Engineering the Bio-Electrode Interface For

Home , Detection limit

ENGINEERING THE BIO-ELECTRODE INTERFACE FOR

ELECTROCHEMICAL BIOSENSORS WITH SENSITIVITY,

ACCURACY AND SIMPLICITY

YIFAN DAI

Submitted in partial fulfillment of the requirements

for the degree of Doctor of Philosophy

Department of Chemical and Biomolecular Engineering

CASE WESTERN RESERVE UNIVERSITY

May, 2020

CASE WESTERN RESERVE UNIVERSITY SCHOOL OF GRADUATE STUDIES

We hereby approve the thesis/dissertation of Yifan Dai candidate for the degree of Doctor of Philosophy.

Committee Chair Chung Chiun Liu

Committee Member Robert F. Savinell

Committee Member Harihara Baskaran

Committee Member Blanton S. Tolbert

Date of Defense 03/2020

*We also certify that written approval has been obtained

for any proprietary material contained therein.

Great thanks to my families and friends and all the inspiring mentors!

I want to thank my advisor, Prof. Chung Chiun Liu. I started doing research with Prof. Liu since I was an undergraduate. I have been deeply appreciated for the trust and opportunities he gave me. He guided me through all my major decisions in my early career with great patience. Prof. Liu shows me more than how to be a delicate scientist. Most importantly, he shows me a great demonstration on how to be a nice individual with patience and love. It has been a great pleasure and honor for me to work with him in the past years.

I also want to thank all my collaborators and professors who inspire me a lot. I had no background about biology or biochemistry before I got into the biosensing field. All the knowledge was learnt from my collaborators and many great papers published by the experts in the field. Those works make learning so enjoyable and have taught me so much throughout my career.

Lastly, I want to thank my great friends. Yu Fei and his families have made my time in Cleveland so joyful. I am very lucky to have you guys. Special thanks to my best friend of life, Xintong Cao. You are always such an inspiring person for me with great courage and kindness. Thanks for always being there for me. Our time together has always been enjoyable. Finally, I want to thank my parents for giving me such a happy life. I love you all.

Yifan

Table of Contents

Abstract ...... 5 Chapter 1 Introduction ...... 6 Chapter 2 CRISPR based Biosensing Platforms ...... 8 2.1 Development of CRISPR based Universal Biosensing System ...... 8 2.1.1 Introduction ...... 8 2.1.2 Verification of E-CRISPR on Nucleic Acid Detection ...... 11 2.1.3 Evaluation of the Optimized Condition for On-chip Trans-Cleavage Activity ...... 12 2.1.4 E-CRISPR for Nucleic Acid Detection ...... 19 2.1.5 Aptamer based E-CRISPR Cascade for Protein Detection...... 21 2.1.6 Conclusions ...... 25 2.1.7 Experimental Procedures...... 26 2.2 CRISPR Mediated E-DNA Sensor ...... 29 2.2.1 Introduction ...... 29 2.2.2 Optimization of the E-DNA Sensor ...... 34 2.2.3 Evaluation of CRISPR Cas9 mediated E-DNA Sensor ...... 36 2.2.4 Comparison of the Cas9 with Cas12a mediated E-DNA Sensor ...... 38 2.2.5 Probing mismatches with CRISPR mediated DNA Electrochemistry .... 42 2.2.6 Conclusions ...... 47 2.2.7 Experimental Procedures...... 47

Chapter 3 Bio-electrode Interface for Protein Detection ...... 50 3.1 Enhancing the Simplicity for the Construction of Bioelectrode Interface .... 50 3.1.1 Evaluation of the Effectiveness of the Bioconjugation Method ...... 52 3.1.2 Surface Analysis of Au-S Formation ...... 55 3.1.3 Electrochemical Detection of TDP-43 ...... 56 3.1.4 Experimental Procedures...... 62

3.2 Enhancing the Sensitivity through Electrostatic Interaction ...... 66 3.2.1 NCI Construction for Analyte with Acidic pI ...... 70 3.2.2 NCI Construction for Analyte with Basic pI ...... 72 3.2.3 NCI Development for HE4 Detection ...... 73 3.2.4 NCI Development for Tau Detection ...... 79 3.2.5 Conclusions ...... 83 3.2.6 Experimental Procedures...... 84 Chapter 4 Conclusion ...... 87

References ...... 89

Engineering The Bio-Electrode Interface For Electrochemical

Biosensors With Sensitivity, Accuracy And Simplicity

Abstract

YIFAN DAI

The rapid development of personalized medicine highlights the need for a robust point- of-care diagnostic system. To address this need, in my doctoral research, I focus on engineering the bio-electrode interface for the development of sensitive, accurate and

simple electrochemical biosensing systems. The developed biosensing strategies have

been applied for the detection of nucleic acids and proteins. These advancements can be

of especially high-utility for the biosensing community, achieving the ultimate goal of

point-of-care diagnosis.

Chapter 1 Introduction

The advancement in understanding of molecular pathways on the regulation of

diseases promotes the development of point-of-care diagnosis.1-12 The information from

broad assessments of biomolecular signatures from nucleic acid, protein and cell

throughout the disease progression is utilized as a general standard to quantitatively

evaluate the severity of the disease.13-17 The existence of specific biomolecules in human

fluids is recognized as the confirmation of a corresponding disease, such as cancers,13, 18,

19 neurodegenerative disorders.20, 21 Therefore, techniques on early detection of selected

biomolecules are extremely critical for medical industry, especially meaningful for

disease diagnosis and drug treatment.22-29

Current techniques for biomolecular quantification involve the use of mass

spectrometry,30 ELISA,31-34 Western blot,35 and Northern blot.36 However, these techniques are expensive, complex and time-consuming, therefore not suitable for the contemporary needs of point-of-care disease evaluation. With the unresolved medical demands on an accurate, simple, cost-effective, time-efficient and sensitive biomarker detection system,

recent scientific endeavors lead to many novel biosensing systems, such as field-effect transistor,37, 38 surface plasmon resonance sensor,39 and electrochemical biosensor.7, 9, 10, 40-

45 Among these emerging diagnostic platforms, electrochemical biosensor stands out for wide developments toward point-of-care diagnostics owing to the advantages discussed

below.9, 46-50 Every sensing system consists of three essential elements: a sensing element,

a bio-recognition element, and a transduction element. For an electrochemical biosensing

system, it possesses a simple and cost-effective sensing element (cost less than $1 for each

sensor), such as a single-use sensor,51 or a sensor array.52 The processing method of bio-

recognition element in an electrochemical biosensing system is relatively simple and

efficient, delivering an effective bio-interface without complex operations. Furthermore,

the cost-effectiveness of the electric (signal) transduction system, such as linear sweep

voltammetry and differential pulse voltammetry, making the whole sensing system applicable in undeveloped area. Thus, the development of electrochemical biosensing system has created a positive and robust impact on global health.

For the development of ideal integrated point-of-care system, many technical challenges remain lingering, limiting the utilizations of electrochemical biosensing system. First, for the detection of early-stage disease, the abundance of biomarkers in human fluid is relatively low comparing with that of irrelevant biomolecules, hence, requiring high-sensitivity and high-differentiation ability for accurate analysis. Second, the matrix effect caused by the biomolecules (other than the target) from human fluid interferes the target recognition process (during mass transportation), delivering a high-

possibility of false positive result. In addition, for practical applications, fast turnaround

time is required, therefore the fabrication method of biosensor needs to be simple and

efficient.

In this doctoral research, the construction of the bio-electrode interface for electrochemistry based biosensor is studied in order to advance the sensitivity, accuracy, and simplicity for biomolecular detection.51, 53-60 Specifically, two classes of biomolecules,

nucleic acids and proteins, are selected to examine the development of the biosensing systems.

Chapter 2 CRISPR based biosensing platforms

2.1 Development of CRISPR based Universal Biosensing System

2.1.1 Introduction

An accurate, rapid, and cost-effective sensing strategy for the quantification of disease

biomarkers is vital for the development of early-diagnostic point-of-care systems, further

leading to personalized medicine and benefiting overall human health. Electrochemistry

based biosensing platforms have been widely developed, owing to its rapid signal

readout, affordable transduction element and simple sensing platform.55, 61-68 One of the

critical challenges for such sensing system is its accuracy. Recent robust developments of

CRISPR (clustered regularly interspaced short palindromic repeats) based gene editing

systems demonstrated the accuracy of the CRISPR system in targeting nucleic acids

owing to the complementarity dependent CRISPR cleavage event.69-75 Utilizing the Cas- crRNA target recognition-and-cleavage event induced collateral (trans) cleavage effect of

the nonspecific ssDNA reporter, CRISPR type III, V, VI RNA guided nucleases (Csm6,

Cas12a, Cas13) have been applied for the detection of nucleic acid (RNA/DNA) through

fluorescence transduction system.76-79 Herein, we extend the application of CRISPR-

Cas12a (cpf1) system to the development of an electrochemical biosensing system, owing

to its relative cost-efficiency and portable nature of the transduction system comparing

with those of the fluorescent transducer. Moreover, for a comprehensive and robust

point-of-care system, a potential universal biosensing platform that can detect different

categories of analytes is of high importance for the clinical applications. Therefore, other

than detection of nucleic acid, we further design a Cas12a based biosensing cascade as a

strategy for the detection of protein. Our study demonstrates a CRISPR-Cas12a (cpf1)

based electrochemical biosensing platform (E-CRISPR) for the detection of the major

categories of biomolecules, providing a potential deployable point-of-care system for the

healthcare industry.

The E-CRISPR introduces a simple transduction method for CRISPR type III, V, VI

nucleases based sensing systems and provides a new liberty in the classes of analytes the sensing system can detect. A disposable, micro-fabricated three-electrode sensor with

thin gold film as working and counter electrodes and Ag/AgCl as the reference electrode was applied for the development of this electrochemical biosensor.53, 56 Cas12a-crRNA

duplex was designed to specifically recognize and cleave target nucleic acid strand based on the protospacer adjacent motif (PAM) sequence of the target and crRNA sequence

(Figure 1A).80, 81 The PAM recognition depends on the specific 5’ TTTN nucleic acid sequence located at the opposite strand (blue strand) of the recognition strand (orange strand). Only upon the recognition of the PAM sequence by the Cas protein, the Cas protein, acting as a DNA helicase, would unwind the target DNA. After the separation of the target strands, the complementarity (between crRNA and target) dependent cleavage activity is further activated.82 To achieve the electrochemical transduction of

CRISPR detection signal, the target cis-cleavage initiated trans-cleavage (collateral cutting) effect of Cas12a on the nonspecific ssDNA is probed through an electrochemical method. A nonspecific ssDNA reporter is designed with a methylene blue (MB) electrochemical tag for signal transduction and a thiol moiety to tether on the sensor surface in order to acquire the signal electrically (Figure 1B).83 Consequently, the electron

transfer process between the gold electrode and the redox active species on the ssDNA is

electrochemically initiated and transduced. With the presence of the target, the Cas12a

trans-cleavage activity is activated, cleaving the MB-ssDNA reporter off the electrode surface, therefore decreasing the MB signal transduced (Figure 1C). Without the presence of the target, the Cas12a trans-cleavage activity is silenced, retaining the MB-ssDNA

reporter on the surface (Figure 1D). A representation of electrochemical signal output

based on the conditions without/with target is shown in Figure 1E. The design of the MB- ssDNA reporter covered electrode is generally applicable for any CRISPR type III, V and

VI systems as a simple and cost-effective signal transduction strategy. Based on the

concept of E-CRISPR, we developed this platform as a universal biosensing strategy for

the detection of nucleic acid, protein and small molecule.

Figure 1. Principle of E-CRISPR. A) Cas12a (cpf1) performs crRNA guided cis-cleavage (specific target) initiated trans-cleavage activity (nonspecific ssDNA). B) Nonspecific ssDNA reporter with methylene blue tag immobilized on the gold electrode. C) With the presence of the target, Cas12a-crRNA would initiate the trans-cleavage activity on nonspecific ssDNA reporter, resulting a low electrochemical current of methylene blue. D) Without the presence of the target, Cas12a-crRNA would not initiate the trans- cleavage activity on nonspecific ssDNA reporter, resulting a high electrochemical current of methylene blue. E) A representation of electrochemical current outputs based on the without & with target conditions. 2.1.2 Verification of E-CRISPR on Nucleic Acid Detection

To examine the feasibility of the E-CRISPR on nucleic acid detection, a human papilloma virus (HPV) subtype, HPV-16, which is critical to carcinogenesis84, was

selected as the target. A target sequence in the L1-enconding gene of HPV16 was identified based on the TTTN PAM sequence required by the Cas12a endonuclease.85 The

electrochemical biosensing platform was initially developed based on the Cas12a

endonuclease from Acidaminococcus sp (AsCas12a).86 We first investigated the on-chip

collateral cleavage performance based on the AsCas12a-crRNA duplex targeting the

HPV-16 sequence. After assembling the HPV-16 and the AsCas12a-crRNA, the triplex

complex was directly incubated onto the ssDNA reporter covered electrode. Square wave

voltammetry (SWV) was applied to evaluate the MB signal, which was decreased only in the presence of the cognate target with corresponding AsCas12a-crRNA (Figure 2A).

