Thermophoresis Controlled Size-Dependent DNA Translocation

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Thermophoresis Controlled Size-Dependent DNA Translocation through an Array of Nanopores Miao Zhang, Chonmanart Ngampeerapong, David Redin, Afshin Ahmadian, Ilya Sychugov, and Jan Linnros ACS Nano, Just Accepted Manuscript • DOI: 10.1021/acsnano.8b00961 • Publication Date (Web): 12 Apr 2018 Downloaded from http://pubs.acs.org on April 12, 2018

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Thermophoresis Controlled Size-Dependent DNA Translocation through an Array of Nanopores Miao Zhang, *,† Chonmanart Ngampeerapong,† David Redin, ‡ Afshin Ahmadian, ‡ Ilya Sychugov, † and Jan Linnros*,† †

Department of Applied Physics, KTH Royal Institute of Technology, Electrum 229, 164 40

Kista, Sweden, ‡ School of Biotechnology, Division of Gene Technology, Science for Life Laboratory, KTH Royal Institute of Technology, SE-171 65, Solna, Sweden. *Address correspondence to [email protected], [email protected].

ABSTRACT: Large arrays of nanopores can be used for high throughput biomolecule translocation with applications towards size discrimination and sorting at the single-molecule level. In this paper, we propose to discriminate DNA length by the capture rate of the molecules to an array of relatively large nanopores (50 – 130 nm) by introducing a thermal gradient in front of the pores balancing the force from an external electric field. Nanopore arrays defined by photolithography were batch processed using standard silicon technology in combination with electrochemical etching. Parallel translocation of single, fluorophore labelled dsDNA strands is recorded by imaging the array with a fast CMOS camera. The experimental data shows that the capture rates of dsDNA molecules decreases with increasing DNA length due to the thermophoretic effect of the molecules. It is shown that the translocation can be completely

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turned off for the longer molecule using an appropriate bias, thus, allowing a size discrimination of the DNA translocation through the nanopores. A derived analytical model correctly predicts the observed capture rate. Our results demonstrate that by combining a thermal and a potential gradient at the nanopores, such large nanopore arrays can potentially be used as a low cost, high throughput platform for molecule sensing and sorting.

KEYWORDS: silicon, nanopore, array, electrochemical etching, thermophoresis, capture rate, sorting

Over the past two decades, the use of a nanopore to translocate biomolecules such as DNA or proteins has progressed tremendously and further thrived into a broad range of applications in single-molecule sensing.1–6 Characterizations including DNA length,7 single-nucleotide polymorphism,8 protein size and conformation9 are achieved by measuring the ionic current blockade when molecules pass through a nanometer-sized pore. The sensing mainly relies on a statistical analysis of the event duration and the level of blockade of the ionic current during the translocation, which can be used even for DNA sequencing.10 As a consequence, this requires the pore size to be sufficiently small to be comparable to the size of the target molecule.11–13 For single-stranded DNA, used for DNA sequencing, a pore diameter approaching 2 nm is needed while for other molecules such as proteins, pore sizes should be scaled appropriately - from several to tens of nanometers. Currently, solid-state nanopores with even sub-nm precision are routinely fabricated one by one by TEM drilling,14 by Helium ion microscopy15 or by dielectric breakdown.16 For arrays of nanopores, however, device throughput is limited by the timeconsuming serial fabrication approach. Although batch processing of large arrays of pores have


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been demonstrated at wafer scale by using cutting-edge electron beam lithography followed by reactive ion etching (RIE),17 or by low cost electrochemical etching,18 the pore sizes are still too large to result in a sufficiently large blockade or slow molecule translocation velocity that allow an information-rich signal analysis. Moreover, when scaling up the electrical sensing from a single pore to a large array, individual sensing units and electric isolation on each single pore leads to a complex structure and a bulky device. To avoid such complexity and to have a highly compact device, parallel optical sensing seems to provide an alternative route. The fast molecule translocation process (~µs), however, often requires an avalanche photodiode (APD) to properly resolve photoluminescence (PL) events temporally, but can only be used on a single pore. On the other hand, using current imaging technology, the frame rate (~kHz) is too slow to reveal detailed translocation processes. Therefore, a true high-throughput sensing platform that can be realized by large-scale fabrication processing is still highly challenging. In this article, we propose a scheme to distinguish DNA length by their capture rates on an array of relatively large nanopores (50 nm -130 nm) using a local thermal gradient. This provides discrimination according to the molecular size or mass. In diffusion limited regime of electrophoresis of DNA through a nanopore, the capture rate does not depend on the diffusion coefficient of DNA ( D ), but rather its electrophoretic mobility ( µ ).19,20 This is because that the diffusion coefficient is cancelled by the capture radius of the nanopore that scales with D-1. In this case, the DNA capture rate is independent of DNA length, due to the fact that the electrophoretic mobility of DNA in a free solution does not depend on the DNA length.21,22 For the same reason, in electrophoretic DNA or polypeptide separation, a sieving material, such as a gel, is needed to physically hinder molecule migration according to their length.23,24 This idea has been previously brought to nanopores by adding a nanofiber-mesh in front of the pore to


