Review—Catalytic Electrochemical Biosensors for Dopamine: Design, Performance, and Healthcare Applications

Dopamine is an essential neurotransmitter for daily cognitive functions controlling many neurophysiological processes including memory, cognition, and physical control. Development of analytical methods and sensors to detect dopamine is important for health monitoring and neurological research. This review provides an overview of recent advances in the development of electrochemical catalytic biosensors based on enzyme and enzyme-mimetic materials and discusses their potential applications for measurements of dopamine in biological fluids. The first part of the review summarizes and critically assesses the different types of enzymes and enzyme mimetic materials that can be used to catalytically convert dopamine, followed by a discussion of the biosensor’s fabrication, key design parameters, and detection mechanism on various electrode platforms ranging from single-use screen-printed electrodes to microneedles and implantable microelectrodes. The second part provides examples of measurements of dopamine in biological samples, including saliva, urine, serum, cell cultures, and brain tissue. We conclude with a summary of advantages and limitations of these devices in the clinical field, and an outlook to future research towards the implementation and broader adoption of electrochemical biosensors in neurophysiology, pharmacology, and the clinical field.

Dopamine is a catecholamine neurotransmitter widely distributed throughout the central nervous system (CNS).Dopamine influences a variety of rewarding behaviors, attention, motor control, neuronal plasticity, and plays a critical role in learning and memory.2][3] There are two main types of postsynaptic receptors specific to dopamine within the brain: ligand-gated receptors (ionotropic) and G-protein coupled receptors (metabotropic). 4Gprotein coupled receptors can be found in subfamilies of D 2 , D 3 , and D 4 , with D 2 receptors being the major target of neurological sensing of dopamine 5 due to the increased density of dopamine neurons on dendrites and perikarya located within the striatum of the brain. 6ealthy neurons contain multiple synaptic vesicles filled with a variety of neurotransmitters with dopamine accounting for approximately 80% of the total catecholamine concentration within the brain. 7Neurotransmitters are released into the synaptic cleft, allowing for their uptake by ion channels within the postsynaptic receptors and facilitating muscle movement and regulatory functions like breathing.However, diseased neurons have a reduced number of synaptic vesicles with lower concentrations of neurotransmitters, and therefore causing decreased levels to reach out to the postsynaptic receptors, leading to lack of signal processing and neurological disorders.Even minor fluctuations of dopamine at these receptors have been related to systematic dysfunctions along with the possible onset of neurodegenerative diseases such as with Parkinson's disease, [8][9][10][11] Schizophrenia, [12][13][14] and Huntington's Disease. 15,16opamine is generated mainly within the striatum, more specifically at the substantia nigra, ventral tegmental area, and hypothalamus, 17 where accessibility and measurements in these areas are difficult and require invasive procedures.Dopamine can be found in concentrations ranging from 0.01-1 μM in biological fluids for healthy individuals, and in the low nanomolar for individuals with Parkinson's disease. 18Easily accessible fluids such as sweat, blood and urine are more convenient to measure but can have low and variable concentrations and provide less precise measurements as compared to measurements at the production site.Implantable biosensors with minimally invasive microelectrodes can be used in vivo to determine dopamine in the brain and monitor progression of neurodegenerative diseases.
Given the wide range of physiological and pathophysiological effects of dopamine, precise and accurate measurement in biological systems is of great importance.The ability to selectively and specifically measure clinically-relevant levels is critical for many aspects of neuroscience while direct monitoring of dopamine in vivo is important for understanding the mechanisms of drug addiction, learning, and other neurological conditions.However, the presence of neurotransmitters such as dopamine at low and variable concentrations in biological samples and the extracellular space of the CNS provides a great challenge for any system designed to detect neurotransmitters in real-time.Some examples of challenges include cross-talk from other neurotransmitters causing altered dopamine signals, lack of sensitivity to reach detection of physiological relevant concentrations and achieving a small enough size of measurement probes to limit damage when implanted.Measurement of dopamine in vivo has been achieved using microdialysis 3,19 with high-performance liquid chromatographytandem mass spectrometry (HPLC-MS), 20 with some examples of electrochemical, 21 coulometric, 22 or fluorescence 23 detection in the dialysate.Some downfalls of microanalysis includes measurement of dopamine fluctuations on the timescale of minutes to hours, relatively large probes, and complex measurement set-up that is not easily adaptable for real-time monitoring.Similar issues are also seen with coulometry and fluorescence, as such measurements cannot be adapted in vivo and therefore they lack the ability to provide real-time information.
5][26][27] Dopamine is electrochemically active and can be measured directly using an electrode held at its characteristic potential.Electrochemical methods for on-line in vivo analysis of dopamine with carbon fiber microelectrodes (CFMEs) 26 have been described, and the use of methods such as amperometry, [28][29][30] differential pulse voltammetry (DPV), [31][32][33] fast scan cyclic voltammetry (FSCV), 2,34 cyclic voltammetry (CV), 35 electrochemical impedance spectroscopy (EIS), 36 and square wave voltammetry (SWV) 37 is well established.The use of CFMEs in FSCV mode has been a method of choice to measure in vivo levels of brain dopamine due to the versatility of CFMEs and the ability to measure both oxidation and reduction somewhat simultaneously. 26or example, Li et al. describes a potential-gated organic electrochemical transistor used with FSCV in rat brains, successfully determining dopamine release upon stimulation. 38Dopamine sensors and FSCV for brain analysis were largely developed by the z E-mail: eandrees@clarkson.eduECS Sensors Plus, 2024 3 020601 pioneering work of Wightman, 1,39 and several microelectrodes (e.g.carbon fiber) are now commercially available for the direct electrochemical detection of dopamine.FSCV also requires background subtraction and may result in electrode passivation and difficulties differentiating among different neurotransmitters due to overlapping signals from co-occurring species.
The oxidation potential for dopamine is large (∼0.5-0.7 V) and encompasses the oxidation range of many catecholamine neurotransmitters and other substances in the central nervous system such as ascorbic acid, 40,42 uric acid, 41 serotonin, 43,44 epinephrine, 45 norepinephrine, 46 L-DOPA (L-3,4-dihydrophenylalanine), 30 DOPAC (3,4-dihydroxyphenylacetic acid), 47 and many more chemicals present within the complex matrix of neuronal tissue.For this reason, it can be difficult to differentiate dopamine from multiple cooccurring interfering species, such as ascorbic acid, uric acid, and other neurotransmitters. 48,49The modification of electrode surfaces with enzymes or nanocomposite layers is one strategy to enhance selectivity and reject interferences from species like uric acid, ascorbic acid, serotonin, tyrosine, serine, and cysteine when detecting dopamine. 50A common strategy to eliminate interferences is to cover the electrode with permselective membranes such as Nafion 51,52 that is known to exclude negatively charged interferents such as ascorbate. 52However, some of these layers contain materials that can be toxic and not ideal for implantation, limiting their use in vivo.Other biocompatible membranes such as chitosan have been used and provided ascorbic acid rejecting properties, further enhancing the specificity for dopamine.An alternative method to detect dopamine involves the use of biosensors constructed with oxidase enzymes, such as tyrosinase and laccase, or nanoenzyme-mimetic catalysts, such as CeO 2 and other metal oxides, that convert dopamine into dopaquinone. 24,53Dopaquinone can be detected electrochemically, allowing the detection window to shift down from the dopamine oxidation potential of ∼ +0.16 V to a lower operating potential of ∼ −0.1 V, corresponding to the electroreduction of dopaquinone.This strategy reduces the effect of interferences and minimizes electrode passivation problems, addressing some of the specificity issues. 25his review discusses the status and progress made within the past few years on the development of electrochemical biosensors with enzymes and enzyme-mimetics for measuring dopamine in biological samples, including serum, urine, cell cultures, and intact brain, and their applicability in the biomedical field.First, we describe the fabrication and characterization of different electrochemical biosensors and review the different types of enzyme and enzyme-mimetic materials that can be used for the detection of dopamine.In the second part, we provide specific examples of applications and performance characteristics of recently reported dopamine biosensors in different types of biological samples and environments.We conclude with a summary of advantages and limitations of these devices and future outlook outlining the potential of these sensors in clinical diagnosis, neuroscience, and pharmaceutical research.