2.1.3 Evaluation of the Optimized Condition for On-Chip Trans-Cleavage activity

For biosensing application, the detection sensitivity is critical due to the low abundance of clinically relevant biomarkers in human fluids.87, 88 For the E-CRISPR

detection platform, the trans-cleavage activity is the key for signal transduction, and

therefore is critical to the sensitivity performance. We first compared the on-chip trans- cleavage activity of another type of Cas12a protein, Lachnospiraceae bacterium ND2006

Cas12a (LbCas12a),89 with that of the AsCas12a. LbCas12a demonstrated a more

apparent and stable trans-cleavage response within 5 min comparing with that of

AsCas12a (Figure 2B). LbCas12a presented a more robust trans-cleavage activity within the testing period based on the same experimental condition, therefore LbCas12a was selected for further E-CRISPR development. We further evaluated the possible factors that may affect the trans-cleavage activity for on-chip electrochemical test using HPV-16 as the target. The optimized trans-cleavage period was investigated. The ΔI%

continuously increased with the increasing incubation time for the collateral cleavage event (Figure 2c). It is interesting to notice that the trans-cleavage activity is not a simultaneous event of the cis-cleavage activity after the activation of cis-cleavage by the target. The cis-cleavage of target strand is typically finished within 30 min; 76 however,

the trans-cleavage function remained active even after 3 hours (Figure 2C), indicating the target recognition and cis-cleavage activity of Cas12a system is the activator for the

trans-cleavage domain of the Cas12a endonuclease.

Figure 2. A）Representation of square wave voltammetry (SWV) evaluation of E- CRISPR in response to HPV-16. Red curve represents the background signal of 50 nM of

Cas12a-crRNA duplex. Black curve represents the signal generated by the 50 nM of Cas12a-crRNA-target induced trans-cleavage activity. B) Evaluation of 50 nM of Cas12a orthologs from Lachnospiraceae bacterium and Acidaminococcus sp on its activity for on- chip trans-cleavage activity based on the change of current between background signal and target-mediated signal. I% = . (Red line- LbCas12a;

Background signal−Target signal Black line- AsCas12a) C) Evaluation of trans-cleavage activity using 50 nM of LbCas12a- ∆ Background signal crRNA-target triplex. D) Evaluation of the effect of the concentration of divalent metal ions on the trans-cleavage activity of RuvC domain based on 50 nM of LbCas12a- crRNA-target triplex. E) & F) & G) SWV graphs of different lengths of surface ssDNA reporters based on 30 nM of LbCas12a-crRNA-target triplex. H) Comparison of signal change from different lengths of surface ssDNA reporters. SWV graphs in these figures present the result of a single test. Error bars in figures present the standard error (SE) based on at least three individual trails using at least three different sensors.

We further investigated the chemical environment of the Cas12a to optimize the

trans-cleavage performance. An important factor that may affect the Cas12a cleavage activity is the divalent cation Mg2+ concentration in the testing solution.90 Cas12a RuvC

domain is known to cleave ssDNA through the two-metal ion mechanism,91, 92 which

involves the Mg2+ ions to induce conformational coordination of the RuvC domain and

the ssDNA by shifting the spatial distribution of ssDNA around the RuvC active cutting

center. Therefore, we evaluated the effect of concentration of Mg2+ ions in the in vitro

cleavage solution on the performance of trans-cleavage activity. The trans-cleavage

activity was only activated with the presence of the Mg2+ ions in the testing solution

(Figure 2D). Increasing concentration of Mg2+ ions up to 15 mM demonstrated an

enhanced trans-cleavage activity. Hence, an optimized Mg2+ concentration of 15 mM was selected for the preparation of Cas12a-crRNA duplex.

In order to perform an efficient surface chemistry based trans-cleavage, the accessibility of Cas12a endonuclease to the nonspecific ssDNA is important. Thus, we evaluated the effect of ssDNA reporter density on the electrode surface on the variation of electrochemical signal before and after trans-cleavage activity. An ideal surface condition can provide an optimized electrostatic environment for charged phosphate backbones and the hydroxyl groups of the passivation agents to ensure an upright ssDNA surface, facilitating the cleavage activity. The surface density of ssDNA reporter was manipulated by the concentration of the ssDNA reporter incubation solution. As shown in Figure 3A, a high surface density of ssDNA reporter significantly decreased the change of signal, because this high surface density decreased the accessibility of Cas endonuclease to the ssDNA reporter, producing a steric hindrance effect, which limited the trans-cleavage activity.

Figure 3. A) Evaluation of different concentrations of ssDNA reporter prepared interrogating electrode based on the trans-cleavage activity. The 10 nt ssDNA reporter was incubated on the

electrode for 1 hr at room temperature in the dark. 2 mM of 3-mercaptopropanol (3-MCP) in 10 mM Tris buffer (containing 100 mM NaCl) was applied to passivate the ssDNA electrode for 30 min, forming an aligned electrode surface. 50 nM of LbCas12a-crRNA targeting HPV-16 was then applied for the detection of 50 nM of HPV-16 target based on the described procedure in the manuscript. The trans-cleavage activity was then investigated by SWV. The signal change acquired by comparison with the electrochemical signal based on mutated trans-cleavage activity. The ΔI% was compared between different concentrations of ssDNA reporter prepared electrode surface. B) Evaluation of the effect of different passivation agents and different lengths of the ssDNA reporters on the probing of the trans-cleavage activity. 1 µM of the ssDNA reporter was incubated on the electrode surface for 1 hr at room temperature in the dark. 2 mM of 3-mercaptopropanol (3-MCP) or 6-mercaptoheaxnol (6-MCH) or 11-Mercapto-1- undecanol (11-MUD) prepared in 10 mM Tris buffer (containing 100 mM NaCl) was applied for passivation for 30 min at room temperature. 50 nM LbCas12a-crRNA-HPV16 triplex was applied to investigate the trans-cleavage on different performed electrode. The signal changes and standard errors of each passivation agent and ssDNA reporter pair were compared. C) The concentration of Cas12a-crRNA in the 30 uL detection solution was evaluated for the optimized performance of trans-cleavage activity. Based on the optimized ssDNA surface density, a fixed concentration of HPV-16 target of 10x nM was applied as triggering DNA for the trans-cleavage activity on the sensor surface. A 30 min of trans-cleavage period at 37 °C was applied for the test. The change of signal % was compared between different concentrations of Cas12a-crRNA duplex. An optimized trans-cleavage activation concentration was identified to be 30 nM based on the signal stability and the degree of the change of signal. An increasing concentration (>100 nM) of Cas12a-crRNA would lead to significant hindrance in the diffusion and capture of cognate target, therefore decreasing the accessibility of ssDNA reporter to the Cas12a endonuclease. A low concentration (<30 nM) would decrease the available nucleases for trans- cleavage activity, hence decreasing the change of signal. In all, to activate the cleavage function for Cas12a-crRNA in vitro application, it is important to dilute to an ideal concentration for operation.

Other than the surface density, the length of the immobilized ssDNA reporter

was also evaluated. We hypothesized that ssDNA reporters with different lengths

might lead to different cleavage efficiency due to the exposed length difference.

Different lengths of ssDNA probes at the same concentration were evaluated based on

the same reaction condition of E-CRISPR as investigated previously. Moreover, the

effect of passivation agents with different carbon chain lengths may influence the

electrostatic interaction between the phosphate backbones of ssDNA probes, therefore

was also evaluated for optimized cleavage activity (Figure 3B). The selected ssDNA and

passivation agent pairs were then compared through the effect of lengths on trans- cleavage activity. However, different lengths of ssDNA reporter only produce a minute variation (< 5%) of signal change (Figure 2H). We observed that for a short ssDNA reporter (10 nt), the electrochemical oxidation current over the background current was larger than that of long ssDNA strand because of its short contact mediated electron tunneling distance to the electrode resulting a faster charge transfer kinetics (Figure 2E).

Therefore, the short probe possessed a large baseline current. As for long reporters (20 nt & 30 nt), they gave a relative low background current (Figure 2F&G), but the ΔI% of these long reporters were comparable to that of short reporter. 20 nt ssDNA reporter was selected for further application because of its relative greater degree of signal change and smaller standard error (Figure 2H). Moreover, evaluation of a multi- component monolayer system (e.g. ternary self-assembled monolayers) might also be a potential solution for future researches seeking for a higher sensitivity through tuning the surface molecular packing condition.93-95

Another interesting finding regarding to the cleavage accessibility is that Cas12a-

crRNA based trans-cleavage activity is also significantly concentration dependent as is

its analog Cas9.96 Different concentrations of Cas12a-crRNA in response to a same target

concentration were evaluated (Figure 3C). A high concentration level (>100 nM) in a 30 uL sample solution significantly decreased the activity of the Cas12a nuclease to

nonspecific ssDNA reporter, due to the large size of the Cas12a probably caused a

diffusion hindrance effect in the solution. Hence, a relative minor change of current

outputs was observed based on a high concentration level of Cas12a-crRNA. An optimized concentration for Cas12a-crRNA duplex trans-cleavage operation was identified to be 30 nM in a 30 uL solution.

Figure 4. E-CRISPR analysis of HPV-16. A) Dose-response curve of the detection of HPV-16 in different matrixes (green line- 10 mM Tris buffer containing 50 mM NaCl and 15 mM MgCl2; purple line- 100% human serum). B) Selectivity study through

comparison of the signal changes based on non-target nucleic acids (500 nM) with that of 1 nM of HPV-16 (n=3, *P<0.01, target signal vs. non-target signal). C) Target strands with mismatches at different positions, including PAM region and crRNA complement at different positions: 1, 6, 11, 16. D) Evaluation of the influence of mismatches at different positions on the E-CRISPR signal. A target concentration of 1 nM was applied for all the targets (wild type (WT) and mismatched targets).

2.1.4 E-CRISPR on Nucleic Acid Detection

Based on the optimized trans-cleavage condition, we evaluated the E-CRISPR platform on the detection of HPV-16. A broad dynamic range (pM to µM) of more than

three orders of magnitude was achieved with a IC50 value of 0.78 nM based on the

samples prepared in the buffer solution (Figure 4A). The dose-dependent response curve

demonstrated an average standard error (SE) of 2.16% (n=3), indicating a reliable

reproducibility. An experimental limit of detection (LOD) at 50 pM was obtained. Worth

mentioning, this LOD of optimized trans-cleavage activity based E-CRISPR surpassed previously demonstrated LOD for non-enzymatic amplified nucleic acid detection over two orders of magnitude.77 Moreover, the detection performance in complex matrix was

also evaluated. A IC50 value in pooled human serum was 0.68 nM, which was

comparable with the IC50 value (0.78 nM) in buffer solution, indicating a great potential

of E-CRISPR in direct analyzing of biological sample.

We further investigated the accuracy of the E-CRISPR platform. A scrambled

sequence and PB-19 were applied to evaluate the selectivity for HPV-16 detection. 500

nM of scrambled sequence and PB-19 sequence demonstrated a signal change less than

1.5% and 1.7%, which were lower than the standard error of the signal generated by 1 nM

HPV-16 target, indicating a good selectivity of the Cas12a-crRNA duplex on

differentiating HPV-16 from non-target (Figure 4B).

Furthermore, as a biosensing platform, discrimination of mismatches in the

nucleic acid base pairs is of especially importance for the potential application for the

identification of disease related point mutations.97, 98 Thus, we next challenged the E-

CRISPR with artificial mismatched nucleic acid targets (HPV-16). The recognition mechanism of CRISPR-Cas12a involves the identification of PAM region on the target to unwind the DNA target by Cas protein and further hybridization between the crRNA and the target strand.76, 92 Therefore, we designed the mismatches at different positions

on the target (Figure 4C). E-CRISPR signal was obtained based on the detection of 1 nM

of these artificial targets (Figure 4D). Comparing with the wild type (WT) HPV-16

sequence, mutations in PAM region and PAM-adjacent region (position 1) led to

complete diminishment of the trans-cleavage signal. This phenomenon indicates the mandatory requirement of PAM sequence for the Cas12a-crRNA duplex to recognize and cleave the target.99, 100 Moreover, mismatches in the complementary region of crRNA and target demonstrated retarded trans-cleavage activity, consistent with previous mismatch

tolerance study of Cas12a.76, 101 The clear differentiable SWV signal between mismatches

at different positions also suggests that the trans-cleavage activity of Cas12a can be

utilized as a biosensing strategy to identify the position of the mismatched base pairs.

Overall, the developed E-CRISPR demonstrates a sensitive, generalized and cost-effective

platform for nucleic acid analysis.

2.1.5 Aptamer based E-CRISPR Cascade for Protein Detection

We next explored whether the E-CRISPR could be repurposed as a protein detection platform by utilizing the nucleic acid detection capability of E-CRISPR. For

protein detection, ssDNA aptamer was used as the recognition element for a protein of

interest. An aptamer based E-CRISPR cascade was designed for protein detection

(Figure 5), allowing the direct analysis of complex sample without any time-consuming

processing procedures. A fixed concentration of aptamer is firstly applied to treat the

sample directly (Figure 5A). Cas12a-crRNA is designed to specifically recognize the

aptamer. The E-CRISPR is then applied to determine the remaining concentration of

aptamer in the sample (Figure 5B). With the presence of the protein target, less aptamer

is captured and transduced by E-CRISPR, leading to a high electrochemical signal of the

methylene blue from the ssDNA reporter. In the absence of the protein target, the

electrochemical signal is lower due to the activation of trans-cleavage activity by the

target recognition (Figure 5C).

Figure 5. Illustration of E-CRISPR for protein detection. A) Sample containing protein target of interest is firstly treated by a fixed concentration of target specific aptamer (ssDNA). B) A E-CRISPR system is specifically designed for the recognition of the aptamer. The remaining concentration of aptamer is analyzed by E-CRISPR. C) A representation of SWV results based on the with target and without target condition. D) Linear calibration curve of TGF-β1 protein detection with an equation of Y=0.91X+1.79 and R-square value of 0.99 (n=3, SE=1.54%). E) Selectivity study through comparison of the signal outputs based on non-target proteins (10 nM) with that of 10 nM of TGF-β1 (n=3, **P<0.01 versus different interference substances). F) Concentration-dependent signals observed within conditioned medium harvested at two time-points during the chondrogenic differentiation program of human mesenchymal stem cells (hMSCs) containing TGF-β1. The samples were analyzed by three sets of individual experiments using three different sensors (n=3, ***P<0.05, Day 28 vs. Day 2). The horizontal black dashed line represents the average signal variation (n=3) based on the presence of blank conditioned medium.