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slow down the DNA translocation.25 Here, we instead propose the use of a thermal gradient in front of the pores to achieve a size-dependent translocation of DNA molecules in a physiological solution. Molecules are known to migrate along a thermal gradient in liquids via thermophoresis, also referred to as the Soret effect.26–29 This leads to a depletion or an accumulation of molecules locally, characterized by the Soret coefficient, ST

= DT D , in which DT denotes the thermal

diffusion coefficient. Extensive research efforts reveal that

ST depends on multiple variables,

including particle size, Debye length, salt type, temperature, etc.27–31 Most bio-molecules in physiological solutions (e.g. NaCl, KCl) above room temperature, possess a positive

ST value,29

meaning a depletion of molecules from the hot towards the cold region. Therefore, by turning a nanopore into a hot spot, molecules are subjected to a thermal depletion according to their sizedependent ST coefficient, leading to a thermally controlled molecule capture rate into the pore under an external electrical field. In contrast to the short molecule translocation duration, the arrival of molecules is imaged in a relatively large detection volume defined by the focus of the epifluorescence microscope. Therefore, events can be well distinguished by modern CMOS cameras, probing an array of nanopores at the same time. Moreover, since this scheme relaxes the requirement of pore size, it can be easily scaled up to a large array, ensuring a high throughput of single molecule sensing. For a proof of concept, we first mass fabricated nanopore arrays by photolithography on silicon membranes at wafer scale processing followed by electrochemical etching of the silicon membranes on each chip. 30 000 pores were obtained on each chip with 900 pores on each membrane having an initial pore entrance diameter of about 20 nm. The translocations of fluorophore-labelled dsDNA with different lengths are measured optically with a fast CMOS


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camera under 514 nm laser excitation. Here, the thermal gradient is simply induced by the absorption of the very same laser beam in the ~ 2 µm thick silicon membrane at the excitation spot but may equally well be provided by heating of the membrane by other means. Our results show that the capture rate decreases with increasing DNA length. We also observe that compared to shorter DNA, a higher threshold bias is needed for the longer molecules to overcome the thermal depletion to reach the nanopores thus allowing a discrimination of DNA length. In addition, an analytical model is derived to predict the capture rate by adding a thermophoretic component to the current equation. Our results demonstrate that such large nanopore arrays combining a thermal and a potential gradient can potentially be used as a platform for highthroughput biomolecule sensing and sorting.

Figure 1. SEM images of (a) nanopore arrays in a bird’s eye view and (b) from backside of the membrane. Inset image in (b) is a zoom-in image of a single pore. (c) A cross-sectional TEM image of a pore with a diameter of ~40 nm and depth of 1.1 µm. The scale bar in (c) denotes 50 nm. Formation of small branches along the pore is a side effect of electrochemical etching in the breakdown regime.

RESULTS AND DISCUSSION Fabrication of the silicon nanopore array and the fluidic cell setup. The nanopore arrays used in this work are formed by electrochemical etching of free-standing silicon membranes fabricated on a SOI wafer (details in Materials and Methods) with a relatively thick device layer


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resulting in high-aspect ratio pores. Conventional photolithography is used to define arrays of etching pits on membranes at wafer scale processing. The following pore formation step is performed by electrochemical etching in the breakdown regime: Because of the high curvature at the bottom of the etching pit, the electric field is sufficiently high to enable avalanche breakdown,32 thereby initiating pore formation (Figure 1a). Backside SEM imaging shows that the etch-through success-rate of the silicon membrane approaches 100% (Figure 1b). The average pore diameter measured at the backside entrance maintains 18 ± 4 nm [histogram shown in figure S1 in the Supporting Information (SI)]. With the current pore depth of 1.1 µm, we carefully examined the cross-sectional profile of one pore. Figure 1c shows that as the pore grows deeper, the diameter quickly increases to ~40 nm and keeps this value through the main body of the pore. Note that there are some spiking pores perpendicular to the main pore, which is commonly seen as a side-effect of etching in the breakdown regime (pores grow in (100)directions). Nevertheless, spiking pores come shortly to dead ends in silicon, thus we assume they do not affect the DNA translocation significantly. As commonly observed in SiN and SiO2 nanopores, the diameter of the pore enlarges once it is immersed in a buffer solution. The above mentioned silicon nanopores are quite inert in buffer solution; however, we do observe that with potential applied across the membrane and under continuous laser illumination, the pore diameter enlarges at a rate of ~2 nm/ hour under normal experimental conditions (shown in Figure S2). We attribute this to the dissolution of silicon into silicic acid in water, which has been reported previously.

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Note that these very gradual changes do not affect translocation

measurement results, because each measurement takes only a couple of minutes and we only compare data obtained on pores with the same initial diameters.


