Electronic Biosensors Based on III-Nitride Semiconductors.

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Review in Advance first posted online on May 27, 2015. (Changes may still occur before final publication online and in print.)

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Annual Review of Analytical Chemistry 2015.8. Downloaded from www.annualreviews.org Access provided by University of California - San Diego on 06/06/15. For personal use only.

Electronic Biosensors Based on III-Nitride Semiconductors Ronny Kirste, Nathaniel Rohrbaugh, Isaac Bryan, Zachary Bryan, Ramon Collazo, and Albena Ivanisevic Department of Materials Science and Engineering, North Carolina State University, Raleigh, North Carolina 27695; email: [email protected]

Annu. Rev. Anal. Chem. 2015. 8:8.1–8.21

Keywords

The Annual Review of Analytical Chemistry is online at anchem.annualreviews.org

semiconductors, surfaces, biosensors, nitrides, field-effect transistors

This article’s doi: 10.1146/annurev-anchem-071114-040247

Abstract

c 2015 by Annual Reviews. Copyright All rights reserved

We review recent advances of AlGaN/GaN high-electron-mobility transistor (HEMT)-based electronic biosensors. We discuss properties and fabrication of III-nitride-based biosensors. Because of their superior biocompatibility and aqueous stability, GaN-based devices are ready to be implemented as next-generation biosensors. We review surface properties, cleaning, and passivation as well as different pathways toward functionalization, and critically analyze III-nitride-based biosensors demonstrated in the literature, including those detecting DNA, bacteria, cancer antibodies, and toxins. We also discuss the high potential of these biosensors for monitoring living cardiac, fibroblast, and nerve cells. Finally, we report on current developments of covalent chemical functionalization of III-nitride devices. Our review concludes with a short outlook on future challenges and projected implementation directions of GaN-based HEMT biosensors.

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1. INTRODUCTION


A general trend in research and society is the urge to continuously sense, collect, diagnose, and track data. Given the widespread use of smart phones and increasing cost of health care, more decentralized monitoring and interpretation of biomedical data are needed. In addition, a constant threat from biological warfare and epidemic plagues has risen due to globalization, freedom of travel, and terrorism. Sensors that are small, reliable, cheap, disposable, and easy to handle are needed to detect viruses, bacteria, antigens, and cells. A biosensor typically consists of three components: the analyte or receptor, a transducer, and an attached electronic system for analysis (i.e., computer) (1). Different types of biosensors covering all facets of analytical chemistry, material science, and device engineering have been proposed in the literature. Most sensors can be classified as either optical or electronic (2, 3). A typical optical biosensor will detect a biomolecule or cell by measuring the interaction of light with the target (4). Examples of possible interactions can include transmission, reflection, or detection of a Raman signal (5). Optical biosensors are very sensitive and can achieve single molecule detection. However, some optical biosensors rely on special sample preparation, are cost intensive, and need bulky setups including an optical excitation and detection system (6). Consequently, only few optical biosensors have accomplished the transition from the lab to commercial applications. Certain electronic biosensors exploit the interaction of a chemically modified surface and a biological target (7). Devices relying on this mechanism of detection include those containing functionalized metal, insulator, and semiconductor surfaces. In addition, sensors composed of nanoscopic entities have been achieved (8, 9). To date, electronic biosensors using semiconductorbased field-effect-transistors (FETs) with a chemically modified channel surface as a transducer are the most promising and practical. This is primarily due to the established and consistent high-volume manufacturing procedures with which they are fabricated. FET-based biosensors have several advantages such as negligible absorption/diffusion of ions in biological solutions and screening (10). However, sensing modalities employing receptor molecules immobilized on metal surfaces have exhibited degradation in sensitivity. Absorption/diffusion of ions limits the operational lifetime most notably in sensors containing oxide layers. In addition, FET-based biosensors can provide quantitative information on target/receptor interactions and greater sensitivity for analyte detection compared to other electronic biosensor designs (11). When a charged or polar target analyte makes contact with the FET channel surface, the current through the device changes, and this alteration can be directly measured (12, 13). The advantage of this sensing approach is that the target analyte does not need to be labeled, and as long as the device is properly functionalized one can quantify different (bio)molecules, cells, and bacteria (14, 15). Thus, FET-based biosensors can be used in different solutions, utilizing less power, while being mass produced economically and adapted into multiplexed detection strategies. Given these attributes, FET-based electronic biosensors fulfill most demands for broad adaptions. Electronic biosensors research has recently intensified due to significant advances in growth and fabrication of semiconductors, as well as strategies for surface functionalization (16–18). Sensor optimization routes typically focus on the improvement of the analytical parameters such as sensitivity, specificity, stability, and selectivity. Various studies have shown that the complexity of the bioanalytes with clinical relevance can be best addressed using FETs composed of compound semiconductors. Given their extraordinary stability in air, water, and blood environments; excellent biocompatibility (19); and good control of electrical, structural, and surface properties, III-nitride-based FETs show the most promise for compound semiconductor biosensors (20).