Figure 6. A) Dose-response curve for the detection of TGF-β1 aptamer (n=3, SE=1.53%). The detection procedure is the same as described for the detection of HPV-16. B) E- CRISPR SWV evaluation of blank cell lysate sample with 50 nM TGF-β1 protein (red line) and without TGF-β1 protein (black line) after treatment with 50 nM TGF-β1 aptamer. C) Comparison of current outputs difference between 50 nM and 1 nM (detection resolution) based on different trans-cleavage period for Cas12a-crRNA-TGF- β1 aptamer. The current change was calculated through the difference between the ΔI% of 50 nM and 1 nM TGF-β1 aptamer. D) Transcriptome data (RNAseq) for TGF-β1 expression at day 2 and day 28.

This designed E-CRISPR array was evaluated for the detection of transforming growth factor beta 1 (TGF-β1) protein, which is a secreted protein contributing to cell proliferation and differentiation,102 and is also recognized as a biomarker for

hepatocellular carcinoma.103, 104 The dose dependent E-CRISPR for the detection of TGF-

β1 aptamer was first evaluated based on the previous established trans-cleavage

condition (Figure 6A). For proof-of-concept, a fixed concentration of aptamer was first

applied to treat sample with and without TGF-β1 protein. E-CRISPR was then applied

to analyze the samples with and without TGF-β1 protein demonstrating a clear signal

difference (Figure 6B). In order to increase the detection resolution for nano molar

concentration range, a greater degree of current difference between 1 nM and 50 nM is

necessary. Therefore, longer trans-cleavage period was investigated to evaluate whether

a higher current difference could be obtained due to that the trans-cleavage activity is a

multiple-turnover reaction.76 Increasing trans-cleavage period indeed led to a higher detection resolution (Figure 6C), so a trans-cleavage period for protein detection was selected to be 60 min. Therefore, this strategy might be applicable to tune the dynamic range and detection limit of the E-CRISPR platform, enhancing the detection performance. An aptamer concentration of 50 nM was selected for protein sample treatment for 30 min. After the treatment, the sample was evaluated by E-CRISPR. A linear detection range was achieved covering three order of magnitudes with an experimental detection limit tested as 0.2 nM (Figure 5D). The detection specificity was investigated using the conditioned medium from hMSCs chonodrogenesis (a complex matrix) biomolecules, including collagen type II, aggrecan protein and bovine serum albumin. The designed strategy defined a good selectivity on target protein over non-

specific molecules, indicating an excellent specificity of the applied aptamer in the

system (Figure 5E). We further challenged the E-CRISPR platform with samples obtained during the chondrogenic differentiation program of hMSCs, which were cultured in aggregates with complete chondrogenic differentiation medium for 4

weeks.105 TGF-β1 protein was produced during the chondrogenic differentiation

process.106, 107 A clear difference was identified between the conditioned medium

obtained at day 2 and day 28 (Figure 5F). These results are in agreement with the

transcriptome analyses performed during Hmsc chondrogenesis of the same analyzed

sample (Figure 6D), indicating a reliable performance of the designed E-CRISPR array

for protein detection. The nucleic acid based receptor is a generalized recognition

element for both protein and small molecule.108 Hence, the designed E-CRISPR array

can also be extended to a wide variety of analytes.

2.1.6 Conclusions for E-CRISPR

This study introduces a new strategy for the development of electrochemical biosensors

by using electrochemistry to probe the CRISPR cleavage activity (E-CRISPR). Owing to

the high-specificity of target recognition, other than gene editing tool, we utilized the

CRISPR Type V system, Cas12a (cpf1) as an efficient biosensing system, which

translates the target recognition activity into a detectable electrochemical signal through

an interrogating electrode constructed with non-specific ssDNA. Various factors were

investigated to produce an optimized on-chip trans-cleavage activity for a high-

sensitivity E-CRISPR detection platform. Moreover, our preliminary implementation

illustrates that the E-CRIPSR system can be applied not only for nucleic acid sensing;

with the addition of an aptamer based sensing cascade, but also be utilized for protein

detection, providing a generalizable, robust and cost-effective detection system.

2.1.7 Experimental procedure for E-CRISPR

2.1.7a Fabrication of ssDNA Reporter surface

An array containing twenty sensors was first cleaned through an established procedure

using potassium hydroxide, sulfuric acid and nitric acid.53 Thiol linked ssDNA reporter was treated with 10 µM of tris(2-carboxyethyl)phosphine (TCEP) to reduce the S-S bond

for 10 min in the dark at room temperature. The ssDNA reporter was then diluted to 1

µM using 10 mM Tris buffer containing 10 mM EDTA. 20 µL of the 1 µM ssDNA

reporter was then directly incubated onto the gold sensor for 1 hr in the dark at room

temperature. The ssDNA immobilized sensor array was then cleaned by immersing in

10 mM Tris buffer for 5 min. After cleaning, the sensor array was immersed in 2 mM 6- mercaptoheaxnol (MCH) prepared in 10 mM Tris buffer for 30 min to passivate the surface and replace loosely tethered ssDNA reporter, forming a highly-aligned surface

(Operation of MCH related steps should be conducted in a fume hood due to its toxicity). After the MCH treatment, the sensor array was then cleaned by immersing in

10 mM Tris buffer for 5 min. The cleaned sensor array was then dried by nitrogen gas

and ready for treatment by CRISPR system. For a short storage period, the cleaned

sensor array was stored in 10 mM Tris buffer (containing 100 mM NaCl) at 4 °C.

2.1.7b In vitro Digestion of Cas12a-crRNA

Cas12a-crRNA duplex was prepared in a buffer prepared by nuclease free water containing 50 mM NaCl, 10 mM Tris-HCl, 15 mM MgCl2, 100 µg/ml BSA with a pH of

7.9. 30 nM of Cas12a-crRNA was assembled and incubated at 25°C for 10 min.

Typically, for nucleic acid detection, 4 µL of sample was added into 26 µL of the

Cas12a-crRNA duplex to form the Cas12a-crRNA-target triplex and incubated for 10 min at room temperature. 20 µL of the Cas12a-crRNA-target triplex solution was applied to ssDNA reporter covered sensor for trans-cleavage activity at 37°C for 30 min.

80 U/mL of Proteinase k was applied to the CRISPR treated surface at 37°C for 15 min before the electrochemical analysis. For protein detection, 10 µL of 100 nM of aptamer was applied to treat 10 µL of sample (resulting in a 50 nM final concentration of aptamer) and incubated at room temperature for 30 min. E-CRISPR as described above was then applied for protein sample analysis with an elongated trans-cleavage period for 60 min.

2.1.7c On-Chip Electrochemical Analysis

After the on-chip CRISPR reaction, the sensors were cleaned by immersing the sensors into a 10 mM Tris buffer for 5 min. For electrochemical test, a 10 mM Tris buffer containing 100 mM NaCl was applied as the electrolyte. Square wave voltammetry

(SWV) was applied before and after the treatment of Cas12a-crRNA-target triplex to obtain the change of current based on a potential range of -0.6V to -0.1V, a frequency of

25 Hz, an amplitude of 25 mV (variation of frequency (15 Hz-120 Hz) and amplitude (25

mV- 50 mV) did not present significant enhancement of the quantity of signal changed

or the signal stability).

2.1.7d Clinical Sample-Mesenchymal stem cell (MSCs) culture and differentiation:

Cultures of human bone marrow-derived MSCs from healthy de-identified adult

volunteer donors were established as previously described. The bone marrow was

collected using a procedure reviewed and approved by the University Hospitals of

Cleveland Institutional Review Board; informed consent was obtained from all de-

identified donors. Cells were expanded in DMEM-LG supplemented with 10% fetal

bovine serum, supplemented with FGF2 (10 ng/ml of) for 14 days. Cells were

trypsinized and then resuspended in chondrogenic differentiation medium consisting

of DMEM-high glucose supplemented with 1% ITS+,10-7 M dexamethasone, 1mM

sodium pyruvate, 120 mM ascorbic acid-2 phosphate, 100 mM nonessential amino

acids, and 10 ng/mL TGF-β1 protein. Two hundred microliters of this cell suspension containing 250,000 cells was added per well of a 96-well polypropylene V-bottom, multi-well dish (Phenix Research). The multi-well plates were centrifuged at 500 g for 5 min and then incubated at 37 °C. The differentiation medium was changed every other day. Conditioned medium from these pellets was collected at different time points.

Days 2 and 28 were chosen to use in the biosensor platform based on previous transcriptome data (RNAseq) showing a greater difference in TGF-β1 protein expression between days 2 and 28 (Figure 6D). To activate the latent secreted TGF-β1 protein to the detectable form, 20 µL of 1 M HCl were added to 100 µL of conditioned medium and incubated for 10 minutes and then neutralized with 20 µL of 1.2 M

NaOH/0.5 M HEPES. The samples were assayed immediately. This procedure ensures that only the secreted version of TGF-β1 protein assayed.

2.2 CRISPR Enhanced E-DNA sensor

2.2.1 introduction

A simple, rapid detection system that can identify the morbigenous genes will be

of especially high-utility for the design of personalized medicine109. Among the

advancements of various types of biosensing platforms, the electrochemical biosensing

systems have been robustly developed in the last decade 55, 58, 110, owing to the cost-

effectiveness of the transduction system, the time-efficiency and the simplicity for sample

processing, providing a potential deployable point-of-care system61, 63, 67, 111-117. One of the

most successful contributions of the electrochemical biosensing system is the

development of the E-DNA biosensor118, which utilizes the conformational change of the

surface probe upon the recognition of the target, inducing the change of electrochemical

signal through the variation of contact-mediated electron transfer distance119-130. The change of electrical signal before and after the introduction of the target (signal gain) depends on the quantity of hairpin surface probes opened by the target nucleic acid.

Various studies have demonstrated the enhancement of signal gain through tuning different factors, such as voltammetric parameters131, 132, surface passivation condition and probe density,133 stem composition of the hairpin probe134, in order to achieve an ideal

conformational change for a satisfied sensing performance. These detail experimental

tunings might limit the general programmability and applicability of the system and

therefore highlights the need for a strategy that uncouples the challenge for an absolute

signal gain and thus, surpassing the detection limit of E-DNA sensor.

From the idealized chemistry perspective, the ultimate signal gain is the elimination of the electrochemical signal in the presence of the target, indicating that the removal of the electrochemical tag upon the target recognition is desirable. Recent discoveries of clustered regularly interspaced short palindromic repeats (CRISPR) Cas systems actually provide the opportunity to produce an absolute signal gain through the programmable and specific RNA guided cleavage activity69, 76, 77, 135, 136. The CRISPR Cas system, as a

powerful gene-editing tool80, 90, performs the high-specificity recognition activity through

the Cas enzyme-gRNA duplex and the cleavage activity on the target DNA through the

Cas enzyme57, 137, therefore providing the occupancy on the genome to allow the cell to operate its own repair mechanism82, 138, achieving genome editing. We hypothesized that through the utilization of the target recognition induced cleavage activity of CRISPR for

E-DNA sensor, the nucleic acid target and its complement (opened hairpin containing the electrochemical tag) can be cleaved by Cas enzyme, therefore releasing the electrochemical tag from the sensor surface. This leads to an absolute loss of the signal, therefore surpassing the detection limit.

Moreover, a biosensing strategy that can discriminate mutation is of great meaning for genetic disorder analysis. However, current electrochemical mutation detection strategies mostly depend on the variation of DNA medicated charge transport process, which can only indicate mutations at a known target concentration condition139,

140. Because without knowing the concentration of the target, the decreased current output

of E-DNA sensor can be explained as either increased well-matched target concentration

or the presence of the mutation, therefore not practical for clinical application for

mutation evaluation. To resolve this issue, we further hypothesized that the

incorporation of target complementarity dependent CRISPR cleavage activity into the E-

DNA sensor allows a second-time recognition activity with a corresponding complementarity dependent cleavage signal, which can be combined with E-DNA sensor

conformational change induced signal to indicate the presence of mutation without

knowing the target concentration, therefore surpassing the detection accuracy and

providing a unique strategy for mutation analysis.

In this work, we introduces a new concept based on the combination of

electrochemistry based DNA sensor and CRISPR, delivering an integrated sensor and

actuator system as a powerful biosensing platform. The designed CRISPR enhanced E-

DNA sensor was demonstrated for the detection of a ssDNA virus, Parvovirus B19 (PB-

19), which is known to cause erythema infectiosum in children and pregnant women141.

We believe that owing to the generality and the continuous discoveries of the CRISPR

and its analogs, our demonstrations of CRISPR systems on E-DNA sensor can be widely

adopted for future researches to boost the detection limit and accuracy for all the types

of point-of-care systems.

The principle of the sensing chemistry is illustrated in Figure 7. A hairpin probe is

attached to the electrode surface through 5’ end modified thiol moiety physically

absorbed onto gold surface with a 3’end modified methylene blue electrochemical

signaling tag (Figure 7A). The presence of the target ssDNA opens the hairpin, delivering a longer electron transfer distance of the electrochemical tag (Figure 7B). At this stage, a differentiable electrochemical signal is already acquirable before

Figure 7. Principle of CRISPR enhanced E-DNA sensor. (A) Hairpin signaling strand construction on the gold electrode. (B) Addition of target opens the hairpin probe, increasing the electron tunneling distance. (C) Addition of CRISPR-Cas system performs a target dependent cleavage activity, further releasing the electrochemical signaling probe from the electrode surface.( D) Representation of electrochemical current outputs based on conformation change induced signal decrease and CRISPR cleavage induced signal decrease.

and after the introduction of the target nucleic acid to the sensor. For the purpose to enhance the signal gain, RNA guided CRISPR system can be programmed to recognize

the formed dsDNA. Upon the second-time confirmation of the sequence information by

CRISPR, cleavage activity is activated to remove the electrochemical tag conjugated DNA probe off the electrode (Figure 7C), further enhancing the signal gain (Figure 7D).