Principle and Operation of Catalytic Biosensors for Dopamine Detection
Principle of catalytic biosensors for dopamine.-Catalyticbiosensors for dopamine rely on the use of enzymes or enzymemimetic catalysts to specifically recognize and catalytically convert dopamine to dopamine-o-quinone by accelerating the rate of dopamine oxidation.The enzymatically generated dopamine-oquinone can be measured by electrochemistry at a low applied potential, thus reducing the effect of the majority of interfering compounds that are not oxidized or reduced in that low potential range.Catalytic biosensors integrate the biological/biomimetic receptor onto the surface of an electrochemical transducer, typically a glassy carbon electrode (GCE), screen-printed electrode (SPCE), CFME, or other types of microelectrode platforms.In addition to solid electrodes, open-cavity carbon-based nanopipettes can also be used for increasing spatial resolution of dopamine measurements, with demonstrated capabilities on brain slices achieving a detection limit (LOD) of 56 nM using FSCV and amperometry. 54To create dopamine biosensors, electrodes are modified with the selected catalysts (sensing layer), typically immobilized within biocompatible layers or attached though chemical modification procedures.A critical step in the development of catalytic biosensors is to ensure stable immobilization of the bio receptor onto the electrode surface while maintaining its recognition/catalytic properties and minimizing effects of coexisting species.The immobilization layers might contain materials such as metal or metal oxide nanoparticles, carbon nanotubes, or graphene to amplify the electrochemical signal, or materials with perm-selective properties to prevent interferences, such as the ascorbic acid-blocking properties of chitosan.This immobilization step tends to be easier when done with larger electrodes, as compared to small size implantable probes, due to the fragility, size, available active surface area, and stability considerations.A general design of electrochemical biosensors based on enzymes and enzyme-mimetic catalysts is provided in Fig. 1, summarizing the different types of dopamine-specific enzymes and enzyme-mimetic catalysts, electrodes and electrode materials, and immobilization strategies.
Selection of enzymes.-Tyrosinase and laccase are the two most commonly used enzymes for dopamine biosensors to date.These enzymes are able to catalyze oxidation of a broader range of phenolic and catecholamine compounds with varying degrees of selectivity, depending on the pH and the environment. 55Tyrosinase, or polyphenol oxidase, is a copper-containing metalloenzyme that uses oxygen to catalyze conversion of monophenols to o-diphenols via o-hydroxylation, and further to o-quinones. 56Laccases are another family of copper proteins, known as multi-copper oxidases (MCOs), with specificity towards phenolic catechol amines derived from the rate-limiting step of the copper T1 reduction located close to the surface of the enzyme, facilitating the oxidation of phenolic moieties at a low potential. 57Traditionally, tyrosinase and laccases catalyze the conversion of phenolic compounds to their respective quinone derivatives. 55,580][61] These enzymes are able to catalyze the conversion of dopamine to o-dopaquinone, allowing dopamine to be indirectly detected by measuring the reduction of o-dopaquinone back to dopamine via a 2e − /2H + redox process at a potential around −0.1 V vs Ag/AgCl reference under physiological conditions.This potential window significantly reduces redox reactions of interfering substances, such as ascorbic acid and uric acid, 25 providing the mean to selectively detect dopamine in electrochemical biosensors.This approach addresses some of the selectivity issues of electrochemical measurements based on direct dopamine oxidation as the signal from other co-existing species becomes negligible.Another enzyme less commonly used is horseradish peroxidase (HRP), a heme protein that catalyzes the oxidation of catecholamine substrates such as dopamine and norepinephrine. 50,62,63An example of an HRP-based biosensor for dopamine detection was shown by immobilizing HRP in a nanocomposite of multiwall carbon nanotubes (MWCNTs), glycine, and silica sol-gel, enabling measurements of dopamine in the 15-865 μM range with good selectivity against uric acid, ascorbic acid, serotonin, tyrosine, serine, and cysteine. 50Figure 2 shows a schematic of dopamine oxidation by oxidase and peroxidase enzymes for most commonly used enzyme catalysts: tyrosinase, laccase, and HRP.
66][67]68 For example, Li et al. covalently immobilized laccase using a superparamagnetic core-shell Fe/Si support and demonstrate detection of dopamine onto a GCE. 53elatively few biosensors have been adapted to microsensor platforms that can be used in confined environments, cells cultures, tissues, or in implantable conditions. 30An example of a recent implantable dopamine biosensor is shown by Yin et al., using CFMEs with a three-dimensional chitosan-based porous hydrogel network. 69Enzymatic biosensors have several advantages as compared to methods based on direct electrochemical detection (e.g.FSCV with CFMEs); 1,3,34,70 (i) higher selectivity due to enzyme bio recognition and low applied potential, (ii) enzyme biosensors can be used in amperometric mode to measure time-dependent dopamine release in the brain without the need to recondition the electrode or perform background subtraction, and (iii) since the product of dopamine oxidation (o-dopaquinone) is recycled, the electrode is less prone to passivation due to the adsorption or polymerization of dopamine and its oxidation products.Challenges associated with these biosensors include the need to immobilize the enzyme on a very small surface, diffusion limitations, decreased response times due to the immobilized layers, and lowered rates of enzymatic reaction compared to direct dopamine oxidation.The goal of this review is to provide an overview of the recent progress made on the development of catalytic biosensors for dopamine and the challenges that need to be overcome to advance their use in neurophysiology, pharmacology and the clinical field.
Immobilization of enzymes for obtaining stable and sensitive biosensors.-Stableattachment of dopamine-specific enzymes onto an electrode surface is one of the most important steps in the development of a reliable enzyme biosensor for dopamine.This process strongly affects the performance of the biosensor in terms of sensitivity, stability, response time, and reproducibility. 71For applications in clinical samples, particularly those that require in vivo use, the main challenge is to design biocompatible coatings which allow stable attachment of the enzyme to the electrode surface while minimizing diffusion and maintaining the recognition functions and accessibility of their active sites for fast binding and catalytic conversion.Since enzymes require specific conditions to maintain their activity, the selection or appropriate materials and immobilization procedure is essential for ensuring efficient recognition, stability and detection sensitivity at electrode surfaces.
Various materials and immobilization strategies have been used to immobilize tyrosinase and laccase including physical adsorption, physical entrapment, electropolymerization, cross-linking, and covalent binding.The easiest method is to use a biocompatible polymer such as chitosan to entrap the enzyme within the biopolymer matrix by ionic or covalent cross-linking.This method was employed by Njagi et al., demonstrating stable attachment of tyrosinase on a carbon fiber-based microbiosensor and continuous operation in vivo for 2 h in amperometric mode with a storage stability of 2 weeks without loss of activity. 30When a silica sol-gel was used as a matrix, the microbiosensor had a lower sensitivity and higher response time (8 sec for chitosan vs 20 sec for sol-gel) due to higher diffusion though the more compact silica network, showing chitosan to be the better candidate for immobilization of enzymes in the construction of dopamine biosensors.Another benefit of using chitosan over other materials is its film-forming ability, ease-of-use, biocompatibility ECS Sensors Plus, 2024 3 020601 and porosity, making it an ideal choice for the effective immobilization of enzymes and the rapid diffusion of enzyme substrates.
Another way to immobilize receptor materials for dopamine biosensors is to use conductive polymers such as pyrrole and poly (3,4-ethylene dioxythiophene) (PEDOT) electropolymerized on the electrode surface. 72The method provides a favorable environment for the entrapment of tyrosinase and laccase, with controlled thickness and accessibility of substrates to the immobilized enzyme.However, the application of an electrical potential can inactivate enzyme activity, restricting the amount of enzyme to a monolayer thickness at the polymer/solution interface, and further limiting the activity of the immobilized enzyme on the electrode surface. 73Using a different polymer, Dhanjai et al. developed an implantable biosensor based on a polyvinyl acetate modified with polyaniline, cellulose nanocrystals, carbon nanotubes, and tyrosinase for detection of dopamine in brain samples of zebrafish as a stress biomarker.This sensor provided an LOD of 1.57 nM with a large linearity range of 7-1000 mM dopamine. 74Although the detection limit for this biosensor is relatively low, the linear working range does not begin until 7 mM, meaning that any dopamine measurements below this concentration will not follow a predictable trend, therefore rendering the electrode essentially ineffective for measurements within the low nanomolar range. 33ther immobilization methods involve covalent bonding using common cross-linking agents such as glutaraldehyde or carbodiimide via cross linking chemistry using N-ethyl-n'-(3-dimethylamino)propyl)carbodiimide/N-hydroxysuccinimide (EDC/NHS) activation, providing strong binding and increased stability for the enzyme. 75However, these coupling agents can be toxic, inactivating the enzyme and decreasing its activity.This can be challenging to overcome, particularly when these materials are immobilized on microelectrodes with very limited surface area and binding sites.Therefore, when comparing immobilization methods for these enzymes, entrapment methods in conductive polymers or biopolymers, such as chitosan, are more suitable for maintaining enzyme activity during the fabrication of enzyme microbiosensors.Examples of commonly used materials and methods for the immobilization of tyrosinase and laccase for the fabrication of electrochemical dopamine biosensors are shown in Fig. 3.
Improved biosensing designs involved immobilization of enzymes in conjunction with high surface area materials such as graphene, CNTs, metal and metal oxide nanoparticles such as those based on cerium, titania, or iron oxides (CeO 2 , TiO 2 , Fe 3 O 4 ).Carbon nanomaterials and nanostructured oxides provide increasing catalytic efficiency as well the ability to stabilize enzymes through surface adsorption, thus enhancing the overall sensitivity and stability of the biosensors. 72,76,77For example, using both tyrosinase and ceria-based oxide nanoparticles increased detection sensitivity as compared to biosensors prepared with the two materials separately.The enhanced sensitivity enabled in vivo measurements of dopamine.The metal oxides had a catalytic amplification effect and provided a sufficient oxygen supply for the tyrosinase, which requires oxygen as a cofactor for its activity.An additional benefit of co-immobilizing enzymes with cerium oxides is that they enable oxidase enzyme function in media with restricted oxygen, facilitating operation during implantation, 78 and allowing for dopamine detection in hypoxic environments, further expanding its application to oxygen restricted environments.The enzyme and nanoparticle layer was deposited within chitosan, which provided good film forming ability and biocompatibility enabling its use in vivo.Oxidation of dopamine by both the enzyme and the oxide layer is followed by the electrochemical detection of dopaquinone at an applied potential of −0.15 V vs Ag/AgCl; the low oxidation potential for dopaquinone eliminates common interferences, such as ascorbate, that are detected at the larger potential differences used to detect dopamine directly with carbon fiber electrodes.A permselective membrane was not required since there are few common interferents oxidized at this potential.Using the enzyme and a CeO 2 /TiO 2 oxide composite embedded in chitosan, the detection limit for dopamine was 1 nM with linear range spanning five orders of magnitude up to 220 μM and sensitivity of 14.2 nA/μM.There was no response to common interferences such as ascorbic acid and uric acid when these were added to the test solution.Interestingly, the sensor did not respond to serotonin, another catecholamine present in the brain; this can be attributed to the slower oxidation rate of tyrosinase for serotonin as compared to dopamine chitosan.The high sensitivity of the biosensor enabled it to be used for realtime in vivo measurements of dopamine released following electrical stimulation.Selectivity of the signal in vivo toward dopamine was demonstrated with nomifensine treatment. 30Performance of this microbiosensor and examples of in vivo recordings are shown in Fig. 4.
Table I summarizes electrochemical enzymatic biosensors reported in literature based on dopamine/dopaquinone conversion by oxidase and peroxidase enzymes, and provides design specifications including electrode materials, electrochemical detection techniques, and performance for measurements of dopamine in physiological fluids.ECS Sensors Plus, 2024 3 020601 Enzyme-mimetic catalysts for dopamine.-Asan alternative to biological catalysts, synthetic nanocatalysts that have the ability to catalytically oxidize dopamine has been explored as active materials for electrodes for dopamine detection.Enzyme-mimetic materials, or nanozymes, engineered at the nanoscale can provide enzyme-like activities and could in principle be used in lieu of natural enzymes.Examples of enzyme-mimetics include nanoscale materials such as metal (Au, Ag, Pt, Pd) and metal oxide (CeO 2 , TiO 2, Fe 2 O 3 , NiCo 2 O 4 ) nanoparticles, sulfides (FeS, MoS 2 , CuS), carbon-based composites functionalized with redox moieties, MXenes, and metal organic frameworks (MOFs). 107These materials can be custom-made with redox sites specifically designed to undergo oxido-reduction reactions with target analytes, including dopamine.For example, CeO 2 -based nanostructures have a variety of enzyme-like functions mimicking the activity of several enzymes including oxidase, peroxidase, catalase, superoxide dismutase, and phosphatase. 108The origin of the catalytic activity of CeO 2 is in the dual Ce 3+ /Ce 4+ surface valence states and high oxygen vacancies, enabling redox reactions similar to oxidase and peroxidase enzymes.It is known that CeO 2 has high reactivity for dopamine oxidation 109,110 with varying levels of sensitivity against other catecholamines such as serotonin and epinephrine. 111These indicate that nanoscale redox CeO 2 and composites of these materials can in principle be used to replace oxidase and peroxidase enzymes due to their ability to mimic the catalytic functions of natural enzymes, while overcoming drawbacks such as limited stability and sensitivity to environmental changes.Synthetic nanoenzymes are less expensive, have high thermal and pH stability, and can function in environments that are not accessible to natural enzymes, which could greatly reduce the cost of biosensors, increase stability and expand their applicability.However, although these materials are rapidly gaining popularity for sensing, their mechanism of action, selectivity for target molecules, and possible inhibitory effects in the measurement environment which can alter the catalytic efficiency are only beginning to be understood at the fundamental levels. 112The batch-to-batch variability of these materials could also affect reproducibility of these sensors.
A variety of enzyme-mimetic catalysis, also called "synthetic nanoenzymes" have been developed and incorporated into biosensing platforms.Several examples are presented in Fig. 5, depicting enzyme-mimetic materials, microelectrode platforms, calibration curves, and examples of dopamine measurements, including in vivo studies from recently reported work. 113,114Table II provides a list of recent enzyme-mimetic electrochemical biosensors for dopamine detection in physiological fluids, along with the enzyme-mimetic materials used, detection method, and performance of reported sensors.Tyr-tyrosinase, PPO-polyphenol oxidase, Lac-laccase, CNTs = -carbon nanotubes, GA-glutaraldehyde, PPY-polypyrrole, BSA-bovine serum albumin.