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This review discusses the recent development of compound semiconductor biosensors. We focus on nitride-based biosensors. The III-nitrides comprise AlN, GaN, InN, and BN and their ternary alloys (21). These compounds are typically nontoxic, offer low solubility in water and high stability in blood, and have a controllable surface stoichiometry and well-known electronic properties (22). III-nitrides and particularly GaN have been studied since the 1970s. GaN is typically semi-insulating or n-type as a result of the incorporation of oxygen impurities or n-type dopants (23). The rise of the III-nitrides started with the discovery of magnesium as a p-type dopant through the application of a thermal activation process (24). Since then, III-nitrides have been used for a broad range of applications including visible and UV LEDs, blue lasers, amplifiers, and switches including FETs and high-electron-mobility transistor (HEMT), solar cells, and recently biosensors. First, we review the properties of GaN and other III-nitrides with relevance to biosensor design. Second, we examine the principle of operation of FET-based biosensors and different design parameters and review the performance characteristics of demonstrated devices. Third, we evaluate the necessary steps to functionalize and stabilize the biosensor surface. Next, most recent detection results are presented. We conclude with an outlook that discusses the major challenges and opportunities for analytical research with these devices.

2. III-NITRIDE MATERIALS’ PROPERTIES WITH IMPORTANCE TO BIOSENSORS Biosensors containing III-nitride-based FETs utilize the interaction between the surface of the compound semiconductor and the molecules that interact with the surface. III-nitrides can crystallize in cubic and wurtzite lattice structures (25). Because of better control of the electronic properties and higher achievable crystalline quality, most devices use crystals with wurtzite structures grown in the c-direction, as shown in Figure 1a. The unit cell of a wurtzite III-nitride crystal consists of two group-III and two nitrogen atoms. The wurtzite crystal structure is a member of the hexagonal crystal system. Figure 1b shows important planes of a hexagonal III-nitride crystal. Typically, crystals are grown in the c-direction such that the polar c-plane forms the surface (26). More recently nonpolar in the a- and m-direction has been investigated where the surface is formed by the m- ({1–100}) and a-plane ({11–20}), respectively (27). III-nitride semiconductors are typically grown by several different methods and on nonlatticed matched substrates such as sapphire, SiC, or Si (28). A large lattice mismatch between non-native substrates and III-nitride layers leads to high dislocation densities, strain, and reduced crystal quality of the layers. Native substrates such as bulk GaN or AlN crystals have become available only recently (29). Substrate dislocation densities ranging from VG3

Bioassay

VG1 > VG2 VG1 > 0

Drain voltage

VG = 0, I D = 0 Normally-off

Figure 3 (a) Schematic of an AlGaN/GaN HEMT with source, gate, and drain. The 2DEG is due to the piezoelectric polarization of the strained AlGaN layer. (b) Optical image of a processed AlGaN/GaN-based HEMT. To be used as a biosensor, the gate area is functionalized. (c) Band diagram of an AlGaN/GaN HEMT. The 2DEG is generated at the interface between AlGaN and GaN. Width of the 2DEG is controlled by the gate voltage. In a normally-off device, the 2DEG would be nonexistent at VG = 0 V, leading to zero drain current. (d ) I–V curves for different gate voltages. For a normally-off HEMT, the drain current increases with applied gate voltage. Presence and quantity of the target analyte is detected by a change of the drain current. (e) A complete biosensor measuring system. The HEMTs will be integrated in the bioassay. Abbreviations: 2DEG, two-dimensional electron gas; DAQ, data acquisition; HEMT, high-electron-mobility transistor; I–V, current–voltage.

region is formed that repels electrons from the AlGaN and eventually also from the 2DEG. Thus, for a normally-off GaN HEMT, no electrons are available for current to flow between the source and the drain. However, if a potential (voltage) is applied at the gate, the depletion region can be reduced or nullified. As a result, more current will be able to flow from the source to the drain through the 2DEG channel. Figure 3d shows an according current–voltage (I–V; i.e., the voltage applied between source and drain) curve for different gate voltages. The drain current is controlled by the gate voltage. When a device is used as a biosensor, the gate area is modified to contain a biomolecular layer capable of capturing a desired target analyte (38). When a target analyte attaches to the gate area, its dielectric properties will change the potential at the gate, leading to a change in drain current (red arrows in Figure 3d ). Two main configurations have emerged in the literature following extensive research. They contain either metallized or nonmetallized gates. The advantage of a metallized surface is that it provides better control of the operation regime of the HEMT (especially for normally-off devices), and depending on the analyte can lead to better functionalization efficiency. However, the metal gate increases the distance to the channel and affects screening. This results in a decrease in the sensitivity of the biosensor. Figure 3e shows a complete biosensor measuring system. The HEMTs is integrated into a bioassay platform with interconnects to a data acquisition (DAQ) system and a computer for analysis. For a GaN-based HEMT biosensor in operation, any drain and gate voltage can be chosen (Figure 3d ), leading to a well-defined and constant drain current. The presence of the target analyte is detected by the observation of a change of the drain current of 8.6

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the biosensor via the interaction of the analyte and the functionalized gate surface. Further change in the drain current can be used to determine absolute concentrations of the analyte. Analyte quantification requires careful calibration of the HEMT. The limits of the operation are reached when the surface of the gate is saturated such that excess amounts of analytes cannot bind to it.