Moreover, the CRISPR cleavage activity depends on the sequence complementarity with

the guide RNA, so probing the CRISPR cleavage signal can lead to the indication of

mutations in the target sequence, surpassing the detection accuracy of the conventional

electrochemical DNA sensor.

2.2.2 Optimization of the E-DNA Sensor

Before applying the CRISPR system onto the E-DNA sensor, we first optimized the detection performance of the E-DNA sensor through manipulating the surface density of the hairpin probe, the passivation agent, and the parameter of square wave voltammetry

(SWV). SWV was applied to evaluate the methylene blue signal change before and after the incubation of the target ssDNA due to the change in the rate of contact-mediated

electron transfer (Figure 8A)127. The hairpin surface density was manipulated through the change of the concentration of the incubation solution; the greatest degree of change of signal was identified at the incubation concentration of 0.8 µM (Figure 8B). We further examined the effect of passivation agents with different lengths on the enhancement of the signal gain. 11-mercapto-1-undecanol (11-MCD) demonstrated the most stable and the greatest degree of current change before and after the introduction of the target

(Figure 8C). This might cause by the following reason. The hairpin immobilization linker was a 6-carbon chain. A longer carbon chain from 11-MCD can provide an intense electrostatic interaction from the hydroxyl group (of the passivation agent) to interact with the phosphate backbone of the hairpin molecule (Figure 8D). Therefore, the long passivation agent sustained the unfolding process of the hairpin probe more favorably than the short passivation agents, providing a stable, extended duplex structure with the introduction of the target and further delivering a greater change of electron transfer distance (Figure 8D). Thus, 11-MCD was selected as the passivation agent for the

development of the E-DNA sensor. The optimized E-DNA sensor was then applied to

CRISPR treatment for performance enhancement.

Figure 8. A) SWV evaluation of methylene blue signal before and after the introduction of the ssDNA probe onto the E-DNA sensor. This proof-of-concept result was obtained by the following procedure. A concentration of 1.6 µM of hairpin probe was incubated for 3 hr at room temperature in the dark. 2 mM 6-mercaptohexanol was incubated for overnight at 4°C to passivate the surface. A 500 pM of the ssDNA target was incubated for 30 min for SWV evaluation with a frequency of 25 mV and an amplitude of 0.025 V. B) Evaluation of incubation solution ( ) concentration for the optimized change of current. % = . Ibaseline is the 𝐼𝐼𝑏𝑏𝑏𝑏𝑠𝑠𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒𝑒−𝐼𝐼𝑡𝑡𝑡𝑡𝑡𝑡𝑡𝑡𝑡𝑡𝑡𝑡 SWV current before the introduction of target. Itarget is the SWV current after the introduction of ∆𝐼𝐼 𝐼𝐼𝑏𝑏𝑏𝑏𝑏𝑏𝑏𝑏𝑏𝑏𝑏𝑏𝑏𝑏𝑏𝑏 target. The experimental conditions of this evaluation is based on the sensor surface passivated by 2 mM 6-MCH. 0.8 µM hairpin solution incubation for 3 hr with passivation produced the greatest degree of current change based on 500 pM of ssDNA target (n=3). Evaluation of passivation agents with different lengths. C) 3-mercaptopropanol (3-MCP), 6-mercaptoheaxnol (6-MCH), and 11-mercapto-1-undecanol (11-MCD) were evaluated as passivation agents. After incubation of the hairpin probe onto the gold electrode, 2 mM of passivation agents were applied for overnight at 4°C. 500 pM ssDNA target was used to evaluate the change of current (n=3). A significant performance enhancement (over 70%) was observed on 11-MCD using as the

passivation agent comparing with those of 3-MCP and 6-MCH. D) Hypothesized mechanism of the enhancement of signal owing to the longer carbon chain of the passivation agent. Passivation agent with longer carbon chain would interact with the opened duplex DNA more intensively, leading to a longer electron transfer distance based on contact-mediated electron transfer of methylene blue tag. However, passivation agent with short carbon chain would loosely confine the steric movement of the duplex DNA, enhancing the electron transfer rate of the contact- mediated electron transfer process. Therefore, a longer chain leads a more stable formation of the extended, hybridized duplex DNA and greater degree of change of electrochemical signal before and after the introduction of the target ssDNA.

2.2.3 Evaluation of CRISPR Cas9 for the E-DNA Sensor Performance

Enhancement

We first evaluated our designed detection strategy through the classic CRISPR-Cas9

system (from Streptococcus pyogenes), which is a type II CRISPR system142. Cas9 is a dual-

RNA guided DNA endonuclease143, consisting of three elements (Figure 9A): a crRNA: guiding for DNA cleavage, a tracrRNA: non-coding RNA to process the crRNA into the discrete ternary Cas9 complex, a Cas9 enzyme: cleaving target through the HNH domain and RuvC like domain143. The sensing/recognition part of CRISPR-Cas9 system is

achieved by two parts: first, the confirmation of a short motif sequence referenced as the

protospacer adjacent motif (PAM), 5’ NGG; second, the programmable crRNA guided

target sequence recognizes through the complementary base pairs. The integration of

crRNA and tracrRNA is named as gRNA duplex. The actuator part is processed by gRNA

guided Cas9 enzyme to perform the dsDNA cleavage at the position of 3 base-pair upstream of 5’ PAM region and the cleavage activity is probed through electrochemistry

(Figure 9B). Based on the optimized E-DNA sensor condition, SWV was applied to

evaluate the change of current outputs before and after the treatment by CRISPR-Cas9

(Figure 9C). In response to 500 pM ssDNA target, CRISPR enhanced E-DNA sensor

(ΔIb%=59.1%) demonstrated an increasing change of current over the conventional E-

DNA sensor (ΔIa%=49.7%), indicating the effectiveness of the cleavage activity of Cas9

for on-chip in vitro cleavage. We further compared the sensing performance at low

concentration range using E-DNA sensor and Cas9 enhanced E-DNA sensor (Figure 9D).

CRISPR-Cas9 enhanced E-DNA sensor demonstrated a higher detection resolution over

the E-DNA sensor at pico-molar concentration level. Moreover, CRISPR-Cas9 enhanced

E-DNA sensor surpassed the detection limit of the E-DNA sensor by at least two order of magnitudes. Differentiable signal was even acquired between 1000 fM and 100 fM. The dynamic detection range of CRISPR enhanced E-DNA sensor was then examined (Figure

9E). A IC50 value of 66.1 pM was obtained with an average standard error (SE) of 3.2%, indicating an excellent stability and sensitivity of this sensing platform.

Figure. 9. Evaluation of the sensing performance of CRISPR-Cas9 enhanced E-DNA sensor. (A) CRISPR Cas9 system consists of Cas9 protein and tracerRNA and crRNA triplex. The combination of tracerRNA and crRNA is named guideRNA. Recognition of PAM sequence by Cas9 leads to crRNA guided dsDNA cleavage activity. (B) Workflow of CRISPR-Cas9 enhanced E-DNA sensor. (C) SWV (in response to 500 pM of target) comparison of baseline without the target (black), addition of ssDNA target (red) and CRISPR Cas9 enhanced result (blue). (D) ΔI%E-DNA= (I baseline– I addition of target)/I baseline. ΔI%CRISPR enhanced = (I baseline– I CRISPR enhanced)/I baseline. Comparison of detection limit between E-DNA sensor and CRISPR-Cas9 enhanced E-DNA sensor. (E) Dose-dependent response of CRISPR-Cas9 enhanced E-DNA sensor (SE=3.2%). The data represent the mean ± S.E. of three replicates based on three individual sensors.

2.2.4 Comparison of the performance of CRISPR Cas12a and Cas9

We further explored a different class of CRISPR, type V CRISPR-Cas system, Cas12a

to evaluate its effect on enhancing the performance of the E-DNA sensor. Distinct from

CRISPR-Cas9, Cas12a performs dsDNA cleavage with a single RuvC catalytic domain

guided by its own crRNA, providing a simpler system than Cas976. The PAM recognition

motif for Cas12a depends on 5’ TTTN, which is different from that of the Cas9. Most

importantly, upon the recognition of the complementary target through PAM

confirmation and crRNA matching, Cas12a not only performs the cleavage activity on the

target DNA (cis-cleavage), but also unleashes an indiscriminate single-stranded

deoxyribonuclease activity (trans-cleavage) (Figure 10A). This unique evolved property

enhances the breakdown of the invading nucleic acids as an important part of the

bacterial adaptive defense system of type V CRISPR enzyme77, 79. Therefore, we assumed

that through the application of Cas12a’s cis and trans cleavage activities onto the E-DNA sensor, the detection limit and the dynamic detection range of E-DNA sensor can further be advanced.

Cas12a from Lachnospiraceae bacterium (LbCas12a) was selected for investigation owing to its superior activity over the other Cas12a analogs.57 Same experimental setting

was applied for CRISPR-Cas12a enhanced E-DNA sensor as for CRISPR-Cas9 system

(Figure 10B). SWV demonstrated a more significant enhancement of signal gain

Figure 10. Performance of CRISPR Cas12a enhanced E-DNA sensor. (A) CRISPR Cas12a- crRNA (yellow) duplex unit recognizes the PAM region, unwinds the dsDNA and binds the complementary strand (green). Recognition of dsDNA target activates its cis-cleavage activity, generating a 5’-overhang staggered cleavage. Activation of cis-cleavage activity further induces the trans-cleavage activity on non-specific DNA (blue). (B) Workflow of CRISPR-Cas12a enhanced E-DNA sensor. (C) SWV (in response to 500 pM target) comparison of baseline without the target (black), addition of ssDNA target (red), CRISPR Cas9 enhanced result (blue) and CRISPR Cas12a enhanced result (green). (D) ΔI%= (I addition of target – I CRISPR-Cas12a enhanced)/I addition of target. Enhancement of detection signal was observed with increasing period of the Cas12a digestion process. (E) Dose dependent responses based on the detection of ssDNA target in buffer and in pooled human serum (ΔI%CRISPR Cas12a enhanced = (I baseline– I CRISPR Cas12a enhanced)/I baseline, SE=2.73%) The data represent the mean ± S.E. of three replicates based on three individual sensors.

by Cas12a system (ΔIc%=67.5%) comparing with that of Cas9 (ΔIb%=59.1%), indicating the trans-cleavage activity of Cas12a mediated the further decrease of the signal (Figure

10C). Another interesting finding is that the trans-cleavage activity was a continuous

event after the activation of the cis-cleavage activity. The cis-cleavage activity is typically

finished within 30 min76, 77. However, our observation demonstrated that with the application of Cas12a, the current change kept increasing over 1 hr (Figure 10D). This phenomenon suggests that the elongation of the CRISPR-Cas12a digestion period is a

potential mean to tune the sensitivity and resolution for sensing applications. We further

evaluated the dose-dependent response of the CRISPR-Cas12a enhanced E-DNA sensor

for the detection range and the detection limit (Figure 10E). An excellent dynamic range

was found to cross 7 orders of magnitude. A saturated detection concentration was

reached at 0.1 µM and an experimental detection limit was identified as 10 fM for both

buffer and pooled human serum tests with an average standard error of 2.8%, indicating

an excellent and stable detection performance. The comparable IC50 values for buffer

(31.9 pM) and pooled human serum test (51.6 pM) were acquired, indicating a potential

operation of the developed CRISPR enhanced E-DNA sensor directly in complex

biological fluids. The low matrix effect observed may attribute to the CRISPR treatment

process. For CRISPR enhancement procedure, it is necessary to use proteinase K, a serine protease144, to treat the electrode in order to digest the Cas protein, retaining an electrode

surface with nucleic acids only. Owing to the broad specificity of proteinase K as a

proteinase, it would also digest those non-specifically, electrode absorbed proteins from

human serum, therefore providing a comparable detection performance in complex

matrix as in buffer. This observation also suggests that for general affinity based

electrochemical biosensor development, a post-treatment through proteinase K could

improve the detection performance on human biological samples.

Figure 11. Sequence dependent CRISPR cleavage activity for discrimination of point mutations. (A) Different target sequences with mutations at different positions induced current change based on the conformational change of hairpin probe of the E-DNA sensor. (B) Different target sequences with mutations at different positions induced current change based on the CRISPR cleavage activity after formation the target-hairpin dsDNA (n=3, *P<0.05, WT vs. MT1&MT2) (C) Correlation factor is based on E-DNA signal over CRISPR cleavage signal (Correlation factor=(ΔI%E-DNA)/ (ΔI%CRISPR enhanced)). The data for correlation factor was acquired based on triplet repeat of 5 different concentrations of wild type sequence or single mutation sequence testing on the E-DNA signal and CRISPR enhanced signal (n=3, *P<0.05, Single mutation vs. Wild type). The data represent the mean ± S.E. of three replicates based on three individual sensors. 2.2.5 Probing the mismatches with CRISPR enhanced DNA sensor

It would be of extremely high utility for a nucleic acid sensing system to be able to

discriminate point mutation as a potential lab applicable technique for genotype

screening. We firstly evaluated the mismatch tolerance capability of E-DNA sensor based

on only target induced conformational change. We artificially designed mutations at

different locations of the target sequence. 500 pM of each designed target was applied to

the E-DNA sensor and the signal gains were compared without the application of CRISPR

systems (Figure 11A). Same as wild type target, point mutations also led to the increase

of current change of the E-DNA sensor because of the following reasons: 1. hairpin probe

would still undergo conformational change because single mismatch is not able to

significantly impede the nucleic acid hybridization process due to the similar melting

temperatures of the hybridized dsDNAs (Table 1)139. 2. After hybridization, due to the

perturbation of the DNA base stacking by the mismatched base pair, DNA-mediated charge transduction is interfered66, 145, resulting in a decrease of electrochemical current

and hence a big % change before and after hybridization onto the E-DNA sensor. Actually,

the variation of pi-stacking conformation resulted unfavorable electron transfer through

DNA is utilized to identify point mutations in electrochemical biosensing systems66, 140, 146.

