TECHNICAL FIELD OF THE INVENTION

The present invention relates to a method of and apparatus for comparing first optical properties of a first fluid with second optical properties of a second fluid. The first and second optical properties may particularly include a first and a second refractive index of the first and second fluids, respectively.

PRIOR ART

The refractive index is a fundamental quantity, intrinsic to the physical and chemical properties of a substance. Measurements of refractive indices are widely used in the pharmaceutical industry, ¹ environmental monitoring, ² adulteration detection, ^3-5 and biosensing. ^6-10 Several technologies for refractive index measurement have been previously described in the literature, and with the push for further miniaturization, many microfluidic technologies have emerged. A brief summary of these methods is outlined in Table 1 .

For many applications, there is a need to measure small changes in refractive index that are easily overwhelmed by nonspecific background signals. For example, measurements of concentration via refractive index require accurate temperature control, and refractive index based biosensors need to be made insensitive to non-specific binding or bulk refractive index changes. Differential measurements with two devices offer a partial remedy. However, the attainable background suppression is often limited due to alignment errors, fabrication tolerances, and other differences between independent sensors. Diffractive optical microdevices open up a path to differential sensing with direct background cancellation by the interference of light waves. One key advantage of these devices is that the signal of interest and the reference are closely integrated and lie on the same beam path. This principle has previously been used for displacement sensing in micromechanics, ^11-13 for chemical sensing, ^14-16 and for biosensing ^17-19 . I n chemical and biological sensing, many powerful device concepts have been proposed based on micropatterning capture molecules or hydrogels into stimulus-responsive phase gratings. ^17,20 Binding of target molecules or swelling of the hydrogel alters the wavefront of a laser passing through the element, and this can be read out via a consequent change of the diffraction pattern in the far-field. Multiplexing possible by this principle if multiple gratings with different orientation are overlaid. ^21,22

Table 1 : Microfluidic refractive index measurement methods

Technology Detection limit Background Size / Volume Ref.

cancelation/self- referencing

Surface Plasmon no-' - no 20-150 μΙ_ 23 24 23 Resonance 2 ^* 10 ^"5 RIU

Young interferometer 1 .8 ^* 10 ^"8 RIU ²⁵ yes, reference arm 6 μΙ_ ²⁵ 25, 26,

27, 28

Microinterferometric 6.9 ^* 10 ^"9 RIU yes, two capillaries: 50 nl_ 29 backscatter detector fringe subtraction

Ring resonator 3.16 ^* 10- 6 _R | U30 Independent reference 200 x 20 μηι / 30, 31,

3.8 ^* 10 ^"8 RIU ³³ structures: multiple 10 μΙ 32, 33

_/Γηίη ³²

5.0 ^* 10 ^"6 RIU ³² sensors on one chip with Resonator cross

beam splitter ³² section: 68 μηι ³¹

Yes, double resonator ³³

Fiber Bragg Grating 2.2 ^* 10 ^"5 RIU no 3.45 μηη per 34, 35 hole

Fabry-Perot Cavity 1 .7 ^* 10 ^"5 RIU ³⁶ no Cavity 36, 37, width=24.5 μηι ³⁶

Photonic crystal: 7 ^* 10 ^"5 RIU Independent reference Hole diameter 43 Nanoscaled structures: Multiple 200 nm, 250 nm

Optofluidic Sensor sensor structures on one deep, 8 holes

Array (NOSA) chip with one waveguide per sensor

Diffraction grating 1 .9 ^* 10 ^"6 I0 as reference 50 μηη thick fluid 44, 45, 46

RIU ^{44 45} layer ⁴⁶

6 ^* 10 ^"7 RIU

Micro edge 3 ^* 10 ^"3 RIU no 20 μηη deep; 47

250 μηη wide

μ-image defocusing 3- 53.7 pixel/RIU self-calibration: 50 μηη wide; 17- 48 pinhole aperture reference fluids included 82 μηη deep

on chip - one image

measurement

Tapered fiber 1 .42 ^* 10 ^"5 RIU ⁴⁹ no Diameter of 9, 49 fiber 200 μηι;

6 mm long ⁹ OBJECT OF THE INVENTION

It is the object of the present invention to provide a method of and apparatus for comparing first optical properties of a first fluid with second optical properties of a second fluid, which allow for a straightforward implementation and nevertheless allow for a precise measurement of, for example, a difference of refractive indices of the first and second fluids both with regard to sign and magnitude.