4. SURFACE CONTAMINATION OF HIGH-ELECTRON-MOBILITY TRANSISTOR BIOSENSORS As-fabricated III-nitride devices do not have dangling bonds due to the formation of N-N bonds on the surface. Typically all GaN surfaces are Ga-terminated (either Ga- or N-polar) (39). In addition, III-nitride surfaces are reported to be highly reactive with oxygen, carbon, and hydrogen (32). Figure 4a shows the surface molar ratios of oxygen, gallium, and nitrogen as measured with X-ray photoelectron spectroscopy (XPS) on a Ga- and N-polar surface of GaN before (left) and after HCl cleaning. In particular, the N-polar surface has a high amount of oxygen on the surface. Similar results are found for the other III-nitrides. This surface layer hinders the direct covalent functionalization of the device surface. Although very unstable, removal is not trivial and any contact with air or fluids will lead to a new oxide layer. Thus, removal of the oxide/hydroxide surface layer and passivation of the surface is crucial for the development of reliable GaN-based biosensors. Typical treatment of the III-nitride materials includes cleaning in acetone, methanol, and DI water with subsequent etching in acid. For example, GaN etched in ammonium sulfide (NH4 )2 S showed a removal of the oxide layer, which was indicated by an increase of the surface recombination velocity by a factor of 2.2 (40). The observation was confirmed by others, and a decrease of the Fermi energy and decreased contact resistivity upon ammonium sulfide etching was reported (41). The result is even more pronounced if an alcohol solution is used, possibly because of the lower dielectric constant of the alcohol, which increases the efficiency of the sulfur binding to surface atoms. Table 1 summarizes approaches for oxide removal, passivation techniques, and reported changes post etching.

5. BIOCOMPATIBILITY AND STABILITY OF III-NITRIDES BIOSENSORS A major advantage of III-nitrides over phosphides, arsenides, and silicon is its stability in aqueous solutions and its superior biocompatibility. We recently compared the stability of GaN surfaces with different polarities immersed in DI H2 O, pH 5, pH 9, and H2 O2 solutions for 7 days (22). Figure 4b shows the concentration of Ga in water solutions after experiments for several days. Although the concentration was higher for N-polar material, no change over time was observed. It was concluded that N-polar material is less stable compared to Ga-polar GaN. The ability of a material to perform without causing harm to a biological host is defined as biocompatibility. Compound semiconductors containing cadmium, tellurium, or arsine do not perform well with biological host materials. The biocompatibility of GaN and AlGaN has been investigated in various publications, and survival of living cells on these materials for several days has been demonstrated (17, 19, 44). Figure 4c shows the survival rate of PC12 cells on etched and functionalized GaN. PC12 cells are a widely used cell line for studying neuronal differentiation. Silicon was evaluated under the same experimental conditions (16). Whereas the cell density on silicon substrates decreases over time and eventually reaches zero, only a negligible decrease of the cell density is observed on etched GaN. For functionalized GaN surfaces, the cell density actually increased, demonstrating its superior biocompatibility. Similar results were described by Steinhoff et al. (15, 38) who studied the survival of rat-fibroblast (3T3) cells. In that case, cells www.annualreviews.org • III-Nitride Biosensors


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did not even show a decreased vitality as indicated by mikrospikes and lamellipodia. Other cell types cultivated on AlGaN or GaN include cardiac, kidney, ovary, and liver cells with similar positive results (15, 45, 46). Most studies also pointed out that the survival rate of cells on GaN is much higher than that of cells on silicon wafers, indicating superior biocompatibility of GaNbased devices. Podolska et al. (44, 47) demonstrated similar results for AlGaN with up to 35%

0.6

a

0.6 Ga-polar GaN N-polar GaN

Molar fraction

Molar fraction

0.4 0.3 0.2

0.4 0.3 0.2 0.1

0.0

1,000

N-polar GaN

0.5

0.1

Gallium

Nitrogen

0.0

Oxygen

b

Gallium

Nitrogen

Oxygen

ICP-MS N-polar H2O N-polar pH 5 N-polar pH 9 N-polar H2O2 Ga-polar H2O Ga-polar pH 5 Ga-polar pH 9 Ga-polar H2O2

Concentration (ppb)

100

10

1 0.4 0.3 0.2 0.1 0.0

Day 1

Day 3

Day 7

c Clean GaN IKVAV GaN Clean silicon IKVAV silicon

400

Cells (mm2)


0.5

300

200

*

**

*

100

0

*

Day 1 8.8

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*

** * *

Day 6

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Table 1

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Selected etching procedures applied to III-nitrides for cleaning and oxide removal (22, 42, 43)

Material


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Etching procedure

Comments

GaN

(NH4 )2 S

Measured by surface recombination velocity and luminescence; decrease of Fermi energy

GaN

(NH4 )2 S in alcohol

Increased effect compared to pure (NH4 )2 S etching

GaN

HCl

Removes most metals and impurities on surface; lowest percentage of carbon; may leave Cl

GaN

Heterostructure field (HF)

Removes oxygen and carbon

GaN/AlN

acetone, methanol, and water

Removes mainly dirt; little effect on oxides and carbon

AlN

HF

Removes most metals and impurities on surface

AlN

Phosphoric acid

50% reduction of oxygen; surface stable for hours; no difference between the mand c-planes

InN

(NH4 )2 S

Surface passivation

Al-concentration. However, they pointed out that a thin GaN capping layer might help in the proliferation of cells, given Al-concentration seems to increase cell death ratio. Using mammalian cell cultures HEK 293FT and CHO-K1, Cimmalla et al. (48) showed that typical cleaning and wet-etching procedures used in the fabrication of GaN sensors and functionalization of the surface (i.e., KOH, HCl, HF wet chemical etching, SF6 and Cl plasma etching) have no influence on the biocompatibility of the GaN material. This is an important observation, given XPS data on cleaned surfaces have occasionally found residual elements such as Cl from HCl cleaning on the surface (Table 2). Such results support the notion that residual surface contaminants have little effect on cell survival. Therefore, several available and well-studied surface cleaning treatments of GaN are acceptable for maintaining the III-nitride device biocompatibility (22, 49–51).

