CMOS Integrated DNA Microarray Based on GMR Sensors

Shu-Jen Han1, Liang Xu1, Heng Yu3, Robert J. Wilson1, Robert L. White1,2, Nader Pourmand3, and Shan X. Wang1,2

1. Dept. of Materials Sci. and Eng., Stanford University, Stanford, CA94305, USA 2. Dept. of Electrical Engineering, Stanford University

3. Stanford Genome Technology Center, Palo Alto, CA94304, USA

Abstract

A high density GMR sensor array was integrated with a standard CMOS chip for DNA hybridization detection. Absorption of magnetic nanoparticles by the hybridized DNA alters the sensor resistance, and generated electrical signals are directly measured with the on-die circuitry. The proposed biochip can be applied to other bio-reaction detection, e.g. protein assay, through different surface modifications.

Introduction

Molecular recognition is exploited in assay techniques such as those based on DNA hybridization microarrays. Integrating biosensor arrays and other laboratory functions on a single CMOS chip yields a low-cost system (lab-on-a-chip) that constitutes a promising tool for future biological diagnostics. Magnetic biosensors are under active development and may soon rival established biological detection methods that use surface-bond fluorescent tags [1,2]. In a magnetoresistive biosensor detection scheme (Fig.1), single–stranded DNA receptors are immobilized on the surface of giant magnetoresistive (GMR) sensors. Oligonucleotides of unknown sequence are selectively captured by complementary probes. Streptavidin coated magnetic nanoparticles are then introduced and bind to the biotin of the hybridized DNA. Finally, magnetic field disturbances due to the nanoparticles are sensed by GMR devices. In this paper, we have designed and fabricated the first GMR sensor integrated CMOS biochip. It contains >1000 sensing elements within 1mm2, together with low noise, high throughput readout channels for high sensitivity DNA detection on the same chip. Compared to complex and expensive optical detection systems, the GMR biochip measures electrical signal directly from the sensor, and makes a low-cost, highly portable device feasible.

DNA Microarray Structure

One difficulty in the design of magnetic biosensor arrays is achieving a good balance between high sensitivity (i.e., small sensors) and low binding delay (i.e. large sensors) [3]. In this work, we address this issue through a multidivided array structure. With this approach, the sensitivity to low concentration samples can be improved by combining several sensor pixels per biological sample spot. Furthermore, large-scale genetic information can be accessed using a single pixel per DNA sequence through in-situ probe synthesis [4]. Fig.2 shows a block diagram of the proposed magnetic biochip. 1008 GMR biosensors are divided into 16 subarrays, and each subarray occupies an area of 120µm x 120µm, which is compatible with state-of-the-art DNA spotters. To reduce the signal read out time from each pixel, we utilized both frequency division multiplexing (FDM) and time division multiplexing (TDM) techniques. Thus the throughput is increased (16 outcomes/reading). The final output is digitized by an

Fig.2: Magnetoresistive biosensor array chip architecture. Multiple spin valve sensors form a subarray corresponding to 1 DNA spot, and 16 subarrays share the same control bus. Each subarray has single column of reference sensors covered by photoresist (or thick oxide) for the fully differential measurement. 4 subarrays are frequency division multiplexed to form a channel (Fig.3). All 4 channels connect to a 4:1 multiplexer to enable time division multiplexing.

Fig.1: Principle of using magnetic biosensor and magnetic nanotags. The resistance of the sensor is altered by the magnetic field generated from attached nanoparticles.

off-chip ADC, and tone measurements are performed using an FFT. The readout channel is illustrated in Fig.3. For nanometer-size sensors, the reaction rate is usually diffusion limited. Fig.4 (a) shows FEM simulation results of the capture rate per sensor as a function of the sensor pitch. The bulk-dissolved molecules can be captured by the sensor through two diffusive pathways: by direct 3D diffusion to the sensor; or by 3D diffusion to the non-sensor surface, followed by 2D surface diffusion to the sensor [5]. As can be seen from Fig.4 (a), the relative contribution of these two diffusive pathways to the total capture rate highly depends on the 2D diffusion coefficient (D2d). For small D2d, the accumulation time for 1 molecule on the sensor is ~1.9 hrs if we only have a single sensor per spot. To find the optimized number of the sensors per spot, Fig.4 (b) shows the sensor density dependence of the molecule transportation and measurement time. By incorporating 64 sensors per DNA spot, we can reduce the required time to less than 5 minutes.

GMR Device Fabrication and Integration

Our biochip implements spin valve structures, the best known structures for observing the GMR effect. In a spin valve, two magnetic layers are separated by a non-magnetic conducting spacer such as Cu, and the resistance of this sandwich structure depends on the relative orientation of the magnetization in the two magnetic layers. The direction of the magnetization in one of the layers is always pinned by an exchange bias interaction with an antiferromagnetic or ferromagnetic layer. Fig.5 (a) summarizes the details of the fabrication procedure for sensor integration: 1. Passivation thinning using RIE with a SF6 and Ar mixture. The remaining thickness of the passivation layer is around 400nm. 2. Spin-valves with dimensions of 0.3µm x 5µm are

patterned with electron beam lithography (Raith-150) and a lift-off process. Films are deposited by UHV ion beam deposition (thicknesses are in nm). 3. Electrical contact holes are etched by RIE (C2F6). 4. Metal leads composed of Ta/Au/Ta are deposited for interconnection for their high resistance to corrosive chemical solutions. Finally, thin SiO2/Si3N4/SiO2 protection layers are globally coated on the chip (not shown here). The structure of a synthetic antiferromagnetic spin valve is also shown in Fig.5, and is deposited by an ultrahigh vacuum (base pressure <5x10-9 Torr) ion beam deposition system. The performance of the spin-valve is highly dependent on the smoothness of the substrate, and it degrades significantly for a rough surface. Fig.5 (b) shows the surface roughness after a long RIE step (step 1). The surface roughness (Ra) is well below 5 Å, and