SOLUTION According to the present invention, the object of the invention is solved by the features of the independent claims.

Additional preferred embodiments according to the invention are to be seen in the dependent claims.

In the claims, the term fluid may refer to a gas or an aerosol. More particular, it refers to a liquid or a liquid based suspension.

The first and second optical properties may particularly include or be determined by a first and a second refractive index of the first and second fluids, respectively. In another embodiment, the first and second optical properties may be determined by absorbance and/or reflectivity and/or turbidity. That the second transparent grating is arranged at a lateral offset of less than 45% of the grating constant with regard to the first transparent grating means that the overall arrangement of the first and second gratings is not symmetric with regard to any of the grating bars. As a result, the grating constant of both the first and second gratings also is the grating constant of their overall arrangement, and each pair of grating bars of the first and second transparent gratings arranged side by side only form a unit cell of the overall arrangement. Due to the asymmetry of this unit cell, the effect of a difference in optical properties between the first and second fluids on the maxima of the diffraction pattern is also asymmetric. Thus, whether the light intensities of the maxima on the left or right hand side of the diffraction pattern increase over the light intensities of the maxima on the other side depends on the sign of the variation of the optical properties. Thus, with the optical properties being determined by the refractive indices of the first and second fluids, the sign of the difference of the refractive indices of the first and second fluids will be decisive. Further, the magnitude of the increase is directly correlated with the magnitude of the variation of the optical properties, i.e., for example, with the difference of the refractive indices of the first and second fluids.

Particularly, the refractive index difference between the first refractive index of the first fluid and the second refractive index of the second fluid can be calculated from a difference of the two light intensities of two intensity maxima of the same order divided by a sum of the two light intensities of two intensity maxima of the same order and by a constant. This constant will be fixed for the respective embodiment of the present invention.

In one embodiment, the first and the second fluids are transparent or clear liquids or gases. In this case the refractive indices of the first and second fluids will determine their optical properties. In another embodiment, the first and the second fluids are two biological fluids containing cells. In this case, the reflectivity and/or absorbance and/or turbidity and/or refractive index of the cell suspension will determine the optical properties of the first and second fluids.

The cells in the biological fluid may be dead or living. With living cells, a temporal development of the cells in the two biological fluids can be monitored by repeatedly directing the coherent light onto the first and second transparent gratings and measuring the two light intensities of the two intensity maxima of the same order.

With regard to the apparatus for comparing first optical properties of a first fluid with second optical properties of a second fluid, all of the parallel fluidic channels in the microfluidic chip will be arranged in one common plane. If the term spacing is used to define the first set of channels or the second set of channels, this term does not refer to the free lateral distances of the respective set of channels but to a repetition or period length including the width of the channels.

That the microfluidic chip has a reflective front face or that that the microfluidic chip being transparent has a reflective back face reflecting the coherent light may be achieved in various ways which include, for example, a reflective coating or interface and a polished surface hit by the coherent light at an angle beyond the critical angle for total reflection.

Often, the fluidic channels in the microfluidic chip will be true nanochannels, i.e. of a depth in the direction perpendicular to their common plane of less than 1 μηη. Nevertheless, if the term nanochannel is used in the following, this is only intended to indicate fluidic channels of a small depth of not more than 10 μηη but not intended to limit these channels to true nanochannels.

DESCRIPTION OF THE DRAWINGS

In the following, the invention is further explained and described with respect to preferred exemplary embodiments illustrated in the drawings.

Fig. 1 is a schematic showing the diffraction grating with nanochannels with an asymmetric arrangement of detection and reference channels and the intensity profile of the reflected diffraction pattern. The reference channels, the detection channels and the solid walls of the transparent plate are depicted in yellow, green and gray, respectively. The channels may be open or closed. Without loss of generality, they are depicted open here, but they may also be closed with a lid made of a transparent material.

Fig. 2 A) is a schematic of the fabrication process for the nanofluidic assymmetric grating device. B) is a SEM top tilted view of open nanochannels and supply vias. C) is a SEM cross sectional view of a 290 nm deep nanochannel after glass wafer bonding to Silicon-on-lnsulator (SOI) wafers. D) is an optical top view image of the nanofluidic grating. The transmission of light through the 4 μηη thick silicon device layer gives the red coloration to the back side feeding channels.