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Figure 2. Schematic of the optical and the electrical system for DNA translocation measurements. A fluidic cell holding a membrane is placed on an inverted wide-field microscope for optical detection of fluorophore-labelled DNA. A four-probe electrical system is used to apply the desired bias across the membrane by injecting current via the working and the counter Pt electrodes (WE, CE). The cross-membrane-bias is monitored by a pair of Ag/AgCl electrodes as sensing and reference electrodes (SE, RE). Left: A time-integrated image of a recorded video sequence shows fluorescent signals at each pore position. A PL time trace of one pore is shown below (1 kHz frame rate). DNA translocation events were recorded as the peaks of PL intensity. After digitizing, using a fixed threshold level, event duration (ton) and event interval time (toff) are extracted.

To further verify the pore size on a large array, we characterized the conductance of a single array of pores in a fluidic cell in 1 M KCl (details in Materials and Methods). Because a large number of pores are now in parallel, the total resistance of the pore array drops to the same range as the contact resistance of the electrodes (tens of kΩ). In order to add desired bias across the membrane, a four-probe system controlled by a Potentiostat is used, as shown in Figure 2. While a high current is maintained through the working and the counter Pt electrode, bias across the membrane is sensed accurately between a pair of Ag/AgCl electrodes that are inserted in the cisand trans-chambers respectively. The conductance measurements are done on arrays of pores with different diameters in dark. The average pore diameter as the only variable is then extracted by using the simple model of pore conductance.34 It matches well with the average diameter


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obtained by SEM imaging without any fitting parameters for all sets of different diameter samples (see in section 1 in SI). This result confirms that a ~100% etch-through-rate and a narrow size distribution of ~ 900 pores on a membrane are achieved by electrochemical etching. Optical and electrical DNA detection. To take advantage of an array of pores, we have chosen to perform DNA sensing optically by epifluorescence microscopy in an imaging mode. The schematic of our measuring system is shown in Figure 2. Double-stranded (ds) DNA of length from 225 bp to 3360 bp, each labeled with double fluorophores (ATTO 532) were used in the experiments. A laser beam of wavelength 514 nm, matching the absorption band of the fluorophore, is focused on the front surface of the membrane through the objective lens, as the source for fluorescence excitation, as well as for heating. The full-width-half-maximum of the laser spot is about 7 µm, thus 5-by-5 pores are illuminated simultaneously. When fluorophorelabelled DNA molecules arrive at the pores, they are excited through the depth of field (DOF), which is ~ 1 µm for a ×63 objective lens. The entries of fluorophore-labelled molecules into the pores are detected by bursts of photons, which are then collected by a CMOS camera at 1 kHz frame rate to capture the molecular dynamics at the pore vicinity. Since the optical detection is on the cis side of the fluidic cell, we image DNA molecules migrating against the thermal gradient when entering the pores. Intermittent fluorescent signals are observed on each single pore as can be seen in the time-integrated PL image, summing 50K frames of a recorded sequence in Figure 2 (left). Below, an example PL intensity time trace of one pore depicts the translocation duration time (ton) and the event interval time (toff). To verify successful translocation, we also examined the conductance of the pore array. Indeed, decreases up to 60 % of the pore array conductance were observed when “flashing events” (intermittent fluorescent signals) appeared at pore positions (see in SI, section2). Note that the conductance measurement


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here only reflects the ensemble occupation of the pores by DNA molecules in the whole array. DNA translocation on single pores are determined by the fluorescence signals optically. The whole array of pores was then examined optically in the DNA translocation experiments by laterally moving the position of the fluidic cell with micro-motors. Indeed, ~ 100% of the pores allow DNA to translocate. DNA Translocation data in this paper were collected on arrays of nanopores with an average diameter of ~90 nm and ~130 nm. A video of recorded PL signals is included in the SI.

Figure 3. A cross-sectional view of the temperature profile through the detection cell by FEM simulation. Dashed fine white lines indicate the membrane. Excitation was at 60 mW over a ~16 µm diameter (3σ) area.

Light-induced thermal gradient. The thermal gradient is introduced to the nanopore array by the same laser that excites the fluorophores. Although the buffer is almost transparent to the visible light, silicon has a strong absorption at this laser wavelength (photon energy 2.41 eV) due to a relative small energy bandgap (1.11 eV at 300 K). For the wavelength of 514 nm, the intensity of light propagating in silicon drops to 1/e at ~ 1 µm from the surface.35 Therefore, for a 2.3 µm thick silicon membrane, 90 % of the laser power is absorbed within the membrane. This energy converts to heat, raising the temperature of the surrounding buffer. To evaluate the temperature profile around the membrane, we performed simulations by the finite element method using Comsol Multiphysics with a continuous or pulsed laser with powers ranging from 10 mW to 60 mW (details in Materials and Methods). As shown in Figure 3, a strong thermal


