to the waveguide following the standard procedure developed for photonic devices in telecommunications (typical additional losses ~0.5 dB). From calculations of the mode overlap between the fiber (w0 = 3.3 μm) and the waveguide (w0 = 4.0 μm) we estimate coupling losses of 0.8 dB/facet. Considering the propagation losses given in Section 2.2 and the waveguide length of 4 mm, we calculate additional losses of 0.36 dB. This means an overall loss of 1.3 dB from the fiber to the waveguide output (also including 0.15 dB for Fresnel losses). It should also be considered that the power emitted by the waveguide will suffer additional losses when entering the microchannel due to scattering at the walls.
The chip connected to the capillaries and the fibers is also glued by UV-curable resin on a thin glass slide to increase robustness of the connections but still allowing imaging of the channel content with a high magnification objective (Fig. 4(b)).
3. Optical trapping and stretching experiments
3.1 Experimental set-up
The schematic diagram of the experimental set-up used to demonstrate the effectiveness of the integrated optical stretcher is shown in Fig. 5 . A CW ytterbium fiber laser (YLD-5, IPG Fibertech), with an emitting power up to 5W at 1070 nm, is used as light source. The beam coming from the laser is split in two branches by means of a 50%-50% fiber coupler (FC1). The optical power in each arm is then controlled by variable optical attenuators (VOAs) and monitored using the 1% port of a 99%-1% fiber coupler (FC2a); this enables to finely balance the optical power at the output of the two fibers. In order to optimize the light coupling into the chip-integrated optical waveguides, a second 99%-1% fiber coupler (FC2b) is added in the fiber line: the power coupled into one waveguide, transmitted through the microchannel and collected by the second waveguide, is thus monitored in the opposite branch. All the fiber components are single mode at the working wavelength as well as the spliced bare end-fibers (OF) (Hi-1060, Corning). The VOAs (BB-500, OZ Optics) are specified for operation up to 2 W of optical power, while we verified the FCs (Single mode coupler, OPTO-LINK Corporation) up to 4 W. Given the high optical threshold of the fused silica chip, the current set-up can also be used to stretch cells other than RBCs, where higher power may be needed.
The chip is mounted on an inverted microscope equipped for phase contrast microscopy (TE2000U, Nikon). Phase contrast images of optical trapping and stretching are acquired by a CCD camera (DS-Fi1, Nikon). The pixel size for all the employed magnifications was calibrated with a grating; this allows for absolute distance measurements with a resolution of 0.055 μm/pixel in the case of a 40 × objective.
3.2 Experimental results and discussion
The trapping and stretching capabilities of the chip are tested on RBCs. The cell suspension is prepared by diluting 10 μL of blood in 8 mL of hypotonic solution in which the RBCs acquire a quasi-spherical shape with a radius of ~4 µm. The cell suspension is inserted in the microfluidic circuit described in Section 2.4. For an easy imaging of the flowing cells, the typical value of the cell speed is set in the 10-50 µm/s range. RBCs optical trapping is achieved with an estimated optical power at each waveguide output of about 20 mW. Figure 6 demonstrates the capabilities of the monolithic device to trap cells and move them along the beam axis. In Fig. 6(a) a scheme of the trapping region is reported. Figure 6(b) shows a sequence of a few frames demonstrating how the trapped RBC is stable in its position even if a background flow is present (flowing cells are out of focus since they are travelling at different heights in the microchannel). Moreover, Fig. 6(c) shows the controlled movement of two trapped RBCs along the beam axis obtained by unbalancing the optical forces applied on the two sides of the dual beam trap. The force unbalance is easily achieved by varying the output power of one of the two waveguides, which can be finely tuned by adjusting the corresponding VOA (see Fig. 5).
When a single cell is stably trapped in the microchannel it can be stretched along the trap axis by simultaneously increasing the optical forces applied to the cell by the two counterpropagating beams . Experimentally this is achieved by keeping the VOAs set at the initial value for stable trapping; by raising the emitted power from the laser source a simultaneous and proportional increase of the output power of the two waveguides is achieved. Therefore, the trap is still stable and a progressive stretching of RBC is observed. Trapping and stretching are successfully achieved on several RBCs. Figure 7 shows a sequence of frames demonstrating the optical stretching of a single RBC. The cell, initially trapped at low power (20 mW at each waveguide output), can be elongated up to 25% of its initial size when increasing the waveguide output power to 300 mW. However, in order to achieve such a clearly visible elongation the cell is stretched into its plastic deformation regime. By stretching the cell with lower optical power smaller deformations are observed in the elastic regime, in which the cell returns to its initial shape when the optical power is turned down.
The current device still suffers from two limitations for accurate analysis of cells: curvature and roughness of the microchannel. The lens effect induced by the curvature prevents from a reliable retrieval of the cell contour. However, work is in progress to fabricate rectangular cross-section microchannels by exploiting the spiralling technique employed in this work for the access holes fabrication. On the other hand, the wall roughness (which is not significantly affecting the optical trap quality, as discussed in Section 2) does not allow to achieve optimal quality images of the trapped cells. Also in this case the limitation can be overcome by employing suitable etching conditions that provide surfaces with very low roughness ; however, these conditions provide shorter microchannels and therefore a hybrid etching process is under investigation to create a clear window just in the region of optical trapping.
In this paper we have demonstrated for the first time a monolithic chip for optical trapping and stretching of single cells fabricated in glass by fs-laser micromachining. With this innovative technique both the microchannel and the waveguides, creating the dual-beam trap, have been produced. By using a single irradiation step in the fabrication process, the alignment between the optical and fluidic components has been achieved with extremely high accuracy. This optofluidic chip has been successfully tested using red blood cells flowing through the microchannel. Curvature and roughness of the channel prevent from a quantitative analysis of the cells and work is in progress to overcome these limitations. However, optical trapping of single and multiple cells has been demonstrated with optical powers at the waveguide outputs comparable to those used in fiber-based optical traps. Optical stretching of a single cell has been also observed when increasing the laser power, thus proving the effectiveness of this device as a monolithic OS.
The transparency of the glass substrate in the visible and MIR range guarantees an easy coupling to a microscope platform and it will enable Raman and fluorescence imaging of the trapped cells. In addition, the extreme flexibility and the three-dimensional capabilities of femtosecond laser micromachining will allow empowering the device with several additional functionalities integrated on-chip, e.g. multiple parallel analysis and accurate sorting of the cells under test.
The authors thank V. Degiorgio and G. Cerullo for their advice and support to the research. KCV acknowledges Italian Ministry of University and Research (MUR) for the scholarship grant.
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