Fig. 3 is A) a plot of the signal S2 as a function of the measurement time. The definition of S _m (where m is an integer) is stated in terms of the mode intensities in Equation 6. A sudden change in S2 is observed when the glycerol solution is introduced into the detection nanochannels. B) is an experimental calibration curve and analytical solution for small Δη as a function of Δη. (Device dimensions w = 3 μηη, I = 4 μηη, P = 18 μηη).

Fig. 4 illustrates an experimental verification of common mode rejection. Phase 1 : Dl water in reference and detection channel, T = 25 °C. Phase 2: 2% glycerol solution in reference channel, T = 25 °C. Phase 3: 2% glycerol solution in reference channel, T = 33 °C. Phase 4: Dl water in reference and detection channel, T = 33 °C. (Device dimensions w = 4 μηη, I = 6 μηη, P = 18 Mm). Fig. 5 shows signal changes {Si) due to sequential sample measu rement (water and 2% glycerol) in addition to measurement of protein accumulation of avidin in detection channels. (Device dimensions w = 4 μηη, I = 6 μηη, P = 18 μηη).

Fig. 6 shows an embodiment of a setup of measurement. The observed diffraction pattern is generated when a collimated laser beam (1=635 nm, waste diameter 360 μηη) impinges on the nanofluid ic grating and is then reflected back onto a mirror and into a digital camera (for example: Thorlabs DCU223M or Andor iXon Ultra or Andor Neo or PCO Edge 4.2 or other detector with at least two detection elements, or pixels.) that records the intensity. The mirror is made out of a glass wafer with a reflecting layer on top which is made of gold here. A slit (3 mm x 10 mm) in the center of the mirror allows the incoming laser beam to pass, but the diffraction pattern is reflected towards the camera. Only the 0th maximum is not reflected as it also passes through the slit in the mirror. This is beneficial for the signal processing as the 0th maximum is very bright compared to the other maxima and might interfere with the other maxima without containing any information. The signal intensity of each maximum is measured by integrating the grey values over an area large enough to include the entire signal (AOI-area of interest).

Fig. 7 shows how fluid flow in the nanofluidic grating is provided by pressure control. Fluid is pumped continuously through the microchannels by the selected pressure values. During the wash step, the same fluid is pushed from the main microchannel (left) into the nanochannels. During the measure step, different reference (pale green) and sample (blue) fluids are pushed from the respective microchannels into the nanochannels. The mentioned pressure values in Fig. 7 have been used for all measurements with water and glycerol solutions. In case of fluids with higher viscosities, like piranha solutions for cleaning, the pressure values were increased accordingly to ensure fluid flow. Fig. 8 shows an intensity distribution of light propagating through a diffractive element which is formed by a linear microchannel array; and

Fig. 9 shows plots of various values relevant for the theoretical background of the present invention. DETAILED DESCRIPTION

Measuring small changes in refractive index can provide both sensitive and contactless information on molecule concentration or process conditions for a wide range of applications. However, refractive index measurements are easily perturbed by non-specific background signals, such as temperature changes or non-specific binding. Here, we present an optofluidic device for measuring refractive index with direct background subtraction within a single measurement. The device is comprised of two interdigitated arrays of nanofluidic channels designed to form an optical grating. Optical path differences between the two sets of channels can be measured directly via an intensity ratio within the diffraction pattern that forms when the grating is illuminated by a collimated laser beam. Our results show that no calibration or biasing is required if the unit cell of the grating is designed with an appropriate built-in asymmetry. In proof-of-concept experiments we attained a noise level equivalent to ~10 ^"5 refractive index units (30 Hz sa m pl in g rate, 4 m i n measu rement interval). Fu rthermore, we show that the accumulation of biomolecules on the surface of the nanochannels can be measured in real- time. Because of its simplicity and robustness, we expect that this inherently differential measurement concept will find many applications in ultra-low volume analytical systems, biosensors, and portable devices.

More particular, we describe a new diffractive optofluidic device for measuring small differences in refractive index between two fluids that are guided through nanofluidic channels. The device consists of two sets of nanochannels, which are arrayed to form an interdigitated grating with an asymmetric unit cell. We show for the first time that the asymmetry enables linear optical differencing in diffractive sensors even when the optical layer thicknesses (here the nanochannel depths) are fixed and cannot be adjusted to achieve phase quadrature.