However, these assays can only compare the electrochemistry of the mismatched strand

with the well-matched strand based on a known target concentration condition. Because,

at the condition of unknown target concentration, either high concentration of wild type

target or mutated target can result a decrease of current output (Figure 12). Thus, a

general method, which is able to confirm that the decrease of current is created by the

mutation without knowing the concentration, which can be extremely meaningful for all

types of nucleic acid sensors aiming to discriminate mutations.

Table 1. Melting temperatures of dsDNAs based on mismatched positions at 500 pM.

Figure 12. Decrease of current output based on the E-DNA sensor stem loop can be resulted by increased of correct target concentration or the presence of a mismatched sequence. A) The opened stem loop by a well-matched target changes the distance of MB tag to the electrode surface, therefore decreasing the current output. Higher the concentration, more the change of current before and after introduction of the correct target. B) Mismatched target can also open the stem loop and caused a decrease of the current output based on the increased of the distance of MB tag to the electrode surface. Furthermore, the electron transduction is perturbed by the damaged base stacking, leading to a further decrease of current output. Both above described condition can lead to decrease of the current. Therefore, without knowing the concentration of

the target, it is not possible to identify the presence of point mutation. C) Representation of the described conditions resulted current outputs.

Owing to the target complementarity based CRISPR cleavage activity, we

hypothesized that the acquisition of CRISPR cleavage signal may provide a mean to

identify the mutation in the sequence. After the targets hybridized on the E-DNA sensors,

we examined the change of current (ΔI%CRISPR enhanced= (I addition of target – I CRISPR-Cas12a

enhanced)/I addition of target) caused by CRISPR cleavage activity based on different target

sequences (Figure 11B). Mutation 1 (MT1) and mutation 2 (MT2) have mismatched base

pairs at PAM region and the PAM adjacent region corresponding to the crRNA of CRISPR

Cas12a, which demonstrated significant attenuation of CRISPR cleavage activity. The

demonstrated cleavage activity of mutated targets also matches previous presentation on the mismatched tolerance study of Cas12a101. This phenomenon is explained based on the

CRISPR recognition dependent cleavage mechanism as described following: CRISPR

Cas12a-crRNA duplex recognizes corresponding DNA target through a two-step process92. First, the Cas12a endonuclease searches for its corresponding PAM sequence

(5’-TTTN). Upon the confirmation of the PAM sequence, the Cas12a endonuclease would

unwind the double strand DNA. The second step involves the programmable crRNA to

examine the complementarity between itself and the target. Once the target

complementarity is confirmed, the RuvC domain of CRISPR performs the cleavage

activity. Therefore, mutations at PAM or PAM adjacent base pairs can avoid the CRISPR

cleavage activity on the wrong sequence, providing an effective mean to differentiate

mutated sequences from the wild type sequence.

Based on this mutation dependent cleavage performance, we correlated the signal

change of CRISPR cleavage activity with the signal change caused by E-DNA sensor.

Utilizing this numerical correlation, we anticipated to indicate whether a mutated target sequence or a well-matched target sequence would induce the electrical signal change, presenting a generalized method to examine the credibility of the detection result. The signal change based on E-DNA conformational change (ΔI%E-DNA) and the signal change by CRISPR cleavage (ΔI%CRISPR enhanced) of various concentrations (pM-nM) of the wild type target were firstly characterized. A correlation factor based on (ΔI%E-

DNA)/ (ΔI%CRISPR enhanced) was then obtained with an average factor of 0.903 (green box chart, Figure 11C). Same characterization process was also applied to different concentrations of mutation sequences (MT1 and MT2). The correlation factor was obtained with an average value of 11.731 (purple box chart). The significant difference of the correlated factors between the wild type target and targets with single mutations demonstrated the reliability of this method to examine the credibility of the detection result, enhancing the detection accuracy. This strategy simply involves CRISPR to perform a complementarity dependent cleavage activity as a second-time target recognition, which can be generalized to various hybridization based nucleic acid detection platform.

2.2.6 Conclusion

In conclusion, we demonstrate a CRISPR enhanced nucleic acid detection strategy

based on the E-DNA biosensing platform. Our approach exploits the specificity and the

target complementarity dependent CRISPR enzymatic activity, surpassing the detection

limit and most importantly, detection accuracy of the conventional electrochemical DNA

sensors. Integration and correlation of DNA sensor capturing signal and CRISPR

cleavage signal led to a dose-independent indication of the potential single-mutation in

the target sequence, providing a reliable and universal method to examine the detection

accuracy. Overall, these novel insights could be generalized to any type of biosensing

platform, promoting the development of biosensors one step forward toward robust

point-of-care systems and clinical applicable genotype screenings.

2.2.7 Experimental procedures

2.2.7a Fabrication of E-DNA sensor

An array of 20 sensors were cleaned through an established cleaning protocol.53 5’

modified C6 S-S linked synthetic hairpin probe was treated with 10 µM of tris(2-

carboxyethyl)phosphine (TCEP) to reduce the S-S bond for 30 min in the dark at room

temperature. After treatment, the thiol linked hairpin probe was diluted to 0.8 µM

using 10 mM Tris buffer containing 100 mM NaCl, 10 mM MgCl2 and 10 mM EDTA.

The diluted hairpin probe was directly incubated onto the sensor array for 3 hr at room

temperature in the dark. Upon completion of the hairpin monolayer formation, the

electrodes were passivated with 2 mM of 11-mercapto-1-undecanol (11-MCD) overnight at 4°C. The electrodes were immersed and rinsed by 10 mM Tris buffer to remove any non-absorbed residual molecules before applying to target incubation. 20 µL of different concentrations of ssDNA targets were incubated onto the sensor array for 20 min before CRISPR treatment.

2.2.7b CRISPR treatment

Assembly of CRISPR Cas9-gRNA duplex.

100 nM of Recombinant S. pyogenes Cas9 and sgRNA was assembled in 50 mM

NaCl, 10 mM Tris-HCl, 10 mM MgCl2, 100 µg/ml BSA (pH 7.9) at ambient temperature

for 10 min. After formation of the CRISPR Cas9-gRNA duplex, 20 µL of the 100 nM of

Cas9-gRNA duplex was applied directly onto the E-DNA sensor (after target capture)

for incubation of 20 min at 37°C to perform the cleavage activity. After CRISPR Cas9

system treatment, the sensor array was cleaned by immersing in 10 mM Tris buffer and

dried by nitrogen gas. 80 U/mL of proteinase k was applied to the sensor array to digest

the Cas enzymes for 10 min at 37°C, preventing non-specific absorptions on the electrodes. After proteinase K treatment, the sensor array was cleaned by immersing in

10 mM Tris buffer and dried by nitrogen gas and ready for electrochemical test.

Assembly of CRISPR Cas12a-crRNA duplex

100 nM of LbCas12a and crRNA was assembled in 50 mM NaCl, 10 mM Tris-HCl,

10 mM MgCl2, 100 µg/ml BSA (pH 7.9) at ambient temperature for 10 min. 20 µL of the

100 nM Cas12a-crRNA duplex was incubated onto the E-DNA sensor (after target

capture) for incubation of 20 min at 37°C to perform the cleavage activity. After CRISPR

Cas12a system treatment, the sensor array was cleaned by immersing in 10 mM Tris

buffer and dried by nitrogen gas. 80 U/mL of proteinase k was applied to the sensor

array to digest the Cas enzymes for 10 min at 37°C, preventing non-specific absorptions

on the electrodes. After proteinase K treatment, the sensor array was cleaned by

immersing in 10 mM Tris buffer and dried by nitrogen gas and ready for

electrochemical test.

2.2.7c Square Wave Voltammetry (SWV) Investigation

For SWV test, a 10 mM Tris buffer containing 1 M NaCl was applied as the

electrolyte. Square wave voltammetry (SWV) was applied before and after certain

treatments to obtain the change of current based on a potential range of -0.1V to -0.3V

(vs. Ag/AgCl) with a frequency of 5 Hz and an amplitude of 25 mV.

Chapter 3 Protein detection

3.1 Enhancing the simplicity of biosensor fabrication through

bioconjugation chemistry

For electrochemical biosensor fabrication, the preparation process is typically complex.

The process involved the formation of self-assembled monolayers (SAM) for

immobilization and functionalization by thiol group at one end and carboxylic group at

the other end, crosslinking and activation treatment for binding of antibody.9, 147 These

processes typically required a couple of days. The completeness of sensor coverage and

overall sensitivity were always the technical concerns in the common preparation

process, such as pin holes for electrode coverage, the alignment of antibody and

others.10 Therefore, a different and effective preparation process for a simple, cost- effective single-use biosensor fabrication was desirable and firstly evaluated. Instead of modification of sensor surface, bioconjugation chemistry was applied to externally modify the recognition element for the immobilization process. Bioconjugation is a

chemical strategy forming a stable covalent link between biomolecule and organic

molecule. This formation results in a zero length linkage between protein and electrode

for this study. This innovative bioconjugation technique was used for the preparation of

biosensor with the advantages on shortening the preparation process, enhancing the

coverage of sensor surface, minimizing pinhole effect, and improving the practical

usage for clinical application. Antibody and antigen interaction remains to be the

crucial mechanism for ensuring the selectivity of the biosensor. N-succinimidyl S-

acetylthioacetate (SATA) was typically used in preparation of antibody-enzyme

conjugates.53 In this study, SATA was selected to conjugate monoclonal anti-TAR DNA- binding protein, forming a SATA-acetylated antibody (Figure 13). Through the reaction between the SATA-acetylated antibody and hydroxylamine, a thiol labeled mono-clonal anti-TAR DNA-binding protein was produced and then applied onto the gold sensor forming the gold-sulfur (Au–S) bond. Comparing with complex multiple-day preparation process of most biosensors on detection biomarkers, this unique procedure was a single-step process, generating a valid biosensor less than 2 hr. TAR DNA binding protein 43 was applied as a target for evaluation of this biosensor fabrication method.

Figure 13. Scheme of bioconjugation process of antibody.

3.1.1 Evaluation of the effectiveness of bioconjugation method.

In order to examine the efficiency and prove the principle of sulfhydryl modification on

protein, a protocol to monitor the extent of modification by NMR spectroscopy was

developed using the N-terminal fragment of the hnRNP F protein (hnRNP F quasi-RNA

Recognition Mofits 1 and 2, FqRRM12, residues 1-194) as a model system and shown in

Figure 14a.Since Anti-TDP-43 purchased from Sigma Aldrich is IgG type, and its

molecular weight was about 150 kDa, which is beyond the characterization limit by

NMR148. Thus, we introduced the FqRRM12 protein for NMR analysis in order to prove

the principle of the reaction. The 15N-1H Heteronuclear Single Quantum Coherence (15N-

1H HSQC) spectrum of a 15N-labeled protein reports on the chemical environment of each

amide group within a protein, and as such it provides a convenient analytical tool to

monitor post-translational modifications149. Sulfhydryl modification is a classical type of

lysine acetylation; therefore, we followed the reaction of FqRRM12 with SATA by

comparing HSQC spectra of the unmodified and modified 15N-labeled protein. Figure

14b shows the overlay of the 15N-1H HSQC spectra of unmodified and modified FqRRM12.

Comparison of the spectra reveal that the correlation peaks of Lys72, Lys124 and Lys185

completely disappeared in the SATA-modified protein. Interestingly, the signal intensity

of Gly133 and Asp74 were also significantly reduced in the 15N-1H HSQC spectrum as

shown in Figure2b. Analysis of the FqRRM12 three-dimensional structure (PDB ID: 2kfy

for RRM1 domain; 2kg0 for RRM2 domain) showed that Gly133 and Asp74 were in close

proximity to the side chains of Lys171 and Lys72, respectively as shown in Figure2a,

which might account for the observed perturbations to their correlation peaks. Therefore,

SATA modification of the epsilon amine groups of lysine residues resulted in the

formation of new amide bonds that were detectable by NMR.As a matter of fact, the 15N-

1H HSQC spectrum of SATA modified FqRRM12 showed additional correlation peaks

relative to the unmodified protein as shown in Figure 14b. The second step of the

chemical reaction, which reduced the attached SATA to a thiol functional group, was also

monitored by recording 15N-1H HSQC experiments. Comparison of the spectra of the

oxidized and reduced forms of the SATA modified FqRRM12 revealed that the reduction

occurred site-specifically since the NMR chemical shifts of the two proteins were

essentially identical as shown in Figure 14c. The absence of additional chemical shift

perturbations to FqRRM12 upon reducing SATA is expected given that the site of reaction

was more than four bonds away from the amide group.

Electrochemical impedance spectroscopy (EIS) was conducted to confirm the

existence of sulfhydryl modified FqRRM12 by examining the ability of the modified

protein to bind with the gold electrode surface. The gold binding ability of unmodified protein was also evaluated as a negative control. A relative large concentration of

FqRRM12 protein (15µg/mL) with incubation of 1 hour on the gold sensor was used to evaluate the binding effect. Significant impedance difference on sulfhydryl modified

protein immobilized sensor was observed comparing with the impedance on the bare electrode using the same redox solution. In Figure 14d, the blue line represents the surface impedance from the sulfhydryl modified FqRRM12 produced electrode and the black line represents the surface impedance from a cleaned bare electrode. Unmodified protein immobilized sensor shows minor change of impedance comparing with that of bare electrode, confirming the gold binding ability of sulfhydryl modified protein. Therefore, the combined NMR and EIS results confirmed that FqRRM12 was site-specifically modified under our experimental conditions.

