Graphene-Based Biosensors to Detect Dopamine: Comparison
Please note this is a comparison between Version 1 by Ajeet Kaushik and Version 2 by Lindsay Dong.

Parkinson’s disease (PD) is a neurodegenerative disease in which the neurotransmitter dopamine (DA) depletes due to the progressive loss of nigrostriatal neurons. Therefore, DA measurement might be a useful diagnostic tool for targeting the early stages of PD, as well as helping to optimize DA replacement therapy. Moreover, DA sensing appears to be a useful analytical tool in complex biological systems in PD studies. Graphene-based DA sensors are emerging analytical tools for PD diagnostics.

  • dopamine
  • Parkinson’s disease
  • graphene
  • biosensing

1. Introduction

Parkinson’s disease (PD) is the second most common human neurodegenerative disorder, after Alzheimer’s disease (AD) [1]. The disease is diagnosed based on motor impairment, including bradykinesia rigidity or tremor; this is when about 70% of the dopaminergic neurons of the substantia nigra pars compacta are degenerated due to α-synuclein deposits. PD is also diagnosed clinically once the synucleinopathy is already advanced. Researchers and clinicians indicate a potential temporal window before the onset of specific signs and symptoms of the disorder during which potential disease-modifying therapy could be administered to prevent or delay the disease development and progression. Indeed, there is a need for an early diagnosis primarily based on quantifiable measures (i.e., biomarkers) to refine qualitative assessments [2]. From a neurochemical perspective, PD is a neurodegenerative disease in which depletion of the catecholamine DA in the nigrostriatal system appears due to the loss of nigral neurons and striatal terminals. Over the years, the neurotransmitter loss progresses to reach only 3% of normal DA concentration in the putamen of patients with pathologically proven end-stage PD. In untreated PD patients, most studies found significantly decreased DA levels in the cerebrospinal fluid (CSF), reflecting dopaminergic cell loss [3]. Eventually, an individual develops motor symptoms, including bradykinesia, rigidity, tremor, and postural instability, which result from this drop in DA level. This means that DA level measurement might be a useful diagnostic tool for targeting the early stage of the defunctionalization of DA-producing neurons (nigrostriatal dopaminergic denervation) to enable the development of approaches to retard progression or even prevent the disease [4].
Due to high spatial and temporal resolution, high sensitivity and selectivity, and the possibility of direct monitoring at low cost and with the leverage of user-friendly tools, oxidation-based electrochemical sensing platforms are becoming a more popular and developed technique that is being implemented in a biological environment [5][6][7][12,13,14] and also for DA detection [8][15]. Efforts have been made to detect in situ DA, e.g., in the brain or living cells. Asif et al. applied the Zn-NiAl LDH/rGO superlattice electrode to track the DA released from human neuronal neuroblastoma cell line SH-SY-5Y [9][16]. Li et al. demonstrated a developed nanoelectronic biosensor, as shown in Figure 1, for monitoring the DA release from living PC12 cells [10][17]. Figure 1a shows the illustration of a DNA-aptamer modified by a multiple parallel-connected (MPC) silicon nanowire field-effect transistor (SiNW-FET) device, as well as the process of DNA-aptamer immobilization of the MPC SiNWFET. This device detects the DA under hypoxic stimulation from living PC12 cells. This developed MPC aptamer/SiNW-FET device demonstrated a DA detection limit of up to <10−6 M with high specificity when exposed to other chemicals, such as tyrosine, ascorbic acid (AA), phenethylamine, norepinephrine, epinephrine, and catechol. Wu et al. fabricated reproducible miniaturized, multi-layered, graphene-based sensors with astonishingly high sensitivity when compared with other sensors [11][18]. Figure 1b (i) shows the nanofabricated miniaturized multilayer graphene sensor electrodes. Figure 1b (ii) shows the scanning electron microscopy (SEM) image of the top of the sensor array and the AFM image of the sensor surface. Figure 1b (iii) depicts the mechanism behind it. The DA undergoes a redox reaction and is oxidized to dopamine-o-quinone (DOQ) by applying voltage. The sensitivity of the fabricated sensor is monitored by fast-scan cyclic voltammetry (FSCV) measurements. Figure 1b (iv) displays the area-normalized electrochemical current (IEC) curves in response to the DA solution. The fabricated graphene sensor achieved a high sensitivity of 177 pAμm−2μM−1 in response to the DA. It is concluded that the MPC aptamer/SiNW-FET sensor has shown improved specificity and an LOD up to <10−11 M for exocytotic DA detection, as compared to other existing electrochemical sensors. The real-time monitoring of DA induced by hypoxia demonstrates that for triggering the DA secretion, intracellular Ca2+ is required, which is commanded by extracellular Ca2+ influx instead of the release of intracellular Ca2+ stores. Such a device, capable of coalescing with living cell systems, opens a new gateway towards the biosensor for the futuristic studies of clinical disease diagnostics.
Figure 1. (a) DNA-aptamer-modified MPC SiNW-FET biosensor for dopamine; illustration of FET device for detecting exocytotic dopamine under hypoxic stimulation from living PC12 cells; (b) a semi-log plot of response as a function of dopamine concentration [10]. (c) Schematics of a graphene-based electrode used for measurements of DA; graphene electrode is mounted on a SiO2/Si substrate, and a fluidic chamber is filled with PBS solution containing target dopamine; (d) SEM image of the graphene-based sensor array; AFM topographic image of CVD grown multilayer graphene (e) mechanism behind the FSCV measurements of dopamine; and (f) noticeable area-normalized electrochemical current (IEC) response to the dopamine concentrations [11].