Using this device, it is possible to measure small differences in the bulk refractive index of two solutions in the presence of large common-mode fluctuations. Alternatively, the device can measure optical path differences caused by different surface-adsorbed layers. This effect is significant due to the small height of our nanofluidic channels. We envision that the inherently differential measurement and the efficient surface-directed transport of molecules in the nanofluidic channels render this concept interesting for label-free biosensing applications. Optofluidic grating design and theory A schematic representation of the nanofluidic diffraction grating device where reference and detection nanochannels are placed in an asymmetric interdigitating arrangement is shown in Fig. 1 . The asymmetric unit cell of the grating results in the formation of a diffraction pattern which is not mirror-symmetric (l- _m ≠ l _m ). Importantly, this asymmetric distribution is produced only when the refractive indices in the reference (yellow) and in the detection channels (green) are different, Δη = n ¾f - noef≠ 0. This allows changes in the refractive index within the detection nanochannels to be detected by measuring the difference in intensity \ _m = U _m - l _m .

For a correct signal processing, we calculated (see below: Supplement on Theory) the dependence of the intensity distribution of the diffraction pattern produced by the asymmetric grating for small Δη.

Fabrication

The nanofluidic grating sensor was fabricated using standard microfabrication techniques as shown in Fig. 2A. Silicon-on-insulator (SOI) wafers were initially dry thermally oxidized to a depth of 290 nm. Nanochannels were then formed on the S1O2 top layer using photolithography and buffered hydrofluoric acid (BHF) wet etching. The silicon device layer of the SOI wafer acted as an etch stop, resulting in a nanochannel depth (h) equal to the thickness of the S1O2 layer.

For the lateral dimensions of the nanochannels, we selected two combinations: (1 ) w = 3 μηη, I = 7 μηη, P = 18 μηη and (2) w = 4 μηη, I = 6 μηη, P = 18 μηη. These dimensions were chosen to obtain high sensitivity to bulk refractive index changes and to surface-adsorption while maintaining safe tolerances during fabrication. The grating layout consisted of 25 reference and detection nanochannels, respectively, for a total of 50 channels per grating, and each channel was 320 μηη in length. Vias (3 μηη in diameter) on each end of the nannochannels were opened by deep reactive ion etching (DRIE) through the silicon device layer of the SOI followed by BHF to open the buried oxide (Fig. 2B).

Subsequently, a 210 μηη thick Borofloat 33 wafer was bonded to the top side of the silicon wafer to seal the nanofluidic grating. The silicon wafer was then ground to a thickness of 50 μηη and polished. After the silicon was thinned down, microfluidic channels were etched from the back side by DRIE to connect to the vias previously fabricated from the top side. After bonding, a thin residual oxide diaphragm that sometimes remained in the vias was cleared by a vapor-phase HF etch. Thinning the silicon served a dual role in the above process. First, the depth and aspect ratio of the DRI E step was significantly reduced; thus, the etching could be stopped uniformly and with minimal footing on the buried oxide layer of the SOI. Second, the volume of the microfluidic channels connecting to the nanochannels could be kept small in this way. Finally, a 700 μηη thick Borofloat 33 wafer was bonded onto the back side of the wafer to ensure the robustness of the fabricated devices. Before bonding, through-holes (800 μηη in diameter) were opened on the back side Borofloat 33 wafer by femtosecond laser ablation to allow fluid delivery into the nanofluidic system.

Optical and fluidic setup

The diffraction pattern that is generated when a collimated laser beam (λ =635 nm, waist diameter 360 μηη) impinges on the nanofluidic grating is reflected back onto a mirror and into a CCD camera (Thorlabs DCU223M or Andor iXon Ultra). The signal intensity of each maximum (l±m) is measured by integrating the gray values over an area large enough to include all of the signal. A pressure-driven fluidic system was used to introduce fluids into all channels. The reference and detection nanochannels can be supplied with different fluids through separate microfluidic supply channels (Fig. 1 and Fig. 2B), enabling selective functionalization and avoiding contamination between the channels. The fluids are guided through the nanochannels by controlling the pressures on either side.

For each measurement, a baseline was determined by flushing all channels with the reference solution (water or PBS for these studies) for five minutes. At the five minute interval, the pressure difference was switched to introduce the sample solution into only the detection channel and reference solution was maintained in the reference channel . Following the measurement, all channels in the grating were again filled with the reference solution. To avoid clogging of the nanochannels, all solutions were passed through a 300 kDa cutoff filter before use.