Figure 14. (a) Possible locations of amino acids modified by SATA on the 3D structure of FqRRM12.). (b) Overlay 1H-15N HSQC spectrum of FqRRM12 (red) and acetylated FqRRM12 (green). Chemical shift perturbations to D74 and G133 were also detected upon SATA modification, likely due to their close proximity to K171 and K72. (c)

Overlay 1H-15N HSQC spectrum of SATA-acetylated FqRRM12 (green) and SATA modified FqRRM12 with reduced thiol groups (blue) (d) Nyquist plot presents the impedance differences between cleaned bare electrode, unmodified protein bonded electrode and sulfhydryl modified protein bonded electrode. 3.1.2 Surface analysis of Au-S formation

In order to confirm the effectiveness of the SATA and protein reaction for the practice

of the biosensor fabrication, the validity of gold sulfur (Au-S) bond and the coverage of

the working electrode surface were investigated by TOF-SIMS technique. The working

electrode with linked anti-TAR DNA binding protein 43 (anti-TDP43) were used for

analysis. Two biosensors prepared separately by thiol linked anti-TDP43 and traditional

11-Mercaptoundecanoic acid (11-MUA) formed monolayer with cross-linked anti-

TDP43 were analyzed and the electrode coverage of each sample was compared. The

confirmation of Au-S bond on the gold electrode surface indicated the successful

synthesis of thiol-linked protein as shown in Figure 15a. Figure 15b shows the coverage

of Au-S bond of the electrode using 11-MUA prepared monolayer with cross-linked

antibody. The difference in coverage between the two samples was apparent. The

counts percentage of Au-S bond based on total counts in the developed thiol-linked anti-TDP43 electrode was 63.3% higher comparing to that of the 11-MUA linked antibody covered electrode. The results of TOF-SIMS analysis proved the effectiveness of bioconjugation mechanism’s ability on covering electrode surface, ensuring the reproducibility of further antigen incubation and detection.

Figure 15. (a)The total secondary ions acquired at the negative polarity of gold (left) and gold-sulfur ion image (right) using a Ga+ primary source for thiol-linked Anti- TDP43 covered electrode. (b) The total secondary ions acquired at the negative polarity of gold (left) and gold-sulfur ion image (right) using a Ga+ primary source for 11-MUA linked Anti-TDP43 formed monolayer.

3.1.3 Electrochemical detection of TDP-43

Differential pulse voltammetry (DPV) was the transduction mechanism for the detection of TDP-43 in this study. Compared with common electrochemical voltammetry, such as cyclic voltammetry, differential pulse voltammetry applies pulse potential following with a potential drop and an immediate measurement of current outputs, in which the charge current is minimized and the sensitivity of the measurement is enhanced. Owing to its high sensitivity, DPV is generously applied to different electrochemical detection systems150-158.

TDP-43 antigen samples prepared in PBS solution was firstly used for proving the

validity of bioconjugation mechanism prepared biosensor. The detection sensitivity

between conventional 3-MPA and 11-MUA monolayer prepared TDP-43 biosensors and

bioconjugation mechanism prepared TDP-43 biosensor were compared and assessed.

The preparation process of 3-MPA and 11-MUA based TDP-43 biosensors with a total time of approximately 51hours were required to prepare those self-assemble monolayer based TDP-43 biosensors, which were significantly longer than the preparation time of the bioconjugation mechanism based TDP-43 biosensor (3hours for bioconjugation of antibody plus 3hours for biosensor fabrication). In order to compare the sensitivity of these three types of TDP-43 biosensors, human recombinant TDP-43 antigen was diluted by 0.1M PBS solution to multiple concentrations ranging from 0.01µg/mL to 1

µg/mL. The 20µL of prepared antigen solution was then applied onto both the bioconjugation prepared biosensors and the 3-MPA and 11-MUA based biosensors. The incubation time of TDP-43 antigen was 1 hour at room temperature. After incubation, each biosensor was rinsed by 1mL 0.1M PBS solution and dried by nitrogen. DPV was then applied to measure the conductivity on the biosensor surface; different current outputs were due to the impedance difference on the biosensor surface, which was provided by the various amount of incubated TDP-43 antigen. Therefore, the current outputs were used to quantify different concentrations of TDP-43 antigen. Figure 4a shows the DPV measurements for the comparison of the sensitivity of the three

different types of biosensors based on PBS prepared antigen. 3-MPA prepared

biosensor showed a higher current outputs comparing with 11-MUA based biosensors,

because the 3-MPA system processes a shorter chain length providing less impedance

on the biosensor surface. However, the bioconjugation method prepared TDP-43 biosensor showed relatively high current density outputs comparing with that of either the 3-MPA or 11-MUA prepared TDP-43 biosensor based on the same detection conditions, indicating that a higher sensitivity was achieved by the bioconjugation method prepared TDP-43 biosensor over conventional SAM type biosensors.

Also, the current density gradients between different concentrations of TDP-43 antigen were apparently larger from the results of bioconjugation based TDP-43 biosensor comparing with those of the 3-MPA and 11-MUA prepared TDP-43 biosensor. Hereby, we defined the value of the change of current density divided by concentration gradients ( i/ C) as the resolution of the biosensor. This term was used

∆ ∆ to quantitatively compare the ability of each biosensor on differentiating different concentrations. High resolution indicates that the biosensor is less possible to produce error signal under the detection concentration range. Derived from the DPV results shown in Figure 16a, the bioconjugation method prepared biosensor showed the highest i/ C of 35.35 A∙mL/(m2∙µg). On the other hand, 3-MPA and 11-MUA system

∆ ∆ prepared biosensor showed a relatively lower value of i/ C of 6.06 A∙mL/(m2∙µg) and

∆ ∆ 6.25 A∙mL/(m2∙µg), indicating a higher resolution achieved by biocnjugation method

prepared biosensor. Moreover, bioconjugation chemistry produces the direct link of

antibody and gold electrode surface, and no other chemicals (such as

unreacted/unoccupied ester groups from activated 3-MPA/11-MUA system) besides

TDP-43 antibody used to react with antigen ensuring the surface linked antibody is the only binding sites available for TDP-43 antigen to interact with. This simplified surface minimizes the possibility of non-specific binding resulting in a high-sensitivity and high-distinguish detection of biomolecules.

Electrochemical impedance spectroscopy was also conducted to measure the surface impedance in order to verify the principle of the detection results from DPV for

PBS test. The EIS measurement was conducted with the same procedure as described above for the DPV measurement with the usage of a redox probe coupling. As shown in

Figure 16b, dashed line shows the impedance of bare electrode and antibody coated electrode; solid lines show the impedance of different TDP-43 antigen concentrations.

Highest concentration of TDP-43 antigen indicates the highest impedance (biggest circle) on the sensor surface at the low frequency region, which was consistent with the lowest current output through the DPV measurement.

Human recombinant TDP-43 antigen was also diluted by undiluted human serum to multiple concentrations, ranging from 0.0005µg/mL to 2µg/mL. The procedure for TDP-

43 antigen in PBS measurement was also used in the serum test for DPV measurement.

The calibration curve for the DPV measurement is shown in Figure 16c with a linear

relationship of Y=-9.607X+4.662 with a R-square value of 0.986 and RSD of 3.56%,

indicating a high reproducibility with n =5. Comparing with TDP-43 protein in PBS

measurement, the measurement of TDP-43 in undiluted human serum showed

decreased current peak for the same concentration range (1 µg/mL to 0.01µg/mL),

which may due to matrix effect caused by complex composition of undiluted human

serum. Also, the biomolecules contained in undiluted human serum may form

unspecific binding onto the gold surface, and the impedance on the surface would

increase, causing a decreasing of the current output. This unspecific absorption was

caused by the components in the solution medium (undiluted human serum) of the

antigen solution, and these components affected identically on all the measurements in

undiluted human serum as shown in Figure 16d comparing with Figure 16a. This type of interference on detection signal caused by components other than the specific analyte is evaluated as matrix effect159, 160. The relative matrix factor based on undiluted human

serum over PBS on the same concentration range was around 0.43.

From the serum detection test, the limitation of detection of TDP-43 antigen was found at 0.0005µg/mL (pink line) in Figure 16d, which aligned closely to the zero concentration line (dark green line). The saturation limit of detection was found at

1µg/mL (light blue line) in Figure 16d, which overlaid with the 2µg/mL (orange line). In order to assess the specificity of this developed TDP-43 biosensor, β-amyloid 42 antigen and T-tau protein were selected for interference tests of this biosensor. 5µg/ml of β-

amyloid 42 antigen solution in human serum and 1µg/mL of T-tau protein in human

serum were incubated onto the TDP-43 biosensors for 1 hour at room temperature.

After incubation, the biosensors were cleaned by DI water and nitrogen gas, and tested using the redox coupling as described in the previous DPV test. The current outputs

based on the detection procedure are shown as the orange color line and dark blue line

in Figure4d, which overlapped with the detection response of non-TDP-43 antigen solution. Another interference test was conducted by a mixed solution of 5µg/mL of β- amyloid 42, 1 µg/mL of T-tau protein with 0.1µg/mL of TDP-43 antigen. The mixed solution was incubated on the TDP-43 biosensor for one hour. The biosensors were then cleaned and tested as described in the previous DPV test. The current output of DPV measurement of this mixed solution was identical to the current output of DPV measurement of only 0.1µg/mL of TDP-43 antigen as the green line signal in Figure 16d.

These two tests also proved that the non-specific absorption was resulted from the components in the undiluted human serum and also demonstrated the high specificity of the simple bioconjugation prepared TDP-43 biosensor in isolating TDP-43 antigen targets for electrochemical detection.

Figure 16. (a) Sensitivity comparison of 11-MUA & 3-MPA based TDP-43 biosensors and bioconjugation method based TDP-43 biosensor using DPV measurement of TDP- 43 antigen in 0.1M PBS solution. (b) Nyquist Plot of TDP-43 antigen in 0.1M PBS solution. (c) Calibration linear curve based on DPV measurement of TDP-43 antigen in undiluted human serum. (d) DPV measurement of TDP-43 antigen in undiluted human serum with limitation and interference tests. 3.1.4 Experimental Procedures for Bioconjugation

3.1.4a Synthesis of thiol-linked anti-TAR DNA-binding protein 43 (anti-TDP-43)

Thiol-linked anti-TAR DNA-binding protein 43 was synthesized for the biosensor fabrication. 0.5 mg of SATA was firstly dissolved in 1 mL of DMSO. 1 µL of

the prepared SATA solution was mixed with 50 µL of anti-TAR DNA-binding protein

43 product in 0.1 M PBS solution based on a molar ratio between SATA and antibody of

20:1 (Hermanson, 2013), and incubated together for 30 min at room temperature. The

solution was then filtered by Amicon ultra-0.5 10 k filter. The filtered solution was

diluted to 0.5 mL with 0.1 M PBS and centrifuged at 12,000 rpm for 15 min at 10 °C producing a concentrated modified antibody sample of a total volume of 50 µL. This filtered antibody solution was stored at 4 °C condition and ready for further deacetylation process. The deacetylation process aimed at generation of sulfhydryl group linked protein by the reaction with a prepared deacetylation solution (0.5 M hydroxylamine, 25 mM EDTA in 0.1 M PBS solution with pH at 7.2. EDTA was added during the reaction to prevent crosslink between sulfhydryl groups. 5 µL of the deacetylation solution was mixed with 50 µL filtered antibody solution and incubated for 2 h at room temperature. Amicon ultra-0.5 10 k filter was applied again, and the deacetylated antibody solution was diluted to 0.5 mL with 10 mM EDTA in 0.1 M PBS solution and centrifuged at 12,000 rpm for 15 min to a volume of 50 µL. This dilution process was repeated for 3 times to remove excessive reagents. Thiol-linked anti-TAR

DNA-binding protein 43 was produced through this process. The produced thiol-linked antibody can be stored at − 20 °C to maintain the bio-reactivity of the antibody for further usage of TDP-43 biosensor fabrication.

3.1.4b Preparation of TAR DNA-binding protein 43 (TDP-43) biosensor using a micro- flow incubation system

TDP-43 biosensor was fabricated based on a three-electrode sensor prototype with thin gold film working and counter electrodes and Ag/AgCl reference electrode.

The thin gold film was deposited by sputtering technique with a thickness of 50 nm and

the Ag/AgCl reference electrode was produced by thick-film printing technology. The

gold sensor prototype was manufactured using a roll-to-roll process, ensuring the production of biosensor cost-effective. The manufacture cost for 100 sensors was approximately $120. The detail configuration of the thin gold film sensor prototype and its electrochemical characterization of the sensor are given in the Supplementary information. A chemical cleaning procedure was applied to remove any oxides and particles on the biosensor surface decreasing the electrode charge transfer resistance.

Typically, a row of 8–10 biosensors were immersed individually in 2 M KOH solution,

0.05 M H2SO4 solution (95.0–98.0 w/w%), and 0.05 M HNO3 solution (70 w/w%) in sequence for 5 min each. The row of biosensors was rinsed by DI water between each cleaning solution. After cleaning, nitrogen air was used for drying the biosensor. The effectiveness of this cleaning process was demonstrated in a previous study. Prepared thiol-linked anti-TDP-43 (anti-TARDBP43) solution was then diluted by 0.1 M PBS buffer with 10 mM EDTA and 0.15 M NaCl to a concentration of 0.25 µg/mL. A micro-

flow incubation system was applied for incubation of thiol-linked antibody solution.

The configuration of the micro-flow incubation system is shown in the Supplementary

information. Continuous flow incubation process can maximize the surface coverage of

protein and enhance the homogeneity and reproducibility of the incubation results comparing with static dropping incubation. The continuous flow system was made with stainless steel and designed to accommodate 10 biosensors inside the flow system.