6. HIGH-ELECTRON-MOBILITY TRANSISTOR SURFACE FUNCTIONALIZATION The clean gate surface of an HEMT is not selective toward different analytes. An HEMT can function as a versatile biosensor after its gate region is modified by attaching receptors that can interact with specific analytes (17). This step is called functionalization. Functionalization of the surface may include cleaning, adsorption of organic receptors that can interact with the target analyte, and annealing to ensure homogenous distribution of the receptors. As an example of the involved process steps for functionalization, a flow chart for the attachment of peptides is shown in Figure 5 (16, 17). Steps 1 to 3 are typical wafer preparation steps employed in most ←−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−−− Figure 4 (a) Surface molar ratios of oxygen, gallium, and nitrogen as measured with XPS on a Ga- and N-polar surface of GaN, and surface molar ratio of N-polar GaN after HCl cleaning. (b) Gallium concentrations of d-H2 O, pH 5, pH 9, and H2 O2 solutions after 1-, 3-, and 7-day incubation with 3 × 3 mm2 N-polar, Ga-polar, and GaN surfaces as determined by ICP-MS. (c) PC12 cells present on GaN and silicon surfaces at 1 (left), 3 (middle), and 6 (right) days with NGF treatments. The bars on each graph represent etched GaN ( green), IKVAV-functionalized GaN (red ), clean silicon (orange) and IKVAV-functionalized silicon (blue). Abbreviations: ICP-MS, inductively coupled plasma mass spectrometry; IKVAV, isoleucine-lysine-valine-alanine-valine; NGF, nerve growth factor; XPS, X-ray photoelectron spectroscopy. Panel b reprinted with permission from Foster et al. (22). Panel c reprinted with permission from Jewett et al. (17). www.annualreviews.org • III-Nitride Biosensors


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Table 2

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Demonstrated III nitride–based compound semiconductor electronic biosensors, performance, and special remarks

Target

Device/functionalization

Performance

Comments


Biomarkers DNA (56)

3 -Thiol-modified oligonucleotides and complementary target DNA

0.1 M

No nonspecific binding observed

DNA (57)

Amine-modified single-stranded DNA on the self-assembled monolayers of 11-mercaptoundecanoic acid

1 μM

Better specificity than sensors using thiol processing

Perkinsus marinus (14)

Antibody

Streptavidin (58)

3-Aminopropyltriethoxysilane and N-hydroxysulfosuccinimide-biotin

4.7 pM

Operated in subthreshold regime

Src kinase (55, 59)

HEMT fabricated with gold nanoparticles and functionalized with PP58

1 pM in ionic solution

Low reaction to clinically relevant nonspecific protein

Enzymes (60)

Penicillinase immobilized on AlGaN/GaN HEMT

Detection of 2 μM penicillin

Enzyme detected via the enzymatic reaction (indirect)

Chemokines (61)

Monokine MIG, CXCL9 antibody

0.4 μM

Distinguish human and murine cells

Breast cancer antigen (62)

c-erB antibody anchored on the gate surface

0.25 μg/ml

Breast cancer patients have 140–210 μg/ml in serum; can detect clinically relevant concentrations

Prostate cancer antigen (63)

Prostate-specific antibody

1 μg/ml–10 pg/ml

Covers relevant range (∼2.5 ng/ml)

Kidney injury (45)

KIM-1 antibody

1 ng/ml

Brain injury (52)

Brain injury antibody

Botox Cells

Antibody

Mouse fibroblasts (38, 64)

First cells demonstrated to have high survival on GaN and AlN; used for radiation biophysics

Cardiac myocyte syncytium (15)

Signal from spontaneously beating cardiac myocyte syncytium observed; signals with amplitude of 70 mV clearly detected given low background noise of the AlGaN/GaN transistors

SaOS-2 human osteoblast-like cells (65)

K+ ion channel blocker TEA applied and change on signal amplitudes recorded; demonstrated that addition of TEA decreases signal amplitude; was assigned to blocking of outward K+ ion currents

HEK 293 cells from Invitrogen (47)

Calcium dosing experiments performed and significant reactions at the normal physiological level (∼2.5 Mm) observed; difference for inhibitors and activators observed, allowing for monitoring of live cell activity

PC12 cells used to model neural differentiation (19)

Demonstrated survival on GaN is superior to silicon substrates; higher cell density achieved on surfaces modified with recognition peptides

Nerve cells NG 10–15 (66)

Observes effects of neuronal toxicity; highlights possibility of long-term measurements; no coating needed, allowing for high sensitivity

Neuronal cells NG 10–15 (67)

Observed change of the drain current of HEMTs with neuronal cells after exposure to buffer; stronger reaction observed for buffer with Na supplement

Identified and distinguished between healthy and infested clams

1 ng/ml

Abbreviations: CXCL9, chemokine (C-X-C motif ) ligand 9; HEK, human embryonic kidney; HEMT, high-electron-mobility transistor; NG, neuroblastoma x rat glioma hybrid; TEA, tetraethylammonium.