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gradient is created in the buffer solution with the maximal temperature at the center of the heated spot reaching 83 oC at 60 mW. The vertical thermal gradient in the liquid at the vicinity of the membrane is about 10 oC/µm at 60 mW (temperature profiles shown in Figure S5). Despite of a much-localized heating within the laser spot, the average temperature across the whole membrane surface also increases by 0.2 oC/ mW. As temperature increases in a liquid system, the solution viscosity changes accordingly. 36 This fundamentally affects all mass transport processes in the solution, including ion transport, and also the DNA translocation. To verify the simulated light-induced local heating on the membrane, we first examined the pore conductance of the nanopore array under laser illumination with powers ranging from 10 mW to 60 mW. Figure 4a shows a significant change in conductance of the pore array (∆G), linearly dependent on the laser power. Although the number of illuminated pores is only ~3 % of the total number of pores that contribute to the measured conductance, the simulation suggests that because of the heat transfer, the temperature increase occurs all over the membrane. The conductivity of the buffer increases in a linear fashion with temperature up to 90 oC.10,11 Therefore we can estimate the temperature-induced conductance increase on the pore array by using the average temperature of the laser illuminated surface. However, in that case the measured conductance increase is only ~60 % of the predicted value using a buffer conductivity of 1 M KCl. We attribute this to the thermal depletion of ions by the thermal gradient. K+ and Cl- have positive Soret coefficients,30 meaning that both cations and anions migrate away from the heated membrane, resulting in a lower concentration of ions in the pores. Similar results have been reported in a simulation study.39 In fact, the pore array conductance measurements provide a method to indirectly monitor the temperature gradient at


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the pore vicinity. Yet, an analytical model of pore conductance temperature-dependence needs to be carefully derived, where other factors may also contribute. Previous works have outlined two main contributions to the pore conductance: pore volume conductance ( ∝ d 2 ) and surface-charge-related pore surface conductance ( ∝ d for a given pore length).

40,41

While temperature has negligible influence on the surface-related conductance

components on large pores, absorption of light itself can excite surface charge as has been reported on SiN nanopores using a 532 nm laser.40 Nevertheless, we observe that the array conductance response rate (∆G/P) scales linearly with cross-sectional area of the pore ( ∝ d 2 ), as shown in the inset in Figure 4a. This suggests that the pore conductance change is induced dominantly by the thermal effect in the volume of the pore, rather than by a surface effect. Indeed, the surface effect plays only a minor role on large pores because of a small surface/volume ratio. Moreover, as the laser power sweeps with a 10 mW step, the ionic current responds positively with a time constant close to 1 ms, which is longer than the previously reported time constant of a surface charge induced response.40 The time-dependent simulation shows a similar transient time for the system to establish a new thermal equilibrium (see Figure S6). This further confirms the light-induced temperature increase across the membrane.

Figure 4. (a) Pore conductance responses to the laser power on membranes with average pore sizes of 56 nm, 87 nm and 116 nm respectively. The inset shows that the laser induced conductance increasing rate with laser power has a linear relation with the square of the pore diameter. (b) In DNA translocation experiments, each events duration (On time) decreases with increasing laser excitation power. No significant difference is observed between 450 bp and 3360 bp DNA. The relative water viscosity to 20 oC is overlaid on top (blue curve) suggesting that


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event on-times are reduced due to faster migration in the liquid. Data were collected on pores with average pore diameter of ~ 90 nm.

To investigate the detection of molecules under different thermal gradients, we measured DNA translocation at a fixed cross-membrane bias of ~0.9 V and swept the laser power from 6 mW to 60 mW. The most noticeable change of the detected events is that the event duration reduces significantly with increasing laser power, as shown in Figure 4b. (Extractions of event time are discussed in section 3 in SI). An overlay of water viscosity that is relative to that at 20 oC is added on top of the Figure 4b graph according to the corresponding central temperature of the laser illuminated surface simulated at each laser power. There is a clear correlation between the viscosity and the event on time. Here, we model the molecule translocation process without linearization of DNA, because of the weak mechanical confinement of the molecules in the large pores.

42

Under the same electric field and travelling distance, the event on time is inversely

proportional to the DNA electrophoretic mobility, thereby scales with the buffer viscosity

ton ∝η(T) (detailed discussion in section 3 in SI). As the temperature increases, the water inside the pore becomes more fluidic and the event duration becomes shorter. It clearly indicates that the volumetric thermal effect plays a dominant role on the molecule translocation rather than the surface effect, which would instead create an extra electro-osmotic flow slowing down the DNA translocation.40 Without linearization of DNA, the difference in the event on time between 450 bp and 3360 bp DNA is insignificant and may not be used to distinguish the two from one another. Therefore, we shift our focus to the capture rate (event interval time) to discriminate DNA sizes.