Measurements of Dopamine in Biologically Relevant Environments
Synthetic and pharmaceutical samples.-Characterization,optimization, and performance of dopamine biosensors is most often established on synthetic and pharmaceutical samples due to their simple composition and well-established concentration ranges that ease standardization.These experiments are typically performed in environments such as biological buffers (ex.0.1 M PBS) around pH 6.5-7.5.li et al. developed an electrochemical biosensor through covalent immobilization of laccase onto a superparamagnetic core--shell Fe/Si support on a GCE in 0.1 M PBS (pH 6).This proof-ofconcept work demonstrates the ability of these materials to efficiently detect dopamine.The core-shell structure provided a good microenvironment for preserving enzyme activity and facilitated fast electron transport, allowing for detection of dopamine in the 1.5--75 μM range with an LOD of 177 nM DA. 53 While the linear range of this sensor is limited, the concept of the core-shell structure demonstrates potential for achieving enhanced enzyme-like activity.Laranjo et al. immobilized tyrosinase in 5% chitosan onto a silica/ titania graphite ceramic composite achieving detection in the 40--350 μM range via CV and amperometry, 96 which enable real-time measurements.Utilizing a gold-doped La 2 O 3 nanocomposite on an indium-tin-oxide electrode, Srivastava et al. reported a lower range of 2-100 μM with an LOD of 258 nM dopamine. 90Decarli et al. demonstrated detection in pharmaceutical samples via SWV, CV, and EIS using a laccase-based carbon paste electrode modified with zwitterionic surfactants and halloysite nanotubes achieving a linear range and LOD of 0.99-67.8μM and 252 nM DA, respectfully 91 Pimpilova et al. reported an LOD of 37 nM DA through use of goldmodified GCE with cross-linked laccase via self-assembled cystamine, depicting the benefits of gold to increasing biosensing performance.However, the large electrode size renders the electrode unusable for measurements in confined environments such as those in in vivo conditions or cell culture environments. 98Moreover, properties of these materials and their electrochemical responses may vary based on electrode design.For example, a procedure that allows for sensitive and reproducible results on a GCE may yield poor results in terms of signal-to-noise ratios and reproducibility when applied on a CFME due to changes in electrode surface area, hydrophobicity, capacitive current, and inhomogeneous coating of the electrode surface with the active layers.
Other recent works concentrate on the application of enzymemimetics.Since these sensors do not involve the use of a biological receptor, they tend to be more stable and less expensive, but may lack selectivity and sensitivity.Olejnik et al. fabricated a dopamine biosensor through chemical vapor deposition of boron-doped carbon ECS Sensors Plus, 2024 3 020601 ECS Sensors Plus, 2024 3 020601 nanowalls functionalized with an electropolymerized polydopamine/ polyzwitterion complex in which the zwitterion complex functioned as an enzyme-mimetic catalyst for dopamine.Mechanistically, the catalytic activity of this material was attributed to intermolecular hydrogen bond, cation-π, and π-π interactions facilitating the transport of dopamine to the electrode surface and creating a molecular pocket where the transition states between dopamine and its quinone derivative are stabilized, causing the oxidation of dopamine to become thermodynamically favoured, further enhancing the sensitivity of the biosensor.Moreover, the incorporation of zwitterions provided antifouling functions, making this a good candidate for preventing non-specific adsorption and achieving increased accuracy in biological environments.This biosensor was characterized by an LOD of 89 nM with a linear range of 0.7--20 μM, comparable to some reported enzymatic biosensors. 125ome of the most common enzyme-mimetic materials for dopamine oxidation involve the use of single or mixed metal oxides.For instance, Wi et al. used MnFe 2 O 4 /MoS 2 as an oxidase-mimetic to design a dopamine biosensor on a SPCE platform, enabling multianalyte detection of dopamine, uric acid, and ascorbic acid with LODs of 410 nM, 140 nM, and 170 nM, and linear ranges of 1-100 μM, 1-100 μM, and 200-1,000 μM, respectively. 155Gutoiu et al. used a manganese-doped crystalline copper oxide drop-coated onto a SPCE achieving an LOD of 30.3 nM and linear ranges of 0.1-1 and 1-100 μM dopamine.The biosensor demonstrated a fivetimes amplification of signal with the use of these enzyme-mimetics when compared to the bare SPCE. 116Having a low nanomolar LOD and effective working range from 0.1-100 μM, this oxidase-like material offers a promising substitute for enzymes when detecting dopamine in non-invasive ways.Oxides derived from sustainable sources have also been reported, such as a recently extracted carbon catalyst derived from Parthenium hysterophorus leaves.The electrochemical sensors built with this material achieved an LOD of 600 nM and linear range of 0-10 μM DA. 35 Although the sensor does not show impressive performance, such green-derived materials can be modified with catalytic materials for enhanced performance for dopamine measurements.
Nanostructured hybrids and composite materials, typically consisting of high surface area catalysts incorporated within or functionalized with conductive materials have shown enhanced performance for dopamine oxidation as compared to materials based on single components.For example, high oxidase-like activity was achieved with the use of a hybrid nanozyme of cobalt/nitrogen Codoped carbon nanotubes and nanosheets deposited on a GCE along with a metal organic framework (MOF) (ZIF-L@ZIF-67), resulting in an LOD of 9 nM and linear range between 0.03-710 μM measured by CV, EIS, and amperometry.The sensor showed good performance in synthetic cerebrospinal fluid and serum samples, further justifying this hybrid nanozyme as a suitable candidate for implementation on more complex electrode designs, such as implantables. 133In another design, Zhang et al. used a conductive Ti 3 C 2 T x MXene with high surface area deposited on a GCE with electrochemically reduced graphene, providing significantly enhanced electrochemical performance while also preventing electrode biofouling.The sensor exhibited high sensitivity with an LOD of reaching an LOD of 41 nM and a linear range of 1-40 μM.Lu et al. 164 deposited WO 3 -SnO 2 nanoflakes onto a fluorine-doped tin oxide platform, achieving an very low LOD of 0.8 nM and significant linear range of 5 nM-1.75 μM in PBS.Rizalputri et al. 165 deposited gold nanobipyramids on a SPCE, achieving an LOD of 37 nM and an increased linear range of 0.1-100 μM.Although this design allows for a wider working range, the biosensor lacks the ability to measure dopamine reliably at concentrations lower than 100 nM.
To reduce sample volume to less than 1 microliter and increase applicability, Nuh et al. 166 developed a microfluidic platform with an integrated SPCE electrode functionalized with OMC-PEDOT-PSS, enabling detection of dopamine with an LOD of 21.6 nM and a linearity range up to 10 μM,, allowing measurement in a very small sample volume.To increase reproducibility and ease of manufacturing of biosensors, a recent trend uses 3D printing to fabricate electrodes.Wan et al. demonstrated the 3D printing of nanocarbon electrodes made of MXene-Quantum dots (MXene-QDots) with a diameter of 6 mm as electrochemical sensors for dopamine.The sensor provided an LOD of 3 nM and a linear range of 0.01-20 μM (Figs.6C, 6D). 167Applicability was demonstrated in commercial dopamine hydrochloride injection, proving that 3D printed electrodes are a viable platform for dopamine biosensors with high sensitivity and appropriate working range for pharmacological applications.
Measurements in biological fluids.-Demonstration of sensing capabilities in real samples and biological fluids is critical to establish the potential of this technology as a detection tool in biology and medicine.The best performing biosensors currently in literature are those based on multi-layered materials such as MXenes, graphene, or carbon-based composites decorated with single atom catalysts or metal/metal oxide nanostructures.For example, an MXene, Ti 3 C 2 T x , demonstrated enzyme-mimetic properties for dopamine oxidation with the ability to detect dopamine in serum when deposited on a GCE using Nafion, producing an LOD of 3 nM and linear range of 0.015-10 μM. 121In another work, a carbon cloth functionalized with aminated Nb 2 CT x MXene and MXene-MoS 2 heterostructures (Nb 2 CT x @MoS 2 ) obtained by modification with ethylene diamine (EDA) demonstrated superior performance as a biomimetic material for dopamine detection in human serum samples, giving an LOD of 0.3 nM and a linear range of 0.001-100 μM, and down to 0.23 fM LOD when Nb 2 CT x @MoS 2 was used (Fig. 7A). 122,168ugo et al. 169 reported a molecularly imprinted microneedle array coated with PDMS/CNT/CNC for multiplexed detection of dopamine, pH, epinephrine, and lactate in human sweat (Fig. 7B).The LOD for dopamine detection was 2.1 nM with a linear range of 73.9-763 nM, demonstrating the potential of microneedle technology for wearable sensors.In spite of this, the linear range fails to cover the useful dopamine concentration range in biological fluids.Liu et al. 170 demonstrated amperometric detection in real human blood serum using an SPCE coated with a gold nanoparticlepeptide hydrogel with an LOD of 0.12 nM and linear range of 0.0002-1.9μM, showing potential for clinical testing ex vivo.Other materials such as Dy 2 MoO 6 nanozyme 150 and barium zirconate perovskite crystallite 140 also show oxidase-mimetic properties and the ability to detect dopamine in real human urine and blood serum, with an LOD of 8.3 nM, linear range of 20-340 μM, 150 and 5 pM with linearity up to 50 μM, 140 respectively, showing promise for use as clinical biosensors.