Results and Discussions

Device Characterization

To validate our theoretical model, we conducted calibration measurements using glycerol solutions in Dl water in the concentration range of 1-30% (w/w). The refractive index of glycerol depends strongly on the solution concentration. ⁵¹ All glycerol solutions were also characterized with a pocket refractometer (Atago USA, Inc.) prior to analysis with our new method. For each measurement, all channels were initially filled with Dl water, and the resulting signal was taken as a baseline. When the glycerol solution was then introduced into the detection channels at around 4 minutes, the difference in refractive index between detection and reference channels resulted in a change in the signal S2 (Fig. 3A). The same procedure was then repeated in triplicate for every concentration of glycerol , thus demonstrating the repeatability of the measurements and stability of the system. I n all measurements the same baseline was recovered when water was again introduced into the detection channels.

Values of S2 are plotted as a function of the independently measured Δη in Fig. 3B. For comparison, Fig. 3B also shows the theoretical prediction calculated according to equation 7. The slope s = 3 S23 Δη corresponds to the sensitivity of our device. Based on the geometry and the wavelength we find s = 2.408 RIU ^-1 . I mportantly, the linear response of the sensor corresponds accurately with the analytical solution for the entire range of Δη measured. Linear regression on the data shown in Fig. 3B yields a slope s = 2.416 ±0.01926 RIU ^"1 (correlation coefficient of R = 0.99954). This is in good agreement with the analytical prediction.

The noise floor is dominated by low-frequency fluctuations. Over the first four minutes of measurement, the refractive index equivalent of the standard deviation in S2 is σ(Δη) = 1 .3 ^χ 10 ^"5 RIU. For comparison with other methods that are summarized in Table 1 , our limit of detection taken as three standard deviations above the noise corresponds to ~ 4 ^χ 10 ^~5 RIU.

Common mode rejection analysis

The accuracy of refractive index measurements can be significantly affected by temperature fluctuations. Here, we show that the differential design of our sensor greatly reduces the influence of thermal drift on the measurement result. Fig. 4 shows an experiment where common refractive index changes due to temperature changes are efficiently suppressed. The experiment is divided into four phases. In the first phase, the temperature of the device was set to 25 °C and Dl water was introduced into both channel types. I n the second phase, a 2% glycerol solution was introduced into only the detection nanochannels. This induced a decrease in both intensities I-2 (black) and I+2 (orange). A Δη = 0.0025 was ca lculated starting from the measured S2 using the experimentally determined sensitivity of 2.416/ARIU. In the third phase, the temperature of the device was increased by 8 °C inducing a change in the refractive index in both reference and detection channels. Although this can be noted as an increase in the individual mode intensities I+2 and I-2, the effect of the temperature change is significantly suppressed in the measured Δη. A small decrease of Δη of about 2 ^χ 1 0 ^"4 RIU was measured in the third phase, around an order of magnitude smaller than the change induced in the refractive index of water by the 8 °C increase of the temperature. ⁵² In the fourth phase, water was introduced again to the reference channel, and the measured Δη subsequently decreased back to values close to zero, despite the temperature remaining 8 °C higher than the initial starting temperature. These results show that common-mode changes in refractive index are suppressed by at least one order of magnitude.

Protein detection

To test whether the nanofluidic grating sensor can be used to detect changes in refractive index due to surface binding, we immobilized avidin to the channel walls (Figure 5). Prior to the experiment all channels had been activated by flowing Piranha solution (1 :3 HbC^HbSCU) through the system, leaving the channel walls with a thin surface oxide layer. Avidin forms positive clusters that readily attach to the negatively charged oxide layer through electrostatic interactions. ⁵³

Phosphate buffered saline (1 x PBS) was used as the reference and wash solution. After 5 minutes of acquiring a baseline, the solution in the detection channel was exchanged to water in order to verify that the system was correctly responding to bulk solution changes. After another 5 minutes, PBS was again flowed through all channels and the baseline signal was recovered. To confirm the ability of the device to sequentially measure changes in refractive index of multiple solutions, 2% glycerol was then introduced. As expected, the signal increased. Next, PBS was again introduced to all channels and the baseline signal was recovered.