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Figure 5. Schematic of DNA capture mechanism. Close to the pore, the movement of DNA is governed by both the electric field (black line) and the thermal gradient (red: hot, blue: cold).

v

v

The thermophoretic force and the electrophoretic force are denoted as Ft and Fe , respectively. The pore diameter d and length l and the absorbing radius r0 are also shown. The physical geometry is not to scale.

Molecule-size-dependent capture rate. The light-induced heating of the membrane creates a thermal gradient that extends from the pore into the bulk liquid, influencing migrations of molecules outside the pores, thereby controlling the molecular capture rate of a nanopore. Figure 5 shows a schematic of the capturing mechanism of a DNA molecule in our system. In addition to the random Brownian motion and viscous drag against the molecule moving, molecules are

uuv subjected to two competing forces, the electric force ( Fe ) towards the pore and the

uv thermophoretic force ( Ft ) repelling molecules from the pore. The capture rate of molecules is then dependent on the outcome of these competitive forces. The electroosmotic flow is neglected here since the flow rate is negligible outside the pore.43 In fact, the temperature rise enhances both processes with a decreasing viscosity (Fig. 4b), i.e. an increasing electrophoretic mobility, and by an increasing concentration depletion. Therefore, instead of varying the temperature we fix the laser power at 60 mW and investigate the capture rate by varying the applied bias. Increasing the bias is equivalent to extending the probed volume around the pore by changing the balance between the two forces depicted in Figure 5. Since D , µ ,

DT are independent of the

electric field in the range applied, varying the bias does not affect the thermophoretic term.


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The translocation experiments of dsDNA with lengths of 225 bp, 450 bp, 1000 bp, 2000 bp and 3360 bp, respectively, were conducted. The results reveal two characteristics of the nanopores combining thermo-electrophoresis. First, as shown in Figure 6 (a), the “threshold bias”, indicating the start of DNA translocation determined by the decreasing pore array conductance and the onset of fluorescent signals, scales with DNA length. A set of data as an example for “threshold bias” extraction was shown in Fig. S4 in SI. Considering the relative large ramp-up steps (~ hundreds of mV) in threshold finding experiments, the “threshold bias” shown in Figure 6 (a) has relative large uncertainties. Therefore, we have put quotation marks on the term with respective error bars in Figure 6 (a). Second, as shown in Figure 6 (b), the capture rate of the shorter DNA strands versus bias rises faster than the longer molecules in the investigated bias range (Extraction of the characteristic capture rate from the optical measurement is discussed in details in SI, S4). Such experimental observations can be well explained by thermophoresis of the molecules in the thermal gradient at the pore vicinity. The Soret coefficient of DNA, which characterizes the degree of thermal depletion, scales with the square root of the DNA length in number of base pairs, ST ∝ NB .29 Longer DNA molecules have a higher ST than the shorter ones, thus being more depleted from the pore in a thermal gradient. This leads to a higher threshold bias and a lower capture rate for the longer DNA. It is also noticed that the threshold bias difference is more pronounced for the shorter molecules, as predicted by the square root dependence. To further confirm that the thermophoretic effect is induced by laser illumination, we switch off the laser excitation while applying a constant bias below the “threshold bias”. Now without the laser induced thermal gradient, the molecules should come to the pores by electric attraction. Indeed, the pore array conductance gradually decreased by ~50 % after the laser was switched


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off, indicating an increasing number of translocation events. The presence of the molecules was confirmed by their fluorescent signals at the pores when switching back the laser. The DNA translocation events rate then reduced gradually to zero again because of the re-established thermal gradient (data shown in SI, section5, Fig. S12). Such molecule behaviors indeed prove that the thermophoresis plays central role in our experiments.

Figure 6. (a) The “threshold bias” for translocation of DNA with different lengths, determined by the pore array conductance decrease and the onset of fluorescent signals, during bias rampups. Square dots are experimental data. Red line is extrapolated bias from the analytical rate equation (2) when capture rate equals to 1.5 Hz. (b) Capture rate versus bias for 225 bp, 450 bp, 1000 bp, and 3360 bp DNA, respectively as extracted from optical intensity trace measurements. The curves derived from the analytical model (solid lines) are fitted to the experimental data. For this set of data, the laser power was set at 60 mW. All DNA samples have an initial concentration of 100 pM in the cis chamber. The data were collected on pores with an average pore diameter of ~ 130 nm. Theoretical analysis. Now, let us derive the capture rate quantitatively. Previous works have derived an equivalent capture radius used in the classic Smoluchowski rate equation to predict the molecule capture rate by dividing the volume outside the pore into a diffusion-governed regime and a capture regime when either an electric field or a thermal gradient is present at the nanopore.19,20,44 However, in our case, the opposing thermal gradient extends tens of micrometers into the bulk liquid, far beyond the capture radius induced by the electric field. Thus we need to use a more general approach to describe the molecule capture rate into a nanopore by directly solving the differential equation of the molecular flow density. Here we consider DNA


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molecules as charged particles where the particle flow density ( j ) in our system is governed by Brownian diffusion, electrophoresis and thermophoresis, which is determined by the gradients of particle concentration c, potential V , and temperature T , respectively,

j = −D(T)∇c + cµ(T)∇V − cDT (T)∇T .