A graphene oxide-based biosensor with self-assembled FeTBPP and p-tert-butyl cup octaaromatic hydrocarbons on an indium tin oxide platform designed by Gao et al. provided detection capabilities in the 0.1-1,000 μM range and functionality in real human urine and serum samples. 135Fu et al. demonstrated the use of hierarchically porous carbon functionalized with iron disulfide as an enzymemimetic for dopamine.The sensor developed on a GCE electrode demonstrated an LOD of 15 nM with linearity range of 0.05-1,000 μM, and good functionality in human urine as tested through EIS, CV, and DPV.Due to the increased surface area of the porous carbon, this platform allowed for efficient loading of the catalytic material, enhancing the oxidase-like activity and ability to sensitively detect dopamine within human biological samples. 157afique et al. demonstrated the use of Ag-chitosan nanoparticles on a GCE for dopamine detection in urine, with an LOD of 10 nM and an applicable linear working range of 0.1-60 μM measured by DPV. 137This work demonstrates the sensitivity achieved through DPV measurements.
Other designs incorporate oxidase-mimetics based on copper nanocubes grown on top of multi-layered graphene deposited on SPCEs (Cu-Gr/SPCE), as seen in the work by Dinu et al. (Fig. 7C).The high surface area-to-volume ratio and enhanced catalytic activity provided increased sensitivity to the enzyme-mimetic sensor, giving an of LOD of 0.33 nM and a wide linear range (0.001-100 μM) with the capability to measure dopamine in human serum. 136Besides the copper nanocubes, the materials for this sensor are fairly simple and cost-effective, making this a promising candidate for application in human serum samples.The two main characterization methods used for this design included EIS and CV, with characterization of dopamine oxidation via the sensors response using DPV.By using the oxidase-like activity of boron-nitrogen Codoped carbon nanotubes, Liao et al. achieved selective detection of dopamine and uric acid without interferences from each other or other molecules.This GCE-based sensor was only able to detect down to 0.8 μM in a linear range of 1-630 μM dopamine. 147Lu et al. used reduced graphene oxide composites of CoO NPs and Ndoped carbon sheets on a GCE, providing an LOD of 150 nM and a linear range of 0.5-110 μM. 159Another cobalt-based nanozyme with a bamboo-like nitrogen-doped carbon nanotube structure demonstrated an LOD of 34.2 nM and linear range of 0.5-150 μM DA, and was successfully applied for detection of dopamine in human serum. 156Ma et al. was able to further increase their sensors detection capabilities by using bifunctional nanozymatic hollow core-shell CuS@CuSe heteromicrocubes leading to an LOD of 0.38 nM and linear ranges of 0.002-0.44 and 0.52-1.66μM DA. 162 An more sensitive sensor was described by Sun et al. through the use of nitrogen-doped graphene with dispersed iron sites, reaching an LOD of 27 pM.However, this biosensor had a narrow linear range spanning from 50 pM to 15 nM in human serum, severely limiting its abilities. 160ome materials recently found to deliver high sensitivity include single atom catalysts deposited on carbon supports.Xie et al. achieved simultaneous and selective detection of dopamine and uric acid using single-atom ruthenium on carbon nitride supports (Ru-Ala-C 3 N 4 ).The use of single atom catalysts enabled detection of as low as 20 nM with a linearity range between 0.06-490 μM in real human serum samples due to the enhanced surface area-tovolume ratio allowing for optimal electrocatalytic activity and oxidase-like properties for dopamine detection.The single-atom catalyst shows to be effective in dopamine oxidation, giving a very wide linear range covering the majority of dopamine concentrations in fluids. 143Bushira et al. used a nitrogen-doped single iron atom catalyst (Fe-N5) on a GCE for detection of dopamine in human blood serum, reaching an LOD of 7 pM and a linear range spanning from 0.005-500 μM with great stability.This sensor was also able to selectively detect uric acid with an LOD 27 pM and linear range 0.01-480 μM with potential for many biomedical applications. 117ncorporating a more complex electrode, Ohshiro et al. reported an extended-gate OFET biosensor with a laccase-linked N-ethylphenazonium moiety mediator for detection of dopamine in real urine samples with an LOD of 190 nM and linear range from 0-4 μM (Fig. 7D), showing the materials applicability to other platforms. 106urther detection of dopamine in cerebrospinal fluid and blood in a mouse model of Parkinson's disease was demonstrated with a microfluidic electrochemical device incorporating a gold electrode.The microfluidic platform enabled measurements in ultralow amounts of samples of 2.4 μl with an LOD of 0.1 nM and working linear range of 0.001-1 μM in samples collected from parkinsonian rats. 171ther unique designs use a combination of enzyme-mimetic catalysts and enzymes such as laccase or tyrosinase, to synergistically amplify dopamine oxidation through both chemical and biological means.An example of this can be seen from Swetha et al. 97 who used a Donor-Acceptor-Donor molecule BTCN via π-π stacking co-immobilized with tyrosinase on a GCE electrode, reaching an LOD of 0.47 nM and a linear range from 0.001-500 μM, demonstrating the capabilities of the sensor in serum samples from healthy individuals.The high sensitivity was explained by the in situ formation of a complex between the BTCN molecule and the copper center of the enzyme along with excellent charge transfer provided by the tunable holes in BTCN.In another design, multi-segment 1D Ni/Au/PPY-COOH nanowires co-immobilized with Trametes versicolor laccase within Nafion on an SPCE showed great selectivity against ascorbic acid, uric acid, L-Cysteine, serotonin, and glucose with the ability to maintain its sensitivity and reusability after longterm storage.This work depicts the enhancement of dopamine detection when combining enzymes and oxidase-like materials on readily available electrodes. 93asurements in cell-lines.-Althoughtesting in biological fluids is an important way to provide non-invasive dopamine detection, cell lines are one step closer to working with organs and larger biological systems that facilitate neurotransmitter monitoring at the production sites.Huang et al. demonstrated subnanomolar detection of dopamine in PC12 cells using a GCE coated with highly ordered mesoporous iron oxide (Fe 3 O 4 −40) on a hard silica KIT-6 template.This configuration provided enzyme-mimetic properties due to the high surface area and large pore volume, imparting good sensitivity to the biosensor for an LOD of 0.8 nM and linear range of 0.002-0.6μM. 120Shu et al. measured dopamine along with H 2 O 2 in PC12 cells using an N-doped carbon decorated with isolated cobalt atoms (Co-N x sites) for an LOD of 40 nM and a linearity range of 0.06-1,200 μM at 0.4 V vs Ag/AgCl electrode (Fig. 8).This nanozyme (Co-N-C-800) demonstrates excellent stability and intrinsic peroxidase-and oxidase-like activities due to its Co-N x sites similar to that of natural enzymes, making it a great biomimetic catalyst for dopamine detection. 154Zou et al. utilized Co 3 O 4 and nitrogen-doped graphene aerogel deposited on the surface of a GCE for real-time monitoring of secreted dopamine from living PC12 cell lines, with an LOD of 12 nM and linear working ranges of 0.05-200 and 200-1,500 μM.A similar design, consisting of cobalt oxyhyroxide and carbon black nanoflake nanocomposites demonstrated oxidase-like activity and the capability to detect dopamine from cell-secretion with an LOD of 660 nM and linear range from 1-950 μM. 152Shetty et al. measured dopamine secretion from SH-SY5Y neuroblastoma cell lines using a polyoxometalate-ϒ-cyclodextran MOF nanozyme deposited on a laser-scribed graphene electrode.The sensor had an LOD of 10 nM and a linear range of 0.05-1,000 μM DA.In addition to electrochemical detection, this platform also demonstrated capabilities as a user-friendly colorimetric read-out that can be used as a point-ofcare device. 134evisiting the use of single-atom catalysts for biosensing design, Shetty et al. demonstrated the use of Fe single atom catalyst along with the Ti 3 C 2 T x for improving sensitivity of dopamine detection ECS Sensors Plus, 2024 3 020601 with an LOD of 1 nM and wide linear range of 0.01-200 μM, enabling square wave voltammetry measurements (SWV) measurements in urine and PC12 cells (Fig. 8). 132Xu et al. 172 fabricated an electrode array with microneedle decorated with 4-MPBA, ACBK, and gold nanoparticles for DA measurement in PC12 cell lines.The biosensor demonstrated an LOD of 140 nM and a linear response from 0.5-1 μM.Nekoueian et al. 173 developed a sensing platform based on tetrahedral amorphous carbon thin films and PECVDgrown carbon nanofibers and molecularly imprinted polypyrrole as an artificial dopamine receptor.The sensor enabled dopamine detection in F-12 K cell medium with an LOD of 5.43 nM DA and a working range of 0.01-10 μM.
Recent work demonstrates detection of dopamine secretion in low volumes in cultures confined within fluidic channels through integration of electrochemical sensors within microfluidic platforms.To illustrate, Li et al. 174 developed an integrated microfluidic sensor with a chip-like microelectrode array decorated with a nanocrystalline indium phosphate interfaces and PANI to detect dopamine released from neuro-2A cells (Fig. 8).Detection was achieved via the crystalline structure of an organic-inorganic hybrid molecule made of indium atoms, phosphate ligands, and 1,2,4,5-tetrakis (imidazol-1-ylmethyl)benzene linkers (NTOU-7).The sensor enabled detection of dopamine in concentrations as low as 0.183 aM, the lowest reported sensitivity for sensors of its kind.Nonetheless, this sensor was only capable of a working range from 1 aM to 10 pM, severely limiting the application of this biosensor to extremely low concentrations.Sen and Avci 175 developed a redox-cycling microfluidic device to test dopamine released from PC12 cell lines using indium tin oxide electrodes coated with gold nanoparticles.These works demonstrate the potential of on-chip electrochemical detection of dopamine released from cells which can be applicable to a wide range of in vitro studies to investigate the response of dopaminergic cells and dopamine neurotransmission.