Fig.3: Schematic drawing of readout channels, with expanded view of the mixer and low noise amplifier. A single 42-kHz master clock is used to generate 7-kHz, 5.25-kHz, 4.2-kHz, and 3-kHz carrier frequencies for the mixers. The LNA has a gain of 18, and the programmable gain amplifier (PGA) with gains of 1, 10, or 100 is designed to maximize the dynamic range of the system.

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Fig.4: Optimization of the sensor density per DNA spot. (a) Numerical simulation results of the capture rate per sensor as a function of the sensor pitch. The molecule concentration is 1fM, and the sensor acts as a perfect sink. It shows that non-specific absorption followed by 2D surface diffusion starts to dominate the capture rate when the 2D diffusion coefficient (D2d) approaches the 3D diffusion coefficient (D3d). (b) Experiment time versus number of the sensors per spot. Based on this optimization curve, the sensor density in our prototype chip is chosen to be 64 sensors per DNA spot.

no post-RIE residues are observed. The sensitivity of the sensor can be expressed by its magnetoresistive (MR) ratio, which is the maxima resistance change (in percent) due to the applied magnetic field. Fig.6 shows a large improvement of device performance by inserting a thin nano-oxide layer during deposition [6]. The oxidation of a rough transition metal surface tends to remove bumps and spikes, hence it smoothes the outer surfaces and enhances the specular reflectivity. The nano-oxide layers were formed by exposing the fresh metal surfaces to pure oxygen in a separated load-lock UHV chamber without breaking vacuum. The prototype chip was fabricated in 6-metals 0.25µm NSC BiCMOS technology, and the post-processed die micrograph is shown in Fig.7. Fig.8 demonstrates the sensitivity of the complete GMR biochip. The commercial magnetic nanoparticle (MACS bead from Miltenyi Biotec) with magnetization determined by the Langevin function is used in simulation. These particles are composed of 25 nm

diameter iron oxide cores and polymer matrix, and dynamic light scattering (DLS) shows the average size of the particle is around 100nm in diameter. Measured uniform field dependence of the pre-amplified signal demonstrates the minimum detectable field change is better than 0.1 Oe. A single MACS nanoparticle generates the uniform field of 0.12 Oe over the sensor area (by simulation), which indicates the presented biochip can perform single nanoparticle (molecule) detection.

DNA Hybridization Detection

To demonstrate working biochemistry and spin-valve sensors, the DNA assay was first performed using the discrete type GMR biochip. Fig.9 shows measured signals from both negative (non- complementary DNA) and positive (complementary DNA) sensors. We can clearly see

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Fig.5: (a) Micro- and nano-fabrication steps for chip post-processing. (b) AFM image of the chip surface after step 1. The surface roughness (Ra) is below 5Å.

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Fig.6: Magnetoresistive ratio (MR) and exchange bias field (Hex) as a function of oxygen pressure applied during nano-oxide layer (NOL) formation. The insertion of NOL dramatically increases the MR ratio of spin valve sensors (MR ratio without NOL was 5%) due to the specular effect.

Fig.7: Micrograph of the post-processed die and SEM images of the biosensor array. The prototype chip was implemented in a 0.25µm BiCMOS process (chosen for its availability). Post-processing was performed at the Stanford Nanofabrication Facility (SNF).

the signal change due to the absorption of magnetic nanoparticles on the positive sensor. SEM images show the difference of the coverage on these sensors, and higher particle coverage on the positive sensor can be obtained by improving the surface chemistry.

Conclusion

We have presented a fully integrated biochip based on GMR sensors and magnetic nanoparticles. The chip has been fabricated in a 0.25µm BiCMOS process, and sub-micron sized sensors were successfully integrated with the diced chip. The device shows promising potential for low cost, highly sensitive diagnostic applications.

Acknowledgments

The authors thank National Semiconductor for chip fabrication. This work was supported by DARPA.

Fig.9: (Top) Measured signals (pre-amplified) from the GMR biochip for both complementary and non-complementary DNA. The particle used here is MACS bead from Miltenyi Biotec. (Bottom) SEM images of regions of complementary DNA (left) and non-complementary DNA (right) after the measurement. It is shown that magnetic nanoparticles stick only on the complementary DNA region.

Fig.8: Sensitivity of the post-processed GMR biochip. Measured uniform field dependence of the pre-amplified signal demonstrates the minimum detectable field change is better than 0.1 Oe. A single MACS nanoparticle generates the uniform field of 0.12 Oe over the sensor area (by simulation), which indicates the presented biochip can perform single nanoparticle (molecule) detection.

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