(1)

As suggested previously, the potential profile at the pore vicinity can be modeled as a point charge located at the pore entrance, written as

V (r) = d 2∆V 8lr , where d is the pore diameter,

l is the pore depth, ∆ V is the applied cross-membrane bias and r is the distance from the pore. 20 The thermal gradient, however, is dependent on the pore position relative to the laser spot. For a proof of concept, we solve equation (1) for the pore in the center of the laser spot, where we consider the thermal gradient is approximately spherically symmetric (see in figure 5 for geometry). We used the approach outlined in ref.

19

, with detailed derivations for our case

described in the SI, section 4. Briefly, in a spherically symmetric case, by applying an absorbing boundary condition

c(r0) = 0 and c(∞) = c0 , where r0 is the radius that the molecule being

immediately absorbed into the pore and removed from the system (absorbing radius), we obtain the particle flow, i.e. the capture rate as R=

where

2π D (T0 )c0 , ∞η   (r) ( T ) µ 0 exp  − V ( r ) + ST (T ( r ) − T0 )  dr ∫r0 relative ( r2 D T 0)  

ηrelative =η(T) /η(T0)

is the relative viscosity and

(2)

T0 is the ambient temperature in the

buffer. Now it is clear from equation (2) that the molecule capture rate depends on the competition between V ( r ) and T ( r ) . We assigned values to parameters in equation (2) according to the current experimental data or took them directly from the literature

22, 36,41

(see in SI, section 4). The rates obtained from


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equation (2) are plotted as solid curves in Figure 6 (b) as a function of bias. The curves agree with the experimental data points when the magnitudes of

ST are in the order of 10-1 K-1 for all

the DNA lengths that were investigated. The magnitudes of

ST for different DNA lengths turned

out to be approximately in accordance to the well-known square root relation, ST ∝ NB (Fig. S11).29 Such fitting of the experimental data with only one parameter (ST) yields reasonable values for the Soret coefficient, as discussed in the Supporting Information, section 4. Here, we work in a system with a positive ST , which leads to a depletion of molecules. For

ST < 0, the

model predicts an enhanced capture rate, which indeed has been observed in DNA translocations through a plasmonic heated nanopore in LiCl.45 If we remove the thermophoretic term from equation (2), the curves of different DNA lengths will collapse into one curve, independent of DNA length as previously reported by Wanunu et al. 20 Therefore, we conclude that the thermal gradient at the pore vicinity causes distinct capture rates of molecules with different Soret coefficients. The capture rate in the presence of both a thermal and a potential gradient can thus be well predicted analytically. Here, we want to emphasize that all DNA translocation experiments are performed on arrays of pores with a diameter (50 – 130 nm) which is similar to the radius of gyration of DNA studies in this work, but much larger than the dsDNA strand width (~ 2 nm). Although the pores are too large to linearize DNA strands to distinguish its length by the event duration, by introducing a thermal gradient at the pore vicinity, we can distinguish DNA sizes by their capture rates (inter arrival times). Such nanopore array systems combining electrophoresis and thermophoresis of DNA can be potentially used for sorting molecules. While a potential gradient keeps a steady flow of charged molecules, the thermal gradient selectively blocks them by different degrees of


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thermal depletion at the pore entrance. It is clear that the sorting works most efficiently for the short DNA molecules as the bias difference per unit base increases with decreasing DNA length, as can be seen in Fig. 6 (a). The smallest DNA length difference studied in this work was about 200 bp. However the size resolution can be reduced even further by an order of magnitude for DNA < 100 bp. The bias difference ~ 100 mV would be required to separate a 20 bp length difference, which is a practical limit set by the bias drift across the membrane. Compared to many other methods of molecule size separation, such as capillary electrophoresis (CE)13, deterministic lateral displacement (DLD)46, anisotropic nanofilter array (ANA)47, etc., the size resolution and separation efficiency (flow rate) of the sorting method proposed in this work need to be further improved to be competitive. However, the feature of the sorting mechanism combined to a nanopore platform is the potential of real-time high-throughput single-molecule detection. For example, DNA molecules in a physiological solution with the same length but different sequences can be labelled with different fluorophores, thereby being distinguished individually by fluorescence detection. In fact, when sorting is combined with a nanopore platform that enables single molecule detection, the high flow rate seems not necessary. In this case, the flow rate is actually defined by the detection time resolution. For high throughput sorting utilizing all the pores, a thermal gradient can be simply introduced by resistive heating of the whole membrane by adding an extra metal layer.48 Thereby, one can really take advantage of the large scale fabrication of solid-state nanopore arrays as demonstrated here. A further combination with plasmonic sensing is also possible. 49