In vivo detection.-Real-timedetection of dopamine within the brain can provide understanding of dopamine signaling and insight into brain activity and regulation, helping to better understand the mechanisms of dopamine within the brain and how it relates to neurodegenerative diseases such as Parkinson's.In vivo detection of dopamine can be realized using microdialysis, imaging, and electrochemistry, 176 with microanalysis probes being relatively large and causing damage to surrounding tissues with a set-up that is not readily adaptable to real-time monitoring.Electrochemistry with implanted electrodes can provide detection of dopamine concentrations in real-time (Fig. 9A).Most reported electrochemical methods rely on the use of FSCV through direct measurements of dopamine oxidation with implanted CFMEs, which allows measurements of dopamine uptake for concentrations down to 100 nM with sub-second resolution. 1,39This section discusses the potential of enzyme and enzyme-mimetic biosensors for the in vivo detection of dopamine.The most common examples of in vivo dopamine detection involve modification of CFMEs with electrocatalytic and enzymemimetic materials and coatings to enhance dopamine electrooxidation and improve detection sensitivity and selectivity.Simultaneous detection of dopamine and serotonin release over time in the striatum of mice was demonstrated using a CFME modified with diazonium salt and single-walled carbon nanotubes (Fig. 9B), providing an LOD of 27.3 nM and respectable linear range of 0.4-5 μM. 153urthermore, modification of CFMEs with a conductive MOF and gold nanoleaves was shown to enhance the electron transfer from dopamine to the electrode surface due to the rich electrocatalytic sites provided by the conjugated molecular wires of 4-(thiophen-3ylethynyl)-benzaldehyde (RP1) and the MOF, decreasing the Gibbs free energy of dopamine oxidation and facilitating sensitive detection, as depicted in Fig. 5A.The LOD of this sensor was 1 nM and the linear range spanned over 0.004-0.4μM dopamine measured by CV, DPV, and FSCV, proving the versatility of the sensor.The electrode (Au/RP1/Ni 3 HHTP 2 /CFME) was tested on both acute and subacute parkinsonian mice, demonstrating its capabilities for detecting dopamine at trace levels within the brain and characterizing dopamine release in conditions associated with Parkinson's disease. 113A slightly different approach was taken by Yuen et al. by studying dopamine release in rat brains through onset of cocaine stimulation via FSCV and multi-cyclic SWV using PEDOT immobilized via Nafion onto a CFME.The basal tonic levels of DA within the nucleus accumbens core of Sprague Dawley rats was found to be 134 ± 32 nM (Fig. 9C).The small size of the microelectrode minimized damage to the surrounding tissue, keeping the dopaminergic neurons intact and allowing them to produce more dopamine than previously reported in literature via microdialysis.This study highlights the critical effect of electrode size and coatings to ensure accuracy in measurements. 128In another work, Drosophila (fruit flies) were used as the animal model for in vivo detection of dopamine via CV using CFMEs functionalized with enzymemimetic Cu 2 S/rGO, achieving an LOD 24 nM and working range of 0.1-20 μM, successfully monitoring dopamine over time. 129The Cu 2 S/rGO provided catalytic activity and enhanced selectivity of the CFMEs for dopamine.
Few enzyme biosensors have been implanted and used to measure dopamine in vivo.Real-time in vivo monitoring was demonstrated with a tyrosinase-modified multi-carbon fiber electrode that combined both enzyme and enzyme-mimetic activity, achieving detection sensitivity down to 1 nM and high selectivity towards dopamine through combination of CeO 2 and TiO 2 nanoparticles paired with tyrosinase. 30The biosensor was implanted in the striatum of rats to measure dopamine release upon electrostimulation and depletion upon administration of nomifensine, as shown in Fig. 4C.Since nomifensine is specific towards deactivating dopamine production, this step can provide further confirmation of dopamine detection.In another work, non-CFME enzymatic in vivo detection of dopamine was demonstrated in fish brain with a tyrosinase and gold-coated iron oxide magnetic nanoparticles (Tyrosinase/Fe 3 O 4 @AuNPs) deposited on a conductive gold microelectrode chip based on field-effect transistor (FET).This biosensor was capable of detecting dopamine levels down to 3.3 nM with an uncompetitive linearity range of 1-120 μM, demonstrating functionality in fish brains. 84These works illustrate the potential of enzyme biosensors for real-time measurements of dopamine in vivo.
In recent years, efforts have concentrated on developing miniaturized microelectrode platforms, e.g.implantable microchips, which are small enough to be implanted in vivo and be used as an additional alternative to CFMEs.Several examples of microarray sensors and demonstration of their applicability in vivo are illustrated in Fig. 10.Zhang et al. reported a PtNPs/Nafion functionalized silicon-based microelectrode array fabricated on a 16-site array using micro-electrochemical systems (MEMS) technology for dual mode electrophysiological and electrochemical recordings of dopamine in parkinsonian primates (Fig. 10A).The fabricated probe was 25 × 300 μm long, which enabled implantation.The microelectrode array had a sensitivity of 13.8 pAμM −1 and was able to quantify differences in dopamine release in a normal (332.6 μM) versus parkinsonian (35.2 μM) monkey, explained by a degeneration of dopaminergic neurons of the substantia nigra, proving its versatility for clinical in vivo scenarios. 177Liu et al. monitored dopamine in freely-moving mice using PEDOT: PSS-coated diamond films deposited atop a micro-light emitting diode platform (Fig. 10B).This closed-loop system provided an integrated system enabling data acquisition, real-time detection, and remote signal control of the dopamine release in the ventral tegmental area where most of the dopamine in the brain is produced.An LOD of 100 nM and a linear range of 0-10 μM was obtained, demonstrating potential of this platform for sensing in freely-moving mice.However, limitations of this sensor include its somewhat narrow linear range, preventing the sensor from providing accurate information above 10 μM. 119Stuart et al. developed a wireless battery-free biosensor with a multilayered architecture small enough to be implanted consisting of SWCNTs/ carbon fiber composite electrodes and a wireless battery-free electronic circuit with a microscale light emitting diode (μLED) ECS Sensors Plus, 2024 3 020601 (Fig. 10C). 178The sensor was fabricated by laser patterning, lithography, and thin film deposition on a flexible electronic platform with an overall size of 12 × 8.5 × 3.2 mm 3 and a weight of 49 mg, providing reproducible sensors.Additionally, its sensitivity of 1264.1 μM −1 cm −2 demonstrated the sensor's ability to measure dopamine during subdermal implantation in freely-moving mice upon stimulation.
Xiao et al. 114 reported a platinum/reduced graphene oxide coated microelectrode array (PtrGO) grown on a silicon-on-insulator substrate for electrochemical analysis within living brains of parkinsonian rats.CV, FSCV, and amperometry were used to characterize the electrode prior to implantation, giving an LOD of 50 nM and linear range of 0-20 μM, making this a suitable candidate for real-time neurotransmitter detection in vivo.This study confirms the critical role of signal amplification, oxidase-like specificity, and electrocatalytic behavior of enzyme-mimetic materials for deep brain stimulation within the globus pallidus internal of rat brains for determining dopamine concentration in vivo.
A critical challenge of in vivo measurements with locally implanted electrochemical probes is electrode passivation due to adsorption of proteins or binding of the electro-oxidized dopamine to the electrode, preventing electrode transfer and inactivating the electrochemically active area of the electrode.An innovative design integrated concepts from FSCV with the sensitivity of organic electrochemical transistors (OECT), leading to the development of fast-scanning potential-gated organic electrochemical transistors.In this configuration, shown in Fig. 11A, fast-scanning potential is used as a gating mode with transconductance used as a sensing parameter for dopamine.The new technology was applied to monitor electrically stimulated dopamine release in a living rat brain and was characterized by an LOD of 5 nM with sensitivity of 0.899 SM −1 . 38in et al. developed a strategy to prevent passivation of CFMEs (Fig. 11B) by functionalizing CFMEs with a three-dimensional chitosan-based porous hydrogel network, helping to retain water and provide a biocompatible interface for confining enzymes, while ensuring rapid mass transfers of analytes, such as dopamine, through the porous network. 69The hydrogel was used to construct implantable enzymatic biosensors where the enzyme and MWNTs were immobilized within the chitosan network and deposited on PtNPsmodified CFMEs.This design demonstrated antifouling behavior and excellent anti-passivation properties upon implantation in rat brains.Chitosan is also known to reject ascorbic acid interferences, further increasing the biosensors affinity for dopamine through elimination of interferences. 179Chitosan-based electrode coatings are excellent matrices for immobilizing enzymes and mimetic catalysis, preventing passivation and improving selectivity of implanted biosensors.
An example of recent development to improve sensitivity of electrochemical biosensors for real-time in vivo detection of dopamine is to integrate electrochemical MEs into a 3D printed microfluidic syringe-like device (μSyringe).The work employed a microfluidic device using a graphite pencil electrode as a detection platform to provide an LOD of 0.1 nM and linear range of 0.1 to 500 nM, enabling detection of dopamine in freely-moving mice before and after treatment with amphetamine (Fig. 12A). 180Other miniaturized microfluidic designs were developed with electrodes incorporated in a PDMS chamber, enabling measurements of dopamine levels in volumes as low as 2.4 μl with detectable ranges down to 240 amol measured by amperometry at an applied potential of 0.2 V.This platform has been used to measure real-time dopamine release in cerebrospinal fluid and plasma of a parkinsonian mouse following administration of L-3,4-dihydroxylphenylalanine (Fig. 12  B). 171Another interesting direction is to fabricate support-free PEDOT-PSS fibers and use them as a bioelectronics interface for neurochemical monitoring.The fibers exhibited high conductivity, biocompatibility, and fast electron transfer, enabling electrophysiological recording in the extracellular space for both chemical and electrophysiological recording. 181This design could further be modified with enzymes and nano-enzyme-mimetics for implantable biosensors.