Also, protein is typically sensitive to temperature (Somero, 1995). Thus, an investigation of the relationship between incubation temperature of antibody and the interaction intensity of antigen-antibody is shown in the Supplementary information. An ideal incubation time was identified at 4 °C. The flow rate was set at 80 µL/min with a retention time for 3 h at 4 °C. After incubation, the biosensors were rinsed with 0.1 M

PBS and dried with nitrogen gas and stored at 4 °C.

3.1.4c Electrochemical spectroscopy impedance (EIS)

The EIS was applied at a frequency range of 0.1–10,000 Hz with an amplitude of 0.01 V.

After biosensor incubated by samples and cleaned by DI water and dried by nitrogen gas, EIS test was then applied using 20 µL of a redox solution with 5 mM in each component of K3Fe(CN)6 and K4Fe(CN)6 to measure the resistivity on the gold sensor surface.

3.1.4d Time-of-flight-secondary ion mass spectrometry (TOF-SIMS)

TOF-SIMS was performed under negative polarity to use the better sensitivity of the instrument to Au, S and their fragments. Element maps for Au and Au-S demonstrated the distribution with a primary source of both C60 and Ga ions. The images acquired

with Ga source were selected to demonstrate the electrode coverage with better spatial

resolution. Experimentally, the primary source was a Ga+ beam accelerated to 30 kV

and bunched to a pulse size of 7 ns and an acquisition rate of 8 kHz. At this setting, the

surface of the electrode was mapped with a spatial resolution of 500 nm. Map stitching

was then used to generate ion maps with a total area of 2 × 2 mm.

3.1.4e Differential pulse voltammetry

Differential pulse voltammetry was used as the transduction mechanism for TDP-43

detection. Typically, before testing, the antigen incubated biosensors were firstly

cleaned by DI water and dried by nitrogen gas. For DPV measurement, 20 µL of a redox

probe solution with 5 mM in each component of K3Fe(CN)6 and K4Fe(CN)6 was

applied onto the biosensor. A voltage ranges of − 0.25 V to + 0.35 V was applied for DPV

measurement with increment of 0.004 V, amplitude of 0.05 V, pulse width of 0.05 s,

sampling width of 0.0167 s and pulse period of 0.2 s.

3.2 Enhancing the sensitivity through utilizing the charge of the analyte

In order to enhance the sensitivity for biomolecular detection, we designed a neutral

charged immunosensor (NCI) in order to fully utilize the electrostatic charge on the

analyte for the quantification of biomarkers. Previous studies demonstrated the

utilization of electrostatic interaction for high-sensitivity electrochemical immunosensor by the application of external charge-labeled materials, such as charged gold

nanomaterials161,162, DNA decorated gold nanoparticles163. The construction of such

sensing platforms required external introduction of charged species to achieve

electrostatic interaction with selected redox systems. However, instead of introducing

the labeled-charged materials, we directly utilized the charged analyte as an interaction

target based on a comprehensive understanding of its electrostatic property.

Electrostatic charge of an analyte is determined by its isoelectric point (pI) and the pH

value of the environment (the testing buffer). Therefore, it is pivotal to comprehend the

electrostatic charge of the analyte for general immunosensor application. In order to

electrochemically quantify the electrostatic charge, a redox coupling solution, such as

[Ru(NH3)6]2+/3+, is typically applied to interact with the analyte150, 164-167. However, when

encountering with positive charged analyte, repulsion of the cationic probe in the

liquid-electrode interface leads to an increase of diffusion length and a decrease in sensitivity161, 168. Therefore, in this study, based on the charge condition of the analyte, we demonstrate the proper applications for both cationic and anodic probes for recruiting its corresponding charged substance. Also, for accurate quantification of electrostatic charge, any charged substance (other than the analyte) on the electrode surface is required to be neutralized in order to provide a neutral environment before the introduction of the analyte. For further directing the charge of the analyte, the buffer pH value of the redox coupling solution is altered. Therefore, there are two circumstances of the electrostatic charge based on electrochemical detection using a

redox probe solution. First, for an analyte with an acidic pI value (< pH=7), a testing buffer with a pH over 7 is used to promote a negatively charged analyte. Second, for an analyte with a basic pI value (>pH=7), a testing buffer with a pH below 7 is used to promote a positively charged analyte. Most of the previous studies were limited to investigate the electrostatic performance only based on negatively charged substances169-172.

We aimed to provide a comprehensive and simple approach to meet the detection demands for all different analytes (with acidic/basic pI) by using electrostatic charges interaction. Consequently, with the purpose of detection of a wide range of analytes, we developed this NCI for both described conditions as a universal immunosensor platform.

The same thin gold film based sensor fabricated by chemical vapor deposition was used as the immunosensor developing platform. The detail description of this sensor configuration and characterization of the stability and reproducibility of this sensor prototype were reported in previous studies53, 173. In order to immobilize the antibody onto the gold sensor surface, N-succinimidyl S-acetylthioacetate (SATA) was applied to conjugate the lysine side chain, producing an external thiol linker that reacted with the gold electrode surface. The characterization of SATA based bioconjugation method was describe above and shown in a previous study53.

Figure 17.a) Formation process of the neutral charged immunosensor (NCI) for acidic pI analyte. Bioconjugation reaction through N-succinimidyl S-acetylthioacetate (SATA) to provide an external thiol linked antibody; thiol-linked antibody was directly incubated onto the gold electrode; thiol-linked poly-arginine was further immobilized to neutralize any electrostatic charge through formation of hydrogen bonding; this fabricated NCI platform was treated with biological samples containing target analyte;

Differential pulse voltammetry (DPV) was performed with the presence of

[Ru(NH3)6]2+/3+ at basic pH condition to fully excite the negative electrostatic charge from the target analyte; DPV measurement demonstrated the difference between +analyte and –analyte. b) Formation process of the NCI for basic pI analyte. SATA conjugated antibody under through direct incubation onto the gold electrode; thiol-linked poly- glutamic acid was further immobilized to neutralize positive electrostatic charge through formation of hydrogen bonding; NCI was further treated with biological samples containing target analyte; DPV was performed with the presence of [Fe(CN)6]3- /4- at acidic pH condition to fully excite the positive electrostatic charge from the analyte; DPV measurement demonstrated the difference between +analyte and –analyte. 3.2.1 NCI Construction for Analyte with Acidic pI.

For an analyte with an acidic pI value, the principle that underlies the NCI is shown in

Figure 17a. The bioconjugated antibody was immobilized on the gold electrode surface

through the gold-thiol reaction, forming a single layer of antibody. Therefore, prior to

the introduction of antigen, the charge condition of the immunosensor surface was only

determined by the antibody itself. After incubation of antigen onto the immunosensor,

the surface charge was then controlled by both the electrostatic charges of the antibody

and the antigen. Consequently, for conventional electrochemical sensor development,

tedious study for a balanced antibody concentration was always required in order to

minimize the effects of electrostatic charge interaction within a certain electrochemical

redox system173-175. Therefore, in order to utilize the electrostatic charge for

quantification of the biomarker, the background signal from the antibody is needed to

be reduced or even neutralized to prevent the possibility that the background signal can be dominant comparing with that of analyte150, because the degrees of charges of

antibody and antigen are different. In order to resolve this research difficulty, we utilized a naturally highly-charged amino acids176, poly-arginine as the surface charge

neutralizer for negatively charged antibody under the pH condition of the

electrochemical testing environment. Arginine possessed the highest isoelectric point

(pI) among amino acids177. It maintained a positive charge for pH under 10.76178. After

immobilization of poly-arginine, the electrostatic charge from antibody on the

immunosensor surface was then neutralized. The whole surface charge condition was

now only determined by the electrostatic property of the analyte, and therefore

enhancing the detection accuracy and sensitivity. The whole construction time of this

NCI was approximately 1 hr, agreeing with the requirement for quick turnaround time

for clinical routine applications.

2+/3+ An electrochemical redox system based on the cationic probe, [Ru(NH3)6] was

applied as the indicator of the intensity of negative charges on the immunosensor

surface. [Ru(NH3)6]2+/3+ was commonly used for electrochemical signal readout for

nucleic acids113, 151 owing to its ability on corporation with negative charged nucleic acids backbone. For the detection of negatively charged analyte, we used the electrostatic interaction between [Ru(NH3)6]2+/3+ and negatively charged amino acids

(Asp & Glu). The cationic probe was attracted into the antigen binding layer instead of

being repulsed to the outer layer of the liquid interface. Therefore, the diffusion

distance of the electron was decreased, providing a higher degree of electron

accessibility of the Au electrode and leading to a higher sensitivity168. In order to fully

utilize the negative charges from the amino acids, a basic buffer (pH=9.6), carbonate

bicarbonate, while maintaining the positive charges of poly-arginine, was used as the composition for the redox solution. Differential pulse voltammetry (DPV) was used as high-sensitivity electrochemical transduction mechanism158, 173.

3.2.2 NCI Construction for Analyte with Basic pI.

For an analyte with a basic pI value, the principle of this NCI is shown in Figure

17b. An analyte with a basic pI value possesses positive charges under physiological

condition. Therefore, in order to interact with positively charged targets, a negatively

charged redox system, [Fe(CN)6]3-/4-, was applied. In this case, the anodic probe was

attracted into the positively charged antigen layer, decreasing the electron diffusion

distance and increasing the sensitivity. To further fully utilize the positively charged

amino acids, the anodic probe was prepared in an acidic buffer, Bis-Tris buffer (pH=6.2).

After immobilization of the antibody, under acidic pH condition, the antibody was

partially positively charged. Therefore, in order to minimize the background signal,

negatively charged poly-glutamic acid was applied as positive charges neutralizer

owing to the low isoelectric point of glutamic acid (pI=3.08). Once the NCI was

constructed, the electrostatic charge readout from DPV would only be determined by

the charge interaction between the positively charged analyte and the anodic probe

[Fe(CN)6]3-/4- under acidic condition.

To demonstrate the validity of this NCI platform, we applied this NCI (1) to detect HE4 (WFDC2) protein, a biomarker for ovarian cancer179, based on the

[Ru(NH3)6]2+/3+ redox interaction with negative charged analyte. We also used this NCI

(2) to detect T-tau protein, a biomarker of neuro-degenerative disorders180, based on the

[Fe(CN)6]3-/4- redox interaction with positive charged analyte. The performance of other

non-charged immunosensor platform was also compared with that of NCI,

demonstrating the high-sensitivity and low-detection limit advantages of this NCI

platform. This study provided a novel perspective on a simple development method for

a universal single-use immunosensor platform with highly enhanced sensitivity.

3.2.3 NCI Development for HE4 Detection

First, we demonstrated this NCI for the detection of HE4 antigen, which

possessed an acidic pI value. In order to utilize the electrostatic charge for the detection,

the charge condition of the antibody while testing, needed to be comprehended. The

surface charge of the antibody was derived from the pH difference between its own

isoelectric point and the pH value of the testing buffer. Therefore, in order to fully

induce the negative charges from the antibody, the influence of basicity of the testing

buffer on electrostatic charge was firstly evaluated based on the HE4 antibody. Two

different buffers prepared redox coupling, phosphate buffer saline (pH=7.4) and carbonate bicarbonate (pH=9.6), were investigated using DPV in the presence of 50µM of [Ru(NH3)6]2+/3+. As shown in Figure 18a, using the HE4 antibody layer covered immunosensor, the carbonate bicarbonate dissolved redox coupling resulted in a higher current outputs comparing with that of phosphate buffer saline dissolved redox coupling, indicating that a more basic buffer medium improved the electrostatic charge interaction between the protein and the ruthenium based redox probe. Therefore, CBC buffer based redox coupling solution was chosen for further application in HE4 detection.

Figure 18. a) Comparison of DPV measurements of electrostatic interactions of HE4 antibody with [Ru(NH3)6]2+/3+ in CBC buffer (pH=9.6) and in PBS buffer (pH=7.4). b) DPV 2+/3+ measurement by [Ru(NH3)6] in CBC buffer (pH=9.6) on HE4 antibody with the

addition of poly-arginine as negative charge neutralizer. c) DPV measurements by [Ru(NH3)6]2+/3+ in CBC buffer (pH=9.6) on various concentrations of poly-arginine alone. d) Proposed cation probe interaction models for positively charged protein and negatively charged protein.

As shown in Figure 18b, the blue line (1) demonstrates the electrostatic charge signal based on the negatively charged antibody (2.43µM) covered electrode surface. To investigate the charge compensation ability of poly-arginine, we further incubated

SATA conjugated poly-arginine onto the HE4 antibody covered electrode, producing a neutral charged surface under testing condition. The red line (2) in Figure 6b presents the electrostatic charge signal based on negatively charged antibody electrode surface plus the selected concentration (30.4µM) of positively charged poly-arginine. The decrease of current output was caused by that the negative charged amino acids in the antibody, which were interacted with poly-arginine instead of [Ru(NH3)6]2+/3+, due to the

formation of hydrogen bonding between negative charged amino acids and positive

charged poly-arginine. Consequently, the reduction process of [Ru(NH3)6]3+ to

2+ [Ru(NH3)6] was weakened. To verify the electrostatic interaction was the dominant component for the decrease of current outputs, we further evaluated the interaction of

[Ru(NH3)6]2+/3+ with different concentrations of poly-arginine alone on the gold sensor

surface. As shown in Figure 18c, various concentrations of poly-arginine did not cause

significant decrease of the current outputs based on the cation redox reaction, indicating

that the cation redox reaction process was not impeded by the addition of poly-

arginine. Therefore, we concluded that the decrease of current in Figure 6b was not initiated by the electron transfer hindrance from the addition of poly-arginine.

Furthermore, this phenomenon (Figure 18c) can be explained as that the repulsion between cation probe and positively charged biomolecules creates a gap preventing their interactions.

Based on above observations, we proposed an electrostatic interaction model as shown in Figure 18d. After the redox solution is placed on the sensor surface, when the immobilized biomolecules are oppositely charged, the attractive force moves the cation probe closer to the electrode surface, decreasing the electron transfer distance during the redox reaction. Consequently, the repulsion and attraction of cation probe with differently charged biomolecules leads to different electron diffusion distance, further delivering different intensity of current outputs from the redox reaction.