Avidin adsorption was measured by flowing a solution of avidin in PBS through the detection channel for 5 minutes and subsequently rinsing with PBS for 10 minutes to remove any loosely bound protein. It is important to note that after the 10 minute rinse, the baseline signal was not recovered, suggesting the avidin molecules adhered tightly to the channel wall. We expect that our nanofluidic grating sensor could be functionalized with different affinity reagents for the label-free detection of specific biomolecules. For sensitive detection, the large surface-to-volume ratio of these channels should provide distinct advantages, as molecules would have a high capture probability on their passage through the channel. At the same time, the large number of parallel nanochannels provides a large effective capture area, so that rare molecules could be efficiently concentrated in the sensing area.

Supplement on Theory

I n this section, we calculate the intensity distribution of light propagating through a diffractive element which is formed by a linear microchannel array, as shown in Fig. 8A. For modeling purposes, we will assume that the microchannel array is thin and forms an ideal phase grating wh ich is illu minated from the backside. The reflective grati ng used i n ou r experi ments is accommodated mathematically by setting the model path length d equal to two times the depth of the real channels.

In the following, our aim is to calculate the difference in refractive index between the fluids A and B by analyzing the intensity distribution of light at a d istance D from the grating. For sufficiently large D, the Fraunhofer approximation can be used to relate the complex-valued entrance amplitude f{u, v) to the exit amplitude g{x,y) at distance D from the grating. When a Gaussian laser beam of power P and waist radius coo is used as the source, the entrance amplitude is given by

u ² +v ²

f(u, v) = f ₀ e (1 )

where f ₀ = and φ (u) is a periodic function (period L) describing the phase shift induced

by the grating. I n the simple model depicted in Fig. 8, <p (u) is piecewise constant with values <Pi = k d rii , where n, (/ ^' = 1 , 2, 3) is the refractive index of either fluid A (i = 1), fluid B (i = 2), or the wall material (here S1O2, i = 3); d is equal to the depth of the channels (equal for A and B)

In the Fraunhofer approximation, the complex amplitude in the far field is given by

-ikD

(2) with the Fourier transform

T{f}{k _xi k _y ) = f(u, v) _e k _x u+k _yV ) _{du dv}

The coefficients are given by cos I -irm (4)

where S _m denotes the Kronecker Delta, Αφ = φ — φ ₂ , and φ ₃ = (φ-^— φ ₃ ) + (φ ₂ — <ρ ₃ )- When the diffraction orders are well separated, i.e. ω ₀ « an intensity distribution similar to the one illustrated in Fig. 8 will be obtained. Neglecting the overlap between different orders, this can be approximated as which represents peaks at positions x _m =—D with peak intensities I _m = g \a _m and g ₀ = ^ ^■ J2P πω . Note that the numerical value of \a _m \ ² directly equals the fractional laser power deflected into mode number m.

As primary signal S _m we consider the ratio between the difference I _m

Im ^'■ = m + I- f° ^r mode pairs ±m≠ 0:

^ sin ² ^- cos ² ( -nrn

Note that Equation 6 is only meaningful when cos -nrnj≠ 0, as otherwise no light is diffracted into the modes of order m, i.e. Im

because then the diffraction pattern is perfectly symmetric. With this in mind, I can be chosen more or less freely within the range w < I < L - w, with only a few values prohibited by manufacturing constraints.

In order to select suitable values for I and w, we analyze how these parameters affect the resolution with which small differences in refractive index between the channel sets A and B can be detected.

We define the signal based on the inverse power series expansion of Equation 6 as

^ ^:= -^ ^cot (z™)!' ⁽⁷⁾ where hat symbols are used to designate measured values. Associated with the intensity measurements is an error o _m = Std{/ _m }, which comprises all noise sources, including shot noise, laser power fluctuations, detector noise, speckle, stray light, and uncompensated offset drift. For « I^, we can approximate σ _} - _/} + ¾ ≡_ By inserting this on the left hand side of

Equation 6 and solving to first order in An we obtain the noise equivalent refractive index difference

The signal-to-noise ratio is maximized by choosing the channel width w and channel separation I such that and m - ^l each is an odd multiple of Here the range is obviously limited to w < ^ and w≤ l≤L - w to ensure that both channels fit with in one grating period without overlapping each other. Regarding the channel depth, signal-to-noise maxima are attained when d = (2/ + 1)—— -, for any integer j≥ 0. This will ensure that φ _∑ is an odd multiple of 2π.