CONCLUSION


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In summary, we have demonstrated a platform of nanopore arrays for sorting molecules by size, combining a thermal and a potential gradient at the pore vicinity. The large arrays of pores are fabricated by electrochemical etching of wafer-scale-pre-patterned silicon chips with close to 100% etch-through rate. Laser illumination is brought to the nanopore membrane as a source for fluorophore excitation and for local heating. The light-induced thermal effect on pore conductance and event time are investigated systematically for several different length DNA strands. We show that the local thermal gradient leads to a size-dependent depletion of DNA at the pore entrance where a difference in size can be distinguished by their capture rate into the nanopores. This size selection process can, in turn, be controlled by the external electric field allowing a size-dependent molecular translocation turn-on effect with increasing bias. The capture rate can be well predicted by an analytical model employing the Soret coefficient to characterize the degree of thermal depletion of molecules. This platform can be potentially used for high throughput molecular sorting to take advantage of the large scale fabrication ability of the silicon processing technology.

MATERIALS AND METHODS Nanopore fabrication: Silicon membranes are fabricated on a 100 mm silicon-on-insulator (SOI) wafer by standard cleanroom processing. The wafer has 2.3 µm thick device layer on top of a 1 µm buried oxide layer on a 500 µm thick substrate. 2 µm-by-2 µm squares with pitch distance of 4 µm are patterned in the hard mask (100 nm SiO2) on top of the silicon device layer by photolithography and reactive ion etching (RIE). Similarly, membrane areas (~120 µm-by-120 µm) are defined on the backside of the support silicon layer (handle layer), and subsequently etched through the 525µm-thick handle layer by inductively-coupled-plasma (ICP) etching. The wafer is then immersed


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in KOH (30 %) briefly to etch anisotropically inverted-pyramids in the silicon device layer, serving later as etching pits for the electrochemical etching. The wafer is then cut into 15 x 15 mm chips and anodic etching of the silicon device layer is performed in HF in the breakdown regime. Since the electrochemical etching of silicon requires participation of positive carriers (holes), we choose the device layer to be n-type silicon, with resistivity of 1-4 Ω·cm, in order to have a better control of the pore formation. Before etching, the chip consisting of 32 membranes is put in oxygen plasma for 2 minutes to increase wettability. Then, the chip is mounted to a double electrochemical cell in between two fluidic chambers. The schematics can be found in a previous paper. 18 The front-side of the membrane is in contact with the etchant (5,4wt % HF), while the backside chamber is filled with NaCl solution to form electric contact to the backside of the chip. A potential is applied across the membrane by a potentiostat (VersaSTAT 4, Princeton Applied Research) via a Pt wire that immerses in the electrolyte in each chamber. We apply 10 V for 120 s and then 5 V for another 120 s. By using 10 V, we initiate etching by avalanche breakdown. As etching proceeds, the curvature at the pore-tip grows even higher that allows us to use lower bias to maintain etching only at the highest curvature by band-to-band charge carrier tunneling. To gain better control of the etching, the whole cell is kept in dark to prevent photon-induced free carriers. In addition, the buried-oxide-layer (1 µm) is kept during etching to prevent liquid contact between the two chambers once the device layer is etched through. After electrochemical etching, the buried-oxide-layer is removed in 5% HF. Preparation of DNA: DNA amplicons with designed lengths of 225 bp, 450 bp, 1000 bp, 2000 bp, 3360 bp and 7118 bp were prepared by PCR amplification of oligonucleotide primers with fluorescent labels at the 5’ ends. All PCR amplicons except the 7118 bp product were prepared with 1x Phusion HF


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Mastermix (Thermo Scientific), 500 µM dNTPs (Thermo Scientific), 2.5 ng of plasmid DNA (Bioneer), and 400 nM of each of

forward and reverse primers (see S6. Oligonucleotide

sequences in SI). The PCR protocol started with 1 min at 95 °C, followed by 22 cycles at 95 °C for 20 sec, 65 °C for 30 sec, 72 °C for 30 sec, and ended with 4 min at 72 °C. The 7118 bp amplicon reactions were prepared with 1 x PrimeStar GXL Buffer (Takara), 500 µM dNTPs (Takara), 1 U PrimeStar GXL Polymerase (Takara), 2.5 ng of Human gDNA (Coriell NA12878) and 400 nM each of primers 7118_FW and 7118_RE. The PCR protocol started with 1 min at 98 °C, followed by 22 cycles at 98 °C for 15 sec, 65 °C for 30 sec, 68 °C for 4 min, and ended with 4 min at 68 °C. Following PCR the amplicons were purified by polyethylene glycol precipitation on carboxylic-acid beads using a Magnatrix™ 1200 Biomagnetic Workstation (NorDiag ASA) as described by Lundin et al. 50. DNA product lengths were then confirmed by gel electrophoresis using a 2100 Bioanalyzer (Agilent) instrument and sample concentrations were determined by Qubit 3.0 (Thermo Scientific). Conductance and DNA translocation measurements: The conductance and DNA translocation measurements were carried out in the same fluidic cell, in which a nanopore chip is mounted. Prior to each measurement, the chip is cleaned in a mixture of H2SO4 and H2O2, then by oxygen plasma to ensure wetting. A PDMS sheet with one open channel is aligned to a single membrane on the backside of the chip. Then the membrane is pushed closely to the quartz window with a gap of ~ 20 µm. Both cis- and trans-chambers are filled with 1 M KCl as buffer solution. The electrical measurement of the system is controlled by a potentiostat (Versa STAT 4, Princeton Applied Research) with two Pt wires serving as working and counter electrodes. Two commercially available minimized Ag/AgCl electrodes act as