Challenges to Overcome for Advancing Translation of Dopamine Biosensors from Lab to Clinic
The current development status of electrochemical dopamine biosensors demonstrate many unique capabilities for measurements in a variety of samples ranging from biological fluids, e.g.blood, saliva, to tissues and whole animals.Their small size and high sensitivity are suitable for mechanistic studies to elucidate questions related to the dynamics, concentration and spatial localization of dopamine.Larger electrodes can be used to measure dopamine in biological fluids for disease diagnosis and monitoring progression of treatment.However, despite progress, the application of electrochemical biosensors is still restricted to measurements in standard laboratory conditions. 182Progress has been slow in the translation of this technology from labs to clinic.The transition from standard solutions to in vivo conditions is often difficult due to the complexity of the environment, potential damage during implantation, size considerations for microelectrode design, and difficulty with calibration.Several technical challenges remain to be overcome to ensure their practicality and utility for integration into healthcare.Several of these challenges are listed below: • Development of standardized protocols, calibration and accuracy: Ensuring accuracy is of paramount importance in the medical field and most biomedical devices require regular calibration and maintenance.Ensuring accuracy of measurements and implementing calibration procedures, potentially with the use of self-referencing probes should be an integral part of the sensor development process.Inter-laboratory studies and development of standardized protocols and performance metrics, and collaboration with medical professionals, manufacturers and end-users is vital for moving these biosensors from bench to market.
• Long-term stability and reliability.Biosensors intended to be used in the medical field must exhibit long-term stability and reliability to provide accurate measurements over extended periods of time.This is particularly important for enzyme sensors that require enzyme immobilization.Dopamine-specific enzymes such as tyrosinase and laccase are in general stable and some studies show that the use of biopolymers like chitosan and alginate can extend stability of the enzyme in immobilized conditions to over 250 days at room temperature. 183Procedures used for the widely used commercial glucose meter with glucose oxidase could be adapted to develop standardized protocols for the preparation of dopamine biosensors.
• Sensitivity and specificity in complex biological samples.
Many catalytic biosensors for dopamine have reported improved selectivity through the use of enzymes and custom-made nanoenzyme mimetics as compared to direct electrochemical approaches for dopamine detection.Although the selectivity is promising, many of the reported configurations do not reach sensitivity levels that match the physiological-relevant range of dopamine concentrations associated with normal and pathological conditions, such as Parkinson's disease.Other challenges are to ensure fast diffusion though the immobilized layer and prevent biofouling.Ensuring sensitivity while maintaining accuracy in the presence of interfering substances is especially critical for clinical monitoring and diagnosis.Improving the selectivity, and developing miniaturized and multi-array sensors that can simultaneously quantify dopamine in addition to other clinically important analytes could significantly expand applicability and accelerate adoption into practice.
• Miniaturization and integration.Most reported electrochemical biosensors for dopamine are performed on electrodes that are fabricated one at a time with analysis done on research grade instrumentation.Moving this technology from bench to clinics require further improvements in the manufacturing of these probes to produce reproducible sensors that can be used at large scale.Optimizing detection on portable, wearable and wireless controlled miniaturized potentiostats would ensure seamless integration into healthcare settings.
• Data management, interpretation and standardization.The large data set generated with the biosensors require efficient management and interpretation to extract meaningful clinical information.The integration of machine learning techniques to deconvolute signals, simplifying calibration with pre-recorded calibration data (assuming no calibration drift and stable sensors), employing standardized procedures in accordance to quality control/ quality assurance (QC/QA) criteria 184,185 and developing userfriendly interfaces for real time analysis would speed-up translation into the existing healthcare infrastructure.
• Regulatory compliances and approvals.Sensors intended for clinical use are subject to regulatory approvals by organization such as Federal Drug Administration (FDA) and sensors intended for this market segment should follow guidelines for method performance requirements.
Interdisciplinary collaboration between scientists, analytical and biosensors experts, biochemists, engineers, clinicians, and stakeholders across the healthcare ecosystem will be essential in future research for addressing these challenges and advance the development of these devices for achieving improved measurements, realtime detection and enhanced health outcomes.