With the understanding of the electrostatic interaction, this developed NCI was further applied to investigate the electrostatic detection of HE4 antigen. For electrostatic quantification of HE4 antigen, the amino acids sequence of HE4 antigen were firstly evaluated, which was used for calculation of the isoelectric point of the HE4 antigen.

The pI for HE4 antigen is 4.50, which is lower than the pH value of CBC buffer, indicating that under the CBC buffer based redox coupling solution, it possesses strong negative charges. Multiple concentrations of HE4 antigen in human serum was incubated on the HE4 NCI for 30 min at room temperature. After this short incubation

time, DPV measurements were performed using CBC based [Ru(NH3)6]2+/3+ redox

solution. A quiet time of 15s was selected, allowing the electrostatic cooperation

between [Ru(NH3)6]2+/3+ and negative charged HE4 antigen. As shown in Figure 19a, the

concentrations of HE4 antigen show a positive correlation with the current outputs. A

low detection limit of HE4 antigen was observed at 2.5 pM. Comparing with recent

report on nanomolar level HE4 detection181, the results of this study demonstrated a

detection limit of three order of magnitude lower, confirming the high sensitivity

property of this NCI. The calibration curve is shown in Figure 19b with a R-square value of 0.905 and RSD value of 3.066%, indicating a good linear relationship and the high reproducibility of this detection platform.

Figure 19. a) DPV measurements of various concentrations (2.5pM-250000pM) of HE4

antigen using HE4 NCI. b) Calibration curve for the DPV measurements of NCI.

c)Comparison of detection resolution of NCI with non-neutralized immunosensor in

response to various HE4 concentrations (2.5pM-250000pM). d) Interference study using

CA125 antigen on the HE4 NCI.

To further demonstrate the high-sensitivity property owing to the application of charge neutralizer, we compared the performance of this NCI platform with a non- neutralizer platform, which only possessed the thiol-linked antibody on the gold sensor surface. These two types of immunosensor employed identical antigen incubation

process and tested by the same electrochemical method. As shown in Figure 19c, the

changes of current referring with the base line are presented. The NCI platform not only

demonstrated a higher resolution for the same concentration range of detection, but also

defined a lower detection limit comparing with those of the non-neutralizer platform, indicating a lower possibility of false positive results and a higher sensitivity of this NCI platform.

The specificity of this HE4 NCI platform was further assessed by using CA-125 protein, also a biomarker of ovarian cancer179. Mixed CA-125 (11000pM) and HE4

antigen (2500pM) samples were applied to incubate on the prepared HE4 NCI. We also

applied only CA-125 (11000pM) incubation on the HE4 NCI. As shown in Figure 19d,

the change of current was compared between the interference samples and the only

HE4 sample and non-analyte baseline. The change of current was lower than the RSD of

this NCI platform. Therefore, the interference test demonstrated that there was no

signal based on only CA-125 antigen and also proved that there was no interference

from CA-125 antigen on the recognition of HE4 antigen. These two experiments

validated the high-selectivity using the NCI platform.

3.2.4 NCI Development for Tau Detection

In order to demonstrate the NCI on analysis of analyte with basic pI value, T-Tau

protein was employed as target for the NCI development. We first analyzed the

influence of the pH condition of the detection buffer by comparing the DPV

measurement of the immobilized antibody using PBS based redox (pH=7.4) and Bis-Tris

based redox (pH=6.2). As shown in Figure 20a, with the application of more acidic

buffer, the electrostatic charge performance of the antibody (6.67 nM) was enhanced

based on the DPV measurement using [Fe(CN)6]3-/4-. In order to produce NCI based on a

positively charged condition, a highly-negative charged, hydrophilic amino acid,

aspartic acid was applied as the charge neutralizer. In order to immobilize the poly-

aspartic acid, a bioconjugation mechanism aiming at modification of side chain

carboxylic group was applied by reacting with 1-ethyl-3-(-3-dimethylaminopropyl) carbodiimide hydrochloride (EDC), cystamine, and 2-Mercaptoethylamine•HCl (2-

MEA) sequentially, providing an external thiol linker. As shown in Figure 20b, after incubation of selected concentration (3.10 nM) of thiol-linked poly-aspartic acid, the

DPV signal produced by the electrostatic interaction was significantly decreased based on the positively charged amino acids interaction with [Fe(CN)6]3-/4-. To verify that the

decrease of DPV signal was caused by the formation of salt bridge (between negatively

charged glutamic acid and positively charged amino acids of the antibody at the testing

environment) instead of electron hindrance due to addition of surface molecules, we

further analyzed the interaction between [Fe(CN)6]3-/4- and different concentrations of

poly-glutamic acid. As shown in Figure 20c, the interaction of anion redox probe with

negatively charged poly-glutamic acids did not change significantly with a large

increase of concentration of poly-glutamic acid. This phenomenon first confirmed that the decrease of DPV signal shown in Figure 20b was not caused by increase of surface impedance. Second, the unchanged DPV signal in Figure 8c indicated that the redox reaction process was not interfered by the immobilized poly-glutamic acid, because there was repulsion force between the anion probe and negatively charged biomolecules. Moreover, based on our observations, an interaction model between the anion probe and charged biomolecules was suggested as shown in Figure 20d. We proposed that the difference in electron diffusion distance caused by electrostatic interaction and incorporation led to different current outputs behavior. This fabricated

NCI was further applied for incubation of T-Tau antigen.

Figure 20. a) Comparison of DPV measurements of electrostatic interactions of Tau

antibody with [Fe(CN)6]3-/4- in Bis-Tris buffer (pH=6.2) and in PBS buffer (pH=7.4). b) DPV

measurements by [Fe(CN)6]3-/4- in Bis-Tris buffer (pH=6.2) on Tau antibody with the addition of poly-glutamic acid as positive charge neutralizer. c) DPV measurements by

[Fe(CN)6]3-/4- in Bis-Tris buffer (pH=6.2) on various concentrations of poly-glutamic acid alone. d) Proposed anion probe interaction models for negatively charged protein and positively charged protein.

T-Tau protein has a basic isoelectric point (pI=8.6). Therefore, under pH<8.6, T-

Tau protein is highly positive charged173, 180. Multiple concentrations of T-Tau antigen were incubated onto the prepared T-Tau NCI for 30 min. After incubation, the immunosensors were tested by DPV in the presence of Bis-Tris based [Fe(CN)6]3-/4-

solution. As shown in Figure 21a, the DPV measurements demonstrated a positive

correlation with the current outputs, indicating the positive electrostatic interaction

3-/4- between the T-Tau antigen with [Fe(CN)6] . Comparing with the results of recent

studies on Tau detection that demonstrated a sensitivity at micro/nano-molar level173, 182,

our NCI platform establishes a low detection limit of 0.968 pM in human serum. The

calibration curve is shown in Figure 21b with a R-square value of 0.943 and RSD around

3.6%.

We further compared this NCI with non-neutralized platform. Non-neutralized platform did not incorporate the charge neutralizer, poly-glutamic acid. As shown in Figure 21c, comparing with non-neutralized immunosensor, our T-Tau NCI demonstrated one order

lower of the detection limit, confirming the high-sensitivity of this NCI platform. For

interference study, 9.68pM T-Tau antigen sample was mixed with TAR DNA binding

protein 43 (133pM) and β-amyloid 42 antigen (153pM), which are confirmed biomarkers

for neurodegenerative disorders20, 183, 184. As shown in Figure 21d, all the interference

currents produced by non-target analyte were within the range of RSD from the DPV

calibration curve. Therefore, this T-Tau NCI also confirmed its high-specificity on

electrochemical detection.

3.2.5 Conclusion

In summary, the developed neutral charged immunosensor (NCI) provides a

simple, rapid immunosensor fabrication protocol suitable for routine clinical point-of-

care usages. By triggering the electrostatic charge of the target analyte, we provide two

perspectives for the NCI, covering the detection methods for both negatively charged and

positively charged analytes. Therefore, this NCI platform provides a universal method

for electrochemical quantification of antigen. The picomolar level detection limit of this

simple immunosensor platform is sufficiently low for quantification of most of the

circulating protein biomarkers.

Figure 21. a) DPV measurement of various concentrations of T-Tau antigen (454pM- 0.968pM) using NCI. b) Calibration curve of the DPV measurements of T-Tau NCI. c) Comparison of detection resolution of NCI with non-neutralized immunosensor in response to various T-Tau antigen concentrations of 454pM-0.968pM. d) Interference study using TDP-43 and Aβ-42 antigen on the T-Tau NCI. 3.2.6 Experimental Procedures

3.2.6a Bioconjugation Procedures

N-succinimidyl S-acetylthioacetate (SATA) was prepared to immobilize anti-HE4 on the

surface of the electrode. SATA was firstly dissolved in dimethyl sulfoxide (DMSO) to

achieve a concentration of 12 mg/mL. 2μL of prepared SATA solution was then reacted

with 30μL anti-HE4 (2.336 μg/μL) for 30 minutes at room temperature. The reacted

solution was transferred into an Amicon ultra-0.5mL 10k filter tube for centrifugation

with a rotation speed of 10000 rpm at 5°C for 15 minutes. The filtered solution was

reacted with 5μL 0.5M hydroxylamine and 25mM EDTA in 0.1M PBS solution at room

temperature for 1 hour to deprotonate the protected thiol group before the mixture was

centrifuged again under the same setting. After the second filtration, the thiol-linked anti-HE4 was obtained and stored for future use.

Thiol-linked poly-arginine was prepared in a similar manner. 1 mg of poly-

arginine was dissolved in 400 μL of 0.1 M PBS. 6 μL of SATA solution (5 mg/mL in

DMSO) was reacted with the prepared poly-arginine solution for 30 minutes at room temperature. The solution was then transferred into an Amicon ultra-0.5mL 10k filter tube for centrifugation with a rotation speed of 10,000 rpm at 5 °C for 15 minutes. The filtered solution was reacted with 5 μL 0.5M hydroxylamine and 25 mM EDTA in 0.1 M

PBS solution at room temperature for 1 hour, vortexed every 10 minutes. The solution was then transferred to the Amicon ultra-0.5mL 10k filter tube for centrifuge with the same settings for 15 minutes. After filtration, the thiol-linked poly-arginine was obtained and stored for future use.

3.2.6b Neutral Charged Immunosensor Preparation and Testing Procedures

A row of 8-10 biosensors were immersed in 0.1M KOH, 0.09M H2SO4, and 0.09M

HNO3 for cleaning, consequentially, for 10 minutes. Biosensors were rinsed by DI water

in between the immersions. The cleaned sensors were then rinsed with DI water and

dried gently by nitrogen gas. Thiol-linked anti-HE4 was diluted by 0.15M NaCl and

10mM EDTA in 0.1 M PBS solution to a concentration of 2.43μM.20μL of this solution

was dropped onto the sensing area of each cleaned biosensor and were incubated for 30

min at room temperature. After the antibody incubation, each sensor was rinsed by DI

water and dried by nitrogen flow. The conjugated poly-arginine was diluted in 0.15M

NaCl and 10mM EDTA in 0.1 M PBS solution to a concentration of 30.4μM, then 20μL

of the diluted solution was dropped onto each sensor and was left to incubate for 30

minutes at room temperature. Each of the biosensor was then rinsed and dried.

Different concentrations of HE4 protein solutions were prepared in human serum and

20μL of each HE4 protein solution was dropped onto each biosensor. HE4 protein was

left to incubate for 30 minutes at room temperature. After the incubation, all the

biosensors were rinsed with DI water and dried by nitrogen gas. The differential pulse

voltammetry (DPV) measurement was performed for each sensor using 20μL of the

redox solution, which contained 50μM Ru(NH3)6Cl3, 4mM K3Fe(CN)6, 25mM NaCl,

prepared in 0.01M carbonate bicarbonate buffer. CHI 660C Electrochemical Workstation was used to perform the DPV tests. The initial potential was set at 0V and the final

potential was set at -0.35V. The potential increment was at 0.004V, the amplitude at

0.05V, the pulse width at 0.05s and the pulse period at 0.2s. The quiet time was set at

15s.

Chapter 4 Conclusion

In conclusion, through the construction of the bio-electrode interfaces, multiple

electrochemical biosensing strategies have been developed for the detection of nucleic

acids and proteins. These developments resolve critical challenges in the field of

electrochemical biosensing. First, CRISPR Cas system was utilized as a novel

recognition element for electrochemistry based nucleic acid sensing, enhancing the

accuracy of detection. The combination of CRISPR systems with electrochemical sensor

platform provided a simple and rapid system for probing the enzymatic activity of

CRISPR type V system.57 Second, we incorporated CRISPR into the classic stem-loop based nucleic acid recognition probe, enhancing the detection sensitivity through the

CRISPR recognition induced on-target cleavage. Furthermore, CRISPR was applied to mediate DNA electrochemistry, providing a new mutation analysis strategy which can discriminate single-mutation at a dose-independent manner. For the detection of protein, instead of conventional construction of bio-electrode interface through self-

assemble monolayer modification, bioconjugation chemistry was applied to modify the

recognition element first, providing a thiol-modified antibody able to immobilize on the

electrode through a one-step process, which significantly decreases the turn-around

time for biosensor fabrication.53 Finally, for the electrochemical transduction process using redox probes, we utilized the electrostatic interaction between certain redox probe and target protein, providing a signal-on target quantification strategy.54 Overall, our efforts throughout this doctoral research provided new chemistry and engineering insights into the biosensing research community. Based on these novel concepts, biosensing systems have been developed successfully with capability to quantitative detect nucleic acids and proteins. Future researches can utilize these established biosensing strategies as guidelines for the development of diverse and robust biosensors.

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