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2

Wafer preparation: Growth; cutting; cleaning with acetone, methanol, and DI water; oxide removal; HCl and cleaning

3

Alkene termination: Place samples in PEEK; exposure to THF for 30 min at 60˚C; remove precipitates in THF; cleaning

6

Olefin cross-metathesis: Exposure to Grubbs first-generation catalyst at 40˚C for 30 min in PEEK; cleaning in CH2Cl2; cleaning of PEEK; exposure to heptanoic acid for 2 h at 40˚C


Chloride termination: Submersion into chlorobenzene (C6H5Cl) for 30 min at 92˚C

Hydrogen termination: Exposure to H-plasma for 10 min

5

4

Amine conjugation: Exposure to peptides in a microcentrifuge with solution of Milli-Q water and ethylene dichloride (EDC) for 3 hrs on a rocking platform; cleaning with Milli-Q water and drying with nitrogen gas

Figure 5 Schematic flow diagram of the steps involved in the functionalization of GaN with peptides. Steps 2–6 were performed in a nitrogenpurged glove box. Steps 4–5 were performed in a PEEK. Allylmagnesium chloride in THF was used for the alkene termination. Unless otherwise stated, samples were cleaned after each step in ethanol and dried with nitrogen gas. Abbreviations: DI, deionized; PEEK, polyetheretherketone; THF, tetrahydrofuran.

functionalization procedures independent of the target receptor. Step 5 yields a reactive surface. Here, the catalyst and specifics of the procedure must be adapted to Step 6, in which the target receptor is covalently bonded to the surface. Typical functionalization procedures often include the preparation of a second test, or control surfaces, to validate the efficiency of the functionalization. In addition to yielding a reactive layer and control of charge transport across the surface, the functionalized surface also serves as a protective layer, given biosensors must work under harsh conditions such as constant flows of air, water, blood, or urine. The protection of the device is provided by the formation of a barrier between the ambient and the device surface. The degree of protection and corresponding chemical modification process determine the lifetime of the device, further highlighting the importance of device functionalization. Many of the AlGaN/GaN HEMT biosensors described thus far in the literature do not rely on covalent attachment of molecules to the active region of the device (16). The most common approaches are to grow an oxide, to deposit a metal layer that is typically gold, to adsorb thick multilayers of biomolecules such as proteins, or to grow nanostructures such as ZnO rods (52). The biosensor specificity comes from the adsorbed moieties onto the AlGaN layer. In these approaches, one compromises the sensitivity for analyte detection. This is because the distance separating the sensor channel and the analytes present in solution is increased. The further from the channel a charged analyte is, the smaller the change one observes in device characteristics after binding. To date, very few reports in the literature rely on direct covalent attachment of biomolecules to the gate area (53).

7. DETECTION OF BIOMARKERS Demonstrated GaN-based HEMT biosensors cover the detection of many analytes (54, 55). All data collection involves monitoring the I–V curves of the devices under different conditions and running statistical comparisons to quantify changes. Figure 6a shows I–V curves of an initial and a peptide-treated HEMT. A significant change of the drain voltage is detected upon exposure to the peptide solution. The approach for the detection of biomarkers such as antibodies for virus, www.annualreviews.org • III-Nitride Biosensors


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a 1.0 × 10 –3

Initial Peptide treated

VG = 2 V

ID (A)

8.0 × 10 –4

6.0 × 10 –4

VG = 1 V

2.0 × 10 –4

VG = 0 V 0.0

0

2

4

VD (V)

6

c

1.0

1.0

1.0

1.0

0.8

0.8

0.8

0.8

0.6

0.6

0.6

0.4

0.4

0.6 0.4

0.4

K D1 = 4.404 × 10–11 M 0.2

K D2 = 1.596 × 10–9 M R2

0.0 –13

–12

–11

–10

–9

= 0.9948 –8

–7

K D1 = 7.875 × 10–11 M 0.2 0.0

–6

–5

0.2

K D2 = 4.667 × 10–9 M R2

0.0 –13

–12

–11

–10

–9

= 0.9979

–8

–7

Coverage

b

Δ/Δlmax


4.0 × 10 –4

0.2 0.0

–6

–5

log [peptide] (M) Figure 6 (a) Current–voltage plot of a representative AlGaN/GaN FET showing initial and peptide treated measurements. This plot demonstrates the signal decay following the etch treatment. Vg was swept from 0 V to 2 V along a Vd of 0 V to 6 V. Initial measurements were taken following an acetone, methanol, DI, nitrogen treatment to remove any residue from dicing. Etch measurements were taken following an HCl and H3 PO4 etch, and refrigerated incubation in 5 μL of 0.01 mM Tat-C solution. (b) Average surface coverage ratio as a function of the antigen concentration with equal maximum current changes for the two binding complexes on an antibody. (c) Average surface coverage ratio as a function of the antigen concentration with nonequal maximum current changes for the two binding complexes on an antibody. Abbreviations: DI, deionized; FET, field-effect transistor. Panels b and c were reprinted with permission from Huang et al. (69).

toxins, cancer markers, or DNA and proteins analytes is always very similar. In the case of virus or bacteria detection, the surface is functionalized with an antibody. Interaction between the antibody and antigen contained in the analyte solution leads to a change of the drain current of the HEMT as described above. Thus, the key to the detection of biomarkers is the functionalization of the gate area with the counterpart of the target analyte. Table 2 shows the wealth of analytes studied to date. We highlight some examples that utilize analytes from each target type. Enzyme detection was demonstrated by Baur et al. (60) using 8.12