Future Trends
Enzyme-based electrochemical biosensors, such as the widely used glucose meter, have been successfully implemented and commonly used in the clinical field.With the commercial success of the glucose meter, there exist opportunities to expand the use of enzymatic biosensors to other targets and health conditions with improved sensors.Enzymatic and enzyme-mimetic biosensors for ECS Sensors Plus, 2024 3 020601 dopamine described in this work demonstrate potential for dopamine detection.When carefully designed, these probes can be used to obtain fundamental and practical understanding of the role of the neurotransmitter dopamine on physiology, metabolism, and disease states in living systems.Such methods are of great importance in neurological and biomedical research and can be used in many applications to provide rapid dopamine measurements, understand dopaminergic function, or monitor real-time response following treatment with drugs or electrostimulation.
While enzyme-based electrochemical sensors for dopamine have demonstrated excellent analytical capabilities, their implementation as routine measurement tools requires further refinement to enhance accuracy, selectivity, and stability, particularly if they are designed to measure dopamine in vivo.To address the remaining challenges such as difficulty with calibration, accuracy of measurements, stability, electrode biofouling and reducing risk of potential damage upon implantation, future research is required.We anticipate future development in the following areas: (1) developing novel electrode materials and biocompatible coatings to enhance stability of the immobilized enzyme and maintain the activity and electrocatalytic functions of surface-immobilized biomimetic catalysts, (2) implementing advanced manufacturing concepts for achieving scalable and more reproducible fabrication of microelectrode platforms, such as 3D printing, (3) miniaturization and interfacing with microfluidics, with designs that follows a "lab on a chip" or "lab-in-a syringe" concept, as illustrated recently, 171,175,180 with one example of measurements in freely-moving mice, 180 (3) integration with portable instrumentation, (4) interfacing with machine learning and new developments in AI methods in order to discriminate dopamine signals from interferences such as ascorbic acid and uric acid for improved dopamine detection. 186Finally, (4) in vivo studies with implanted enzyme-based microbiosensors should be validated with pharmacological manipulations to demonstrate accuracy of measurements.These advancements and addressing the remaining technical challenges could enhance performance, improve practicality and accelerate translation of electrochemical dopamine sensors into the healthcare infrastructure.