2. Analytical Performances of DA Graphene-Based Biosensors

Detecting biomolecules in real samples is associated with the interaction of other compounds with similar oxidation potentials during detection [12][20]. Thus, designing sensors for the DA monitoring in biological samples, such as routine clinical ones, is challenging since electrochemically active compounds commonly found in body fluids, such as AA, uric acid (UA), and glucose (Glu), constantly interact with each other during detection due to their similar oxidation potentials. Moreover, the present macromolecules, including proteins, can non-specifically adsorb on the electrode surface, thus hindering the electron transfer rate [13][21]. Thus, the development of electrochemical methods for the analysis of DA in a complex matrix must address all these possible interactions to enable its successful DA detection in a simple, rapid, and highly selective way. The limitation caused by overlapping voltametric signals of compounds with very close oxidation potentials and relatively poor selectivity can be avoided by applying different sensing layers that enable separate detection of the electrochemical signals. Several electrode-modification substances, such as oxides, conducting polymers, and nanomaterial, have been adopted for this purpose. Nanomaterial-modified electrodes, especially with graphene and its derivatives, such as reduced graphene oxide (rGO) and graphene oxide (GO), have recently attracted great focus in electrochemical biosensing approaches [14][12][15][16][17][18][19][7,20,22,23,24,25,26]. Due to their unique structure, graphene-based materials increase the conductivity of the compounds used in electrochemical measurement systems. Owing to their large surface area, they offer a high number of accessible active sites to detect analytes (Figure 2) [17][24]. Graphene is always admired for its excellent properties among the various sensing materials for DA due to its excellent electrical conductivity and π−π interaction between the aromatic rings of DA and graphene. Butler et al. developed a graphene ink-based, ultrasensitive electrochemical sensor for the detection of DA. The lowest limit of detection is reported as 1 nM. This sensitivity and selectivity of the sensor are achieved by tuning the surface chemistry of graphene. Figure 2a shows a schematic illustration of the fabrication of the DA sensor. The curves of Figure 2b depict the effect of annealing the graphene towards the DA response from 55 pM to 50 μM, using DPV measurements. Scanning electrochemical microscopy (SECM) mapping confirmed that the graphene layer (Figure 2d−g) shows higher oxidation at the edges of the flakes. Figure 2d,f display the height maps for two different regions of the graphene ink film-based sensor. Figure 2, for example, shows the electrochemical mapping of the graphene ink with 100 mM DA in PBS. At different concentrations, the total activity is enhanced, as seen by the increased magnitude of the current in the electrochemical response. Considering the 2D defects and the active edge sites of graphene ink, it can be an ideal candidate for printable and low-cost DA sensing devices/systems.
Figure 2. (a) Schematic representation of fabrication and electrochemical testing process of the graphene ink-based DA sensor. (b) Differential pulse voltammogram of the response towards DA detection from 5 pM to 50 μM. (c) Normalized peak current values versus DA concentration. (d) Height map, measured using scanning electrochemical microscopy (SECM) and (e) the corresponding electrochemical map with 1 mMDA. (f) A height map of a different region of the graphene film and (g) the corresponding electrochemical map with 100 mM DA [15].