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reference and sensing electrodes, monitoring the cross-membrane bias. All the DNA translocation experiments were performed on the inverted microscope (Zeiss) with a 514 nm diode laser excitation (Omicron Phoxx) through a window-correction objective lens (Zeiss 63× NA 0.75). The PL signal is then collected by a CMOS camera (Hamamatsu V42, Photometrics Prime) in crop mode at 1 kHz. The optical setup is kept in a closed opaque box. After thermal stabilization, the ambient temperature is 25 oC in the box. Thermal simulation: The laser-induced temperature profile is simulated numerically by the finite element method using the Comsol Multiphysics software. The simulation volume is a cylinder of diameter 400 µm and of height 393 µm. From bottom to top the layers are a 170 µm quartz window, a 20 µm water layer, a 2.3 µm silicon membrane, and a cylindrical trench of diameter 120 µm in the center of the 200 µm thick supporting silicon and 1 µm thick buried-oxide layer. Here, the volume of the pore is negligible compared to that of the membrane, thus we assume the absence of pores in the simulation will not lead to a large deviation in the temperature profile from the real case. The heat transfer by conduction can be described by the heat equation, written as: ρC p (

∂T v + v ∇T ) − ∇( K ∇T ) = H , ∂t

v

where ρ is the mass density, Cp the specific heat, v the velocity of media relative to the heat source, the heat source K the thermal conductivity and the H the heat source. At stationary state, the time-dependent term is zero. The heat source comes from laser absorption in the silicon membrane. For light of wavelength 514 nm, the light absorption in silicon drops exponentially with distance and using an absorption coefficient of ~9×104 cm-1,

35

more than 90 % of the

penetrated light is absorbed by the silicon membrane. In silicon the laser-induced excitation and optical emission rate is low in comparison to the thermalization rate, thus we consider the


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absorbed laser energy as being directly transformed into heat. Here for simplicity, we ignore the light absorption profile, assuming the total energy deposition occurs fully at the membrane surface. The laser beam as a heat source is then modelled as a 2-D Gaussian profile with the integral power as the attenuated laser power at the membrane surface. Considering transmission loss in the optical path and reflection from the quartz-water-silicon multilayers, the attenuation

v

coefficient is 0.55 in our experiment. v is set to zero because of no relative move between the cell and the laser beam. Other parameters in the equation are defined according to the materials. The initial temperature is set as 25 oC. The surface of the simulation cylinder, except the bottom surface, is set as a heat sink, maintaining a constant temperature of 25 oC. The boundary condition at the bottom surface, i.e. the quartz/air interface has an air convection boundary condition with a heat transfer coefficient of 10 Wm-2K-1. The simulation volume is divided into a sufficiently fine mesh with roughly 530 000 domain elements. A finer grid has no significant effect on the solution.

ACKNOWLEDGEMENT: Authors are grateful to F. Pevere for the ImageJ plug-in modification, T. Schmidt for designing the electrochemical etching cell. The project was supported by a grant from Olle Engkvist Foundation and from Carl Tryggers Foundation, as well as by an earlier grant from the SSF foundation (Swedish Strategic Foundation). SUPPORTING INFORMATION: The following files are available free of charge via the Internet at http://pubs.acs.org.


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•

File name1: A video of 450 bp dsDNA translocating through an array pf pores

•

File name2: characterization of pores, electrical and optical signal of DNA translocation, light-induced thermal gradient, extraction of the event on/ off times, derivation of the capture rate equation, discussion on the value of the parameters in the rate equation, laser-off experiments and the oligonucleotide sequences.

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Figure 1


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Figure 2 1176x498mm (72 x 72 DPI)


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Figure 3


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Figure 4


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Figure 5


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Figure 6 345x499mm (72 x 72 DPI)


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Table of Content graphic. Left: A time-integrated PL image of a video of DNA translocation through an array of nanopores. Right: schematic of a pair of forces on a molecule resulted from a electric field and a thermal gradient respectively. The thermophoresis of DNA of different lengths leads to distinguished event frequencies. 559x252mm (72 x 72 DPI)


Thermophoresis Controlled Size-Dependent DNA Translocation

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