Conclusions
The sensitive and accurate detection of dopamine is essential in in neurophysiology, pharmacology, and the clinical field.This review critically discussed the fabrication, applicability and limitations of recently developed electrochemical biosensors based on enzyme and enzyme-mimetic catalysts.Many advances have been made in recent years demonstrating the potential of these biosensors for achieving rapid real-time detection of dopamine with high sensitivity and selectivity in a variety of biological systems.Despite progress, working range, stability and accuracy of these devices still need great improvement in order to improve performance for use in the biomedical field.Moreover, most of these biosensors have only been tested on standard laboratory samples rather than in real samples.Thus, this issue needs to be addressed in future research activities as efforts shift to further demonstrating utility of these devices in more realistic samples and conditions.Establishing capabilities of enzymatic and enzyme-mimetic biosensors in real clinical settings, measuring dopamine in tissues and whole animals, and validating performance is important to fully take advantage of this technology.Once these capabilities are achieved, these techniques can find a myriad of uses in neurophysiology in biomedical research.

Figure 1 .
Figure1.Schematic of general biosensing design using dopamine-specific enzymes and enzyme-mimetic catalysts for the electrochemical detection of dopamine.

Figure 2 .
Figure 2. Examples of enzymes for dopamine biosensors showing enzyme-catalyzed oxidation reactions of dopamine into dopaquinone using (A) tyrosinase, (B) laccase, and (C) horseradish peroxidase (HRP).The reduction potential of dopaquinone to dopamine is about −0.1 V vs Ag/AgCl reference electrode.

Figure 3 .
Figure 3. (A) Examples of electrode materials commonly used for the immobilization of tyrosinase and laccase enzymes for the construction of electrochemical dopamine biosensors; not that these can be used alone or in composite form containing usually enzyme, a polymeric linker, and a high surface area nanostructure.(B) Summary of immobilization methods showing conceptual design for adsorption, cross-linking, entrapment, and covalent coupling.

Figure 4 .
Figure 4. (A) Comparative calibration curves of Tyrosinase-based CeO 2 /TiO 2 /Chit/carbon fiber and Chit/carbon fiber biosensors (n = 3 measurements) showing significant increase in biosensor response due to the synergistic effect of mixed oxides.(B) Amperometric responses obtained with the Tyrosinase biosensor to successive additions of 5 mM dopamine (DA).The amperogram also shows the selectivity to ascorbic acid (AA), uric acid (UA), serotonin (SR), L-dopamine (L-DOPA), norepinephrine (NorEP) and epinephrine (EP); (C) Immunohistochemical images of the brain slices for the stimulating and recording sites.The stimulating site was the median forebrain bundle while the recording site was the striatum with in vivo amperometric responses biosensors in the presence and absence of the metal oxides and the control electrode in the absence of enzyme at an applied potential of −150 mV vs Ag/AgCl during electrical stimulation of the median forebrain bundle.The black arrow shows the onset of the electrical stimulation and the red arrow indicates when stimulation was stopped.(D) In vivo amperometric responses obtained with the implanted biosensors and the control in the presence and absence of nomifensine at an applied potential of −150 mV vs Ag/AgCl during electrical stimulation of the median forebrain bundle.The stars indicate the onset of electrical stimulation while black arrow indicates when stimulation was stopped.Reproduced with permission from Ref. 30.

Figure 5 .
Figure 5. Design principles of several enzyme-mimetic microbiosensors for dopamine measurements.(A) MOF-decorated CF electrode with nicked-based electrocatalysis sites for increased electron transfer for dopamine oxidation, providing a linearity range of 0.004-0.4μΜ with a detection limit of 1 nM.The ME was used to measure dopamine in a mouse model of Parkinson disease (reproduced with permission from Ref. 113); (B) microelectrode array (250 μΜ) modified with Pt-reduced graphene oxide for measuring dopamine levels in the globus pallidus internal (GPi) in a rat model of Parkinson's disease following deep brain stimulation; the electrode was able to detect dopamine in the range of 50 nm-16.3μΜ with a detection limit of 50 nM at an oxidation potential of 0.13 V (reproduced with permission from Ref. 114).(C) Nickel tetrasulfonated phthalocyanine functionalized nitrogen-doped graphene for enhanced electrocatalysis of dopamine oxidation and measurements in the 1 × 10 −7 -2 × 10 −4 M range and a detection limit of 100 nM (with permission from Ref. 115).D) Synthesis of Mndoped crystalline CuO deposited on a CFME and DPV measurements of dopamine with an LOD of 9.1 × 10 −8 M and a sensitivity of 0.02 A M −1 at +0.075 V, (reproduced with permission from Ref. 116).

Figure 7 .
Figure 7. (A) Fabrication of carbon cloth functionalized with aminated EDA-Nb 2 CT x MXene as electrode materials for electrochemical dopamine detection in serum (with permission from Ref. 168) (B) Microneedle array patch detecting dopamine in sweat when placed on human skin.SEM images of (a) microneedle array; (b) microneedle array cross-section; (c) PDMS/CNT/CNC@PDMS patch surface (with permission from Ref. 169); (C) Fabrication and electrochemical testing of Cu-Gr/SPCE for dopamine detection (with permission from Ref. 136); (D) Laccase-based extended-gate OFET biosensor for dopamine detection in human urine (with permission from Ref. 106).

Figure 8 .
Figure 8. (A) (a) Electrochemical dopamine detection in neuro-2A cells using a N7-P-based microfluidic biosensor based on orthahedral building blocks and H1.5PO4 tetrahedra of N7, (b) integrated within a custom-made electrode layer assembled within a microchip with (c) bright field images showing in situ polymerization of PANI, and (d) in situ recording of the electropolymerization of aniline (with permission from Ref. 174); (B) Detection of dopamine using a sensors based on Fe single atom catalyst deposited within Ti 3 C 2 T x showing PC12 cell culture fluid collection; Fe-SACs/Ti 3 C 2 T x /LSGE dopamine detection mechanism; and amperometric response to dopamine (With permission from Ref. 132).(C) Electrochemical detection of dopamine release from SH-SY5Y cells using a laser-scribed graphene electrode modified with POM-ϒCD MOF (shown inset) (with permission from Ref. 134).

Figure 9 .
Figure 9. (A) Diagram depicting typical in vivo set-up for amperometric detection of neurotransmitters in Sprague Dawley rats.The working electrode is implanted into either of the three major dopaminergic epicenters; the striatum (SN), ventral tegmental area (VTA), or hypothalamus (HPA).(B) Schematic showing SWCNT/CFME biosensor implanted into a mouse brain along with electrode modification and DPV recording allowing dopamine and serotonin (5-HT) detection.The CFME was functionalized with diazonium salt-SWCNTs.The CFME's detection limit was 27.3 nM and the linear range spanned between 0.25-5⎕⎕⎕ measured at 0.18 V (with permission from Ref. 153).(C) In vivo voltammetry and rat surgery set-up showing recording and stimulating electrodes inserted unilaterally into the core of the nucleus accumbens and medial forebrain bundle; and design of waveform square wave voltammetry, voltammetric recordings, and calibration curve obtained with a CFME modified with PEDOT/Nafion (reproduced with permission from Ref. 128).

Figure 10 .
Figure 10.Examples of microelectrode arrays for in vivo dopamine detection: (A) implanted into primates showing electrode set up with 16 recording sites and the surface of the bioelectrode functionalized with Nafion and PtNPs (a)-(e) and example of impedance spectra at different stages of fabrication and amperometric response showing selectivity for dopamine and in vivo analysis of dopamine release in the cortex and striatum of normal (blue line) and parkinsonian (red line) monkeys; (with permission from Ref. 177); (B) Schematic of implantable PEDOT: PSS-coated diamond films deposited atop a microlight emitting diode (LED) for wireless detection of dopamine in freely-moving mice.Pictures of implanted biosensor with attached circuit board along with area of implantation (with permission from Ref. 119).(C) Wireless battery-free implantable microsensor for real-time monitoring of dopamine release within the brain of moving mice with examples of in vitro and in vivo recordings of optogenetic stimulation in mice.Data on right panel shows dopamine concentrations upon administration of morphine and naloxone during recording within nucleus accumbens of tethered mice (with permission from Ref. 178).

Figure 11 .
Figure 11.(A) (a) Schematic of dopamine detection within brain of rats with the novel concept of FSP-OECT biosensor (b) showing input FSP gating mode potential waveform and (c) transfer function with g m -v gs curves in response to dopamine (with permission from Ref. 38), (B) Example of electrode modification strategy a CFME with anti-biofouling layer for enhancing selectivity of dopamine detection (with permission from Ref. 69).

Figure 12 .
Figure 12.Examples of microfluidic-based electrochemical sensing of dopamine: (A) μSyringe sensing device with integrated graphite pencil electrode for electrochemical monitoring cerebrospinal fluid (CSF) dopamine in freely-moving mice showing experimental setting and sampling tube inserted into a mouse brain for CSF extraction, locomotor activity counts induced by amphetamine, and dopamine concentrations measured pre and post-test (with permission from Ref. 180); (B) (a) Illustration of microfluidic chamber with integrated electrochemical-sensing device and setup showing connected cables and inlet/outlet, (b) locomotor activity counts of saline-and reserpine-treated mice over 16 h, (c) picture of a PD mouse after the insertion of a guide cannula for CSF extraction and (d) dopamine concentrations measured in CSF of three study groups in the control, only reserpine-treated, and reserpine followed by -DOPA-treated mice (with permission from Ref. 171).

Table I .
Summary of enzyme-based biosensors for dopamine detection.

Table II .
Comparison of enzyme-mimetic biosensors for DA detection.Table II.(Continued).