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immobilized penicillinase on AlGaN/GaN HEMTs. When exposed to a penicillin solution an enzymatic reaction was triggered. This resulted in the formation of an intermediate enzymesubstrate complex, which irreversibly decayed into penicilloic acid and an enzyme. Penicillin concentrations as low as 2 μM resulted in measurable changes in the HEMT drain current. Functionalized HEMTs can be very selective toward different proteins: The surface will only interact with the target species and not with nonspecific proteins. Casal et al. (61) demonstrated that human and murinal proteins can be distinguished by establishing selectivity toward different species. Using AlGaN/GaN HEMTs modified with CXCL9 [chemokine (C-X-C motif ) ligand 9] monokine antibodies (human or murinal), it was shown that the biosensor was able to differentiate between human or murinal CXCL9 chemokines. The selectivity of these nonoptimized devices was approximately four for concentrations as small as 0.43 μM. Early on, DNA analytes were studied not only to allow for detection and simplification of the sequencing process of human DNA but also for functionalization with antigens for bacteria and virus. Kang et al. (56) used results from Baur et al. (60) to demonstrate the feasibility of surface functionalization with complementary target DNA using 3 -thiol-modified oligonucleotides as a binding layer. They reported a great repeatability and recovery of the device and no nonspecific reactions. However, Thapa et al. (57) pointed out that the attachment of a thiolated probe DNA involves the rearrangement of DNA film structure, which may result in an arbitrary change of the surface potential and consequently a decreased specificity of the biosensor. An alternative aminemodified single-stranded DNA on the self-assembled monolayers of 11-mercaptoundecanoic acid was proposed. As a result, a highly specific and sensitive device was demonstrated. The results were superior to those using a thiol-based functionalization process. Sahoo et al. (68) recently showed that DNA concentrations as low as 0.1 aM can be detected with a poly(amidoamine)dendrimer functionalized nanowire device. The promising results support the need for further miniaturization and potential for detection of even the smallest amounts of target DNA. Examples of GaN-based biosensors targeting virus and bacteria are also described in the literature. Wang et al. (14) detected Perkinsus marinus using functionalized Au-gated AlGaN/GaN HEMTs. The device was operated in mimicked seawater and was able to distinguish between tanks with infected and healthy clams. A good recyclability was demonstrated, indicating high potential for commercialization of the device. The detection of Streptomyces was performed using the same approach (58). Additional promising work with GaN-based HEMTs and cancer antibodies are reported in recent publications. Detecting these antibodies is possible if an antigen is available. So far, detection of breast and prostate cancer antibodies has been described (62, 63). Various antigen concentrations were tested using Au-gated HEMT devices modified with the target antigens. Concentrations and signal response were found to be in the range useful for real life detection systems. No indication of the limitation to these cancer types was found, and further investigations are envisioned. Similar positive results were achieved for antibodies released as a consequence of injury, namely brain and kidney injury (45, 52). Clinically relevant concentrations of according antibodies were detected on devices functionalized with the corresponding antigen. Several studies have reported dose response curves to assess the device sensitivity (69–72). The purpose of these plots is to compare a change in analyte concentration relative to a conductance change from an established baseline value. Figure 6b,c shows examples of such plots utilizing an AlGaN/GaN-based biosensor (69). These plots show three distinct regions of device operation. The first region is characterized by no response and a lack of conductance effects. In Figure 6b, this region is seen at concentrations below 1 pM. The second region is referred to as a response region, where over a specific range of concentrations a quantifiable detection of the analyte is possible. In Figure 6b, the response region is between 1 pM to roughly 0.1 μM. Finally, in the third region, a saturation of available recognition sites is reached and the conductance is no longer www.annualreviews.org • III-Nitride Biosensors


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affected by increasing concentrations of the analyte. At higher doses, this region is useful as a qualitative check for presence of a biomolecule and can be seen in Figure 6b at concentrations past 1 μM. Figure 6b,c represent two fittings for a two-binding site antibody model. The plot in panel b shows a fit with the assumption that the two binding events have an equal maximum current change effect on the device. Panel c assumes these current changes are not equal.

8. MONITORING LIVING CELLS


The biocompatibility of GaN and AlGaN was investigated in many publications, and a high survival rate of human and animal cells exceeding that of survival on silicon was found (17). This opens opportunities for new devices combining living cells and III-nitride biosensors (73). Such devices could find multiple applications such as detection of injury and recovery of cells, monitoring of transport processes, and cell semiconductor interfacing, including control of cell action potential. Survival and properties of many different cells on AlGaN/GaN HEMTs has been reported. The first cells demonstrating survival on GaN and AlN surfaces were mouse fibroblasts due to their robustness and accessibility (38). These cells survived without additional surface preparation and further solidified the observation that III-nitrides have high biocompatibility. Additional investigations showed the feasibility of GaN biosensors for radiation biophysics (64, 74). Hofstetter et al. concluded that cell repair mechanisms were not influenced by the GaN interface by directly observing the cells on the GaN surface after irradiation with X-rays and comparing results to glass substrates (64). The first direct observation of the interaction between living cells and GaN HEMTs was published by Steinhoff et al. (15). They cultivated cardiac myocyte syncytium on the device surface of an AlGaN/GaN EGFET array and demonstrated that the device had a very high signal-tonoise ratio and resulted in the observation of cell action potentials from the cardiac cells. The researchers were able to record spontaneous beating of rat cardiac cells and analyzed signals for a duration of 100–150 ms and a potential of 70 μV. Neuronal or nerve cells are more promising cell systems than fibroblasts, osteoblasts, or kidney cells with applications beyond basic scientific research. Envisioned applications include for example a nerve semiconductor interface or neuronal transistors. For neural and nerve cells, similar to cardiac cells, it is known that they can be electrically excited and show a measurable potential that can be recorded with an AlGaN/GaN HEMT. Foster et al. (19) demonstrated the survival of PC12 cells on a GaN surface. These cells are well-known cells for modeling neural differentiation. Further surface treatment using recognition peptides increased the survival rate. Gebinoga et al. (66, 67) demonstrated that the sodium flux of neuronal cells can be recorded using HEMTs. They investigated the cellular response of a mouse neuroblastoma x rat glioma hybrid (NG 108–15) by monitoring the drain current of an AlGaN/GaN biosensor. A decrease of the drain current after provocation of the cells by diisopropylfluorophosphate in different cell culture media (DMEM media, ZB-0 buffer, ZB-Na100 buffer) was observed. The strongest reaction was observed for a buffer utilizing a sodium supplement. The finding was interpreted as an indicator that sodium ion fluxes of the neuronal cells can be directly monitored. Similar results were obtained later by the same group using NG 10–15 nerve cells (66). The researchers observed the cell response to inhibitors and pointed out that a long-term observation of neurotoxins could be possible due to the high stability and low noise of the system.