\n— n ₃ 1

What is attractive about the first order approximation given by Equation 7 is that only an intensity ratio is required. This makes it possible to measure the refractive index difference between fluids A and B robustly without knowing any of the absolute refractive index values in the system. Unfortunately, however, only the linear term can be used in this manner. If we were to continue the inverse expansion of Equation 6, the sum of optical path lengths, φ ₃ , would also play a role. While it is possible in theory to measure this value using the absolute intensities, doing so with precision significantly increases the complexity of the method. How limiting is our restriction to the first order inverse in Equation 7? There are two types of inaccuracies to consider. The first, and perhaps most obvious one is due to deviations from linearity when An becomes large. Fig. 9 shows the resulting error e := £{Δη} - Δη. (10)

This error depends strongly on the average refractive index n = ⁿⁱ +" ² . Any device design is optimized for only one value of n according to the condition for φ ₃ given above. The second important type of systematic error is the caused by an imperfect suppression of common mode variations, i.e. refractive index modulations that are identical in channels A and B. A useful measure of this type of error is the slope ^ ^Δη /^ _η ' ^w h' ^cn would be exactly zero in an ideal difference measurement. Although such perfect common-mode suppression is achieved only at the design value n = n ₀ , several orders of magnitude are still possible over a wide dynamic range, as shown in Fig. 9.

Finally, we discuss the dynamic range of n over which meaningful measurements can be conducted. For this purpose, we assume that An « n, which is the regime where our differential method provides the greatest advantage over conventional single-ended measurements. The widest dynamic range is achieved for d chosen as the lowest order, i.e. j = 0. In this case the n ₃ -n

range between adjacent minima in sensitivity is defined by the condition 0 < ≤ 2. For

\n ₃ -n ₀ \

example, a device designed for water (n ₀ = 1.33) with thermal silicon dioxide as material between the channels A and B (n ₃ = 1.46) would be usable over the range 1.2 < n < 1.46.

Conclusions

Here we report the development of a new nanofluidic device for simple and robust differential refractive index measurements in an ultra-low volume. Optical differencing is performed directly by an interferometer with a common path for the sensing and the reference arm. This is enabled by guiding the sample and reference fluids through two sets of parallel nanochannels arranged to form an interdigitated optical reflection grating. A key innovation in our design is the asymmetric arrangement of the unit cell of the grating, which shifts the operating point of the interferometer to allow sensitive measurements even when the refractive index of the two solutions is nearly matched. The differential design of our system allows measurements that are inherently compensated for common mode variations. Here, we have shown that common mode refractive index changes due to temperature fluctuations can be suppressed by at least a factor of 10 in the differential signal. The noise floor of the device in our current setup is limited by low-frequency fluctuations to 1 .3 x 10 ^~5 RIU (standard deviation over 4 minutes). We expect that this can still be significantly improved by optimization of channel alignment and external measurement components.

The separate fluidic addressing of reference and detection channels allows a specific immobilization of molecules in just the detection channels. The capability of quantifying thin adsorbed protein layers in combination with the above described common mode rejection provides the potential for detecting biomolecules label-free with a uniquely simple, robust, and inherently differential sensor. Future applications as a point-of-care device are therefore very attractive.

Further advantageous developments of the invention result from the claims, the description and the drawings. The advantages of features and of combinations of a plurality of features mentioned at the beginning of the description only serve as examples and may be used alternatively or cumulatively without the necessity of embodiments according to the invention having to obtain these advantages. Without changing the scope of protection as defined by the enclosed claims, the following applies with respect to the disclosure of the original application and the patent: further features may be taken from the drawings, in particular from the illustrated designs and the dimensions of a plurality of components with respect to one another as well as from their relative arrangement and their operative connection. The combination of features of different embodiments of the invention or of features of different claims independent of the chosen references of the claims is also possible, and it is motivated herewith. This also relates to features which are illustrated in separate drawings, or which are mentioned when describing them. These features may also be combined with features of different claims. Furthermore, it is possible that further embodiments of the invention do not have the features mentioned in the claims.

The number of the features mentioned in the claims and in the description is to be understood to cover this exact number and a greater number than the mentioned number without having to explicitly use the adverb "at least". For example, if an element is mentioned , this is to be understood such that there is exactly one element or there are two elements or more elements. Additional features may be added to the features defined in the claims, or these features may be the only features of the respective method or apparatus.

The reference signs contained in the claims are not limiting the extent of the matter protected by the claims. Their sole function is to make the claims easier to understand.

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