9. RECENT DEVELOPMENTS Device modification with appropriate biomolecules, in the right orientation and coverage, has been a central challenge in adapting semiconductor devices into sensing applications that require 8.14

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direct contact with water solutions in addition to sensitivity, selectivity, and reproducibility (75). Published covalent surface chemistry reactions can be difficult to adapt to the modification of devices due to types of solvents used that result in peeling off of contacts and compromise of device integrity. Use of recognition peptides initially showed promising results and resulted in modified devices with excellent characteristics. However, the interactions between the recognition peptide sequences and the device semiconductor surfaces were weak and exposure to solutions resulted in loss of peptide coatings (76). Our (18, 77) lab has recently demonstrated that in situ functionalization using an etchant and a phosphonic acid derivative can be used to functionalize devices without compromising the FETs integrity and offering stability in water immersion. The in situ functionalization was recently adapted for the covalent attachment of peptides to GaN surfaces using a two-step procedure (78). Figure 7a illustrates the functionalization approach. Diced wafers containing the AlGaN/GaN FETs were subjected to an HCl and H3 PO4 etch. I–V measurements were recorded to assess FET integrity after each surface modification step. In situ functionalization was performed in a 1:1 mixture of ethephon and H3 PO4 to terminate the GaN surface on −OH groups. Following this treatment, a −15.46% ± 8.01% signal loss was detected compared to the etch group, which is expected after the covalent attachment of the short ethephon ligand. Devices can be easily modified with peptides using thioether chemistry. A commercially available cysteine terminated peptide sequence CGRKKRRQRRR (AnaSpec) was utilized in proofof-concept studies. Signal change compared to initial values yielded a loss of −15.15% ± 13.90% (Figure 7b). No statistically significant difference was measured following the addition of the peptide adlayer to the already etched surfaces. Current voltage measurements were repeated following a 24-h soak in DI water and a 24-h exposure to ambient conditions. Devices soaked in water demonstrated a −24.69% ± 7.20% change and those left under ambient conditions showed a change of −13.69% ± 6.65%. Analysis of the data revealed statistically significant changes after water exposure only. This was anticipated, given exposure to water is expected to result in protonation of functional groups on the amino acid residues. The electrical conductance of the devices is very sensitive to changes in charge at the device surface. XPS was performed to quantify the amount of peptides present on the surface before and after exposure to water solution (Figure 7c). Data showed a dominant N 1s peak at 400.4 eV, which is characteristic of surface amide groups after the attachment of the peptide to the device surface. The peptide molecules remained on the devices following water exposure. The analysis of the surface immediately following peptide coating showed a 0.75-eV shift toward lower peak values. This is attributed to surface charging effects caused by an excess of peptides present on the surface and is corrected for in the data displayed in Figure 7c. This shift was not observed after the soaking procedure. The XPS data support the notion that the peptide is initially present on the surface as multilayers, many of which are removed during soaking. A significant amount of peptides remain intact on the devices after solution exposure.

10. OUTLOOK The most immediate experiments to conduct to advance III-nitride biosensors are to optimize sensor parameters and to continue with the detection of additional biomolecular analytes. Continuous monitoring of bioanalytes requires operations over extended periods of time in the order of months. No current studies have examined the stability and performance of the discussed III-nitride biosensors over such timeframes or in vivo. Publications are beginning to emerge on the role of oxide formation on the AlGaN/GaN sensor behavior (79). Coupling such studies with investigations of large sets of devices is essential for sensor optimization. Recent detailed work to quantify the long-term stability of nanowire-based FETs in physiological environments confirms the need for passivation www.annualreviews.org • III-Nitride Biosensors


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a

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Figure 7 (a) Attachment of peptides to AlGaN/GaN FETs using covalent chemistry. (b) Percent change in ID values of device after indicated treatments. (c) Four XPS spectra from bottom to top showing devices after indicated treatments. Peaks denoted are N 1 s at 400.4 eV, N-Ga at 397.4 eV, and Ga auger peaks at 395.7 eV and 392.7 eV. The third spectrum from the bottom is shifted by 0.75 eV to the left to compensate for charging effects from the thicker peptide layer present on the surface.

(i.e., functionalization) of electronic biosensors (80). Long-term stability is essential to demonstrate the utility of III-nitride sensors for various health monitoring applications including implantable applications. With a look into the desire to adapt such devices for implantation in the body, the incorporation of several existing technologies is essential before GaN devices can be used in clinical applications. Multiplexed detection can be realized if different receptors are placed on adjacent transistors. Furthermore, the incorporation of the transistor sensors into microfluidic systems will offer additional capabilities associated with delivery of molecules as a result of sensor readings. It is envisioned that the integration of GaN FET biosensors into handheld devices or linked to smart devices will permit the monitoring of bioanalytes with relevance to human health and sustainability. 8.16

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A handheld device can become a reality if an integrated instrument such as the one shown in Figure 3e is first realized and appropriately packaged. The main elements that make up the proposed instrument are the bioassay part that includes the III-nitride sensors and board, the DAQ module as part of the computer/bioassay interface, and a computer (or smart phone) for data analysis and control of the experiments. The bioassay part can contain wells with individual IIInitride sensors functionalized to detect one type or different types of analytes. The DAQ system can be configured to sample all the transistors present in each well in a parallel form for data redundancy. The typical form for the data will come from measuring a source/drain current while applying a given voltage between the source and drain at a given gate potential. Data collection and analysis can be easily realized by developing custom algorithms and user-friendly apps. The main features sought in biosensors are reliable and versatile performance, accessibility, and affordability. In all of the sections of this review, we described recent work that is starting to show that III-nitride sensors can be adapted for various bioanalytical needs. Affordability is realized by state of the art processing capabilities for GaN-based transistors. A GaN-based transistor is fabricated from epitaxial thin films deposited on either a sapphire or silicon substrate. A production cost estimate for an example biochip sensor includes the following pricing elements: substrate, epitaxy, front-end processing, packaging, and testing. For a 6-inch sapphire wafer, in a commercial production environment, this amounts to $1,600 or $0.11/mm2 , assuming a 90% yield. This corresponds to a price of $0.41 per chip that can include seven individual transistors modified to sense a specific analyte. Multiple transistors per chip are needed to acquire enough data for statistics. A printed circuit board (PCB) that accommodates the chips plus a connector costs ∼$50. This low price allows for the III-nitride sensors to be fully disposable by replacing even the entire bioassay board, adding simplicity in experiment preparation and reduction in sensor crosscontamination. Accessibility is realized by the simple integration with a computer or a portable device such as a smart phone. Packaging of the chip and its integration to the PCB can be explored in two convenient forms: design of the PCB for simple substitution of the chip or complete packaging for replacing the entire board for each experiment. Both approaches can be realized by simply disconnecting a cable. One can dispose of the chips on the board or the entire board containing the chips. The fabrication of III-nitride biosensors relies on established, mature, affordable, highvolume techniques, which is in contrast to many of the nanoscopic electronic biosensors reported in the literature. The fabrication advantage of III-nitride devices and recent biosensor developments summarized in this review show high potential for adaptation into commercial products. An area where III-nitride biosensors can be easily adapted into a commercial product is monitoring of water contamination (81). Water contaminants can cause various short- and long-term health problems. The Environmental Protection Agency monitors water quality for acceptable levels of microorganisms, disinfection byproducts, disinfectants, radionucleotides, and inorganic and organic contaminants. In particular, metal monitoring relies on laboratory services, sample collection, sample processing, and data interpretation that can be complex and time consuming. The primary analytical techniques used measure quantity of metal in a specific sample. However, the really dangerous health problems arise from cumulative metal exposure that is best assessed through continuous real-time monitoring. There is a great need for a disposable, user-friendly biosensor to quantify heavy metal contamination in soil and water. More specifically, the sensor should be able to quantify cumulative exposure, rather than amounts of dissolved metal ions in a given sample. This review has described published approaches to tailor III-nitride biosensors to operate over extended periods of time in solution and to sense different analytes. III nitride– based biosensors functionalized with heavy metal binding peptides or nucleic acids (82) have the potential to meet the need to monitor the accumulation of heavy metals in water supplies.

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DISCLOSURE STATEMENT The authors are not aware of any affiliations, memberships, funding, or financial holdings that might be perceived as affecting the objectivity of this review. LITERATURE CITED


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www.annualreviews.org • III-Nitride Biosensors


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Fluorescent Biosensors Based on Single-Molecule Counting.

Biosensors based on cell and tissue material.

Biosensors based on membrane transport proteins.

Peptide-based biosensors.

Intact chemoreceptor-based biosensors.

Highly scalable, uniform, and sensitive biosensors based on top-down indium oxide nanoribbons and electronic enzyme-linked immunosorbent assay.

DNAzyme-based biosensors and nanodevices.

Electrochemical sensors and biosensors based on nanomaterials and nanostructures.

Non-antibody protein-based biosensors.

Biosensors based on electrochemical lactate detection: A comprehensive review.

Recent Advances on Luminescent Enhancement-Based Porous Silicon Biosensors.

Glutamate biosensors based on diamond and graphene platforms.

Biosensors based on porous cellulose nanocrystal-poly(vinyl alcohol) scaffolds.

Tantalum-based semiconductors for solar water splitting.

polyaniline nanocomposite.

Yeast-based biosensors: design and applications.

Scalable Production of Molybdenum Disulfide Based Biosensors.

Nanomaterial-based biosensors for food toxin detection.

Aptamer-based biosensors for biomedical diagnostics.

Last Advances in Silicon-Based Optical Biosensors.

Carbon nanomaterial-based electrochemical biosensors: an overview.

Versatile matrix for constructing enzyme-based biosensors.

Tyrosinase-Based Biosensors for Selective Dopamine Detection.

Recent Progress in Lectin-Based Biosensors.