Tyrosinase-Based Biosensors for Selective Dopamine Detection

A novel tyrosinase-based biosensor was developed for the detection of dopamine (DA). For increased selectivity, gold electrodes were previously modified with cobalt (II)-porphyrin (CoP) film with electrocatalytic activity, to act both as an electrochemical mediator and an enzyme support, upon which the enzyme tyrosinase (Tyr) was cross-linked. Differential pulse voltammetry was used for electrochemical detection and the reduction current of dopamine-quinone was measured as a function of dopamine concentration. Our experiments demonstrated that the presence of CoP improves the selectivity of the electrode towards dopamine in the presence of ascorbic acid (AA), with a linear trend of concentration dependence in the range of 2–30 µM. By optimizing the conditioning parameters, a separation of 130 mV between the peak potentials for ascorbic acid AA and DA was obtained, allowing the selective detection of DA. The biosensor had a sensitivity of 1.22 ± 0.02 µA·cm−2·µM−1 and a detection limit of 0.43 µM. Biosensor performances were tested in the presence of dopamine medication, with satisfactory results in terms of recovery (96%), and relative standard deviation values below 5%. These results confirmed the applicability of the biosensors in real samples such as human urine and blood serum.


Introduction
The detection of dopamine (DA) levels in physiological media, is gaining interest as dopamine is an important neurotransmitter which is linked to a large variety of medical conditions [1]. Selective identification of DA concentrations allows the monitoring of Parkinson's and Alzheimer's diseases, attention deficit hyperactivity disorder (ADHD), and schizophrenia [2][3][4][5], and the hormone is becoming accepted as a prognostic biomarker [6]. DA is one of the most important catecholamines, having an important role in the functionality of the central nervous system (CNS), but also influencing the hormonal, renal and cardiovascular systems [7], due to its mobility through the human body [8][9][10]. The treatment of some of the aforementioned diseases is performed with medication that blocks or activates corresponding receptors, and continuous monitoring in patients is required. Herewith, analysis and control of neurotransmitters is of great importance in the field of neuroscience, where current treatment is arbitrary, mainly based on clinical observations of the patient. For quantification, the most widely-used methods for neurotransmitter detection are in vitro methods like high performance liquid chromatography (HPLC) [11], absorption spectroscopy [12], where both methods need special laboratory setups and are time-consuming; or in vivo imaging techniques, which are very expensive [13]. Electrochemical sensors offer a real-time response with high sensitivity and selectivity, due to the electro-active nature of DA [14].
One of the problems for the electrochemical detection of dopamine is the co-existence of many interfering compounds in biological systems which can be oxidized at similar potentials, resulting

Instrumentation
All electrochemical measurements were carried out in a conventional electrochemical cell containing three electrodes. Gold bulk electrodes covered in Teflon (area 0.00785 cm 2 , eDAQ Pty. Ltd., Denistone East, New South Wales, Australia) were used as working electrodes, with a platinum foil as a counter electrode, and a Ag/AgCl electrode (3.5 M KCl) as reference. Chrono-amperometric, electrochemical impedance and voltammetric measurements (differential pulse and cyclic voltammograms) were performed by using a PalmSens3 electrochemical sensor interface (Palm Instruments BV, Houten, The Netherlands) controlled with PSTrace 4.8 software. Differential pulse voltammograms were recorded in different conditions in the presence and absence of the enzyme, where the constant amplitude of the pulse was set to 50 mV (pulse height) for 0.2 s (pulse width), at a scan rate of 10 mV·s −1 . For impedance measurements, an rms perturbation of 10 mV was applied over the frequency range 50 kHz-0.1 Hz, with 10 frequency values per frequency decade. The obtained spectra were recorded at a potential of −0.2 V vs. Ag/AgCl, and plotted in the form of complex plane diagrams (Nyquist plots) using electrochemical impedance spectroscopy (EIS) Spectrum Analyzer 1.0 software [34]. The surface morphology of the 2,3,7,8,12,13,17,18-Octaethyl-21H,23H-porphine cobalt(II) thin film was investigated by atomic force microscopy (AFM NT-MDT model NTEGRA PRIMA EC). The images were taken in semi-contact mode with "GOLDEN" silicon cantilever (NCSG10, force constant 0.15 N/m, tip radius 10 nm).

CoP-Film Preparation and Construction of Tyrosinase Biosensor
A clean gold electrode was modified with CoP film by immersion in 0.5% solution of CoP dissolved in chloroform, for 10 min. For the next 10 min, the adsorbed film was dried in air and deposited in buffer solution in order to avoid the interaction of the porphyrin film with gas molecules [35]. For the Tyr-based biosensor, after drying the CoP film, an aqueous solution containing 1% Tyr and 4% BSA, was mixed with cross-linking agent 2.5% GA and pipetted onto the electrode surface; thus a concentration of 0.67% Tyr was immobilized. The enzyme layer was left to dry for 10 min, washed, and stored in buffer solution in between measurements.

Characterization of the CoP-Tyr Biosensor
The monitoring of neurotransmitter levels is vague, highlighting the need for rapid and selective tools. Electrochemical sensors have some limitations, especially due to the lack of resolution between DA and other electroactive species coexisting in the cerebral system (the concentrations of some interferent species, such as AA, are much higher than that of DA in the CNS). Biosensor performances offer here an advantage towards selectivity, sensitivity, and surface morphology. This, together with electrochemical characterization, was performed as follows.

Surface Morphology Analysis of CoP Film
Atomic Force Microscopy (AFM) was employed to analyze the surface morphology of the CoP-film. Surface topography is usually described by amplitude parameters such as average roughness and root mean square roughness (RMS) [36]. It was in our interest for the thin film to offer a good support for enzyme immobilization. In this context, it was important for the film to have a large number of active adsorption sites. This can be appreciated from the film roughness. The degree of particle deposition upon a surface, is given by an increasing value of the RMS, as shown in Figure 1. For a RMS of 12.16 nm, the peaks and valleys are more pronounced, offering accessible binding sites on the films for enzyme layer. However, the CoP film was evenly distributed, as seen from the histograms. Atomic Force Microscopy (AFM) was employed to analyze the surface morphology of the CoPfilm. Surface topography is usually described by amplitude parameters such as average roughness and root mean square roughness (RMS) [36]. It was in our interest for the thin film to offer a good support for enzyme immobilization. In this context, it was important for the film to have a large number of active adsorption sites. This can be appreciated from the film roughness. The degree of particle deposition upon a surface, is given by an increasing value of the RMS, as shown in Figure 1. For a RMS of 12.16 nm, the peaks and valleys are more pronounced, offering accessible binding sites on the films for enzyme layer. However, the CoP film was evenly distributed, as seen from the histograms.

Electrochemical Characterization of CoP Film
Cyclic voltammetry (CV) offers an insight into the overall characteristics of a sensor material. Figure 2 illustrates the voltammograms of the modified Au electrode in NaPB, pH 8.0, at a scan rate of 50 mV·s −1 . It can be said that responses of the modified electrodes CoP and CoP-Tyr are bigger compared to bare gold electrode, which may be attributed to the increase in surface area of the modified electrodes. Voltammograms obtained with the CoP-Tyr-biosensor in presence of two different DA concentrations showed a considerable increase in current intensity for the peaks of the DA redox reaction. This suggested an increase in the electron transfer rate. Figure 2 reveals a very well defined anodic peak assigned to the oxidation of DA, which was paired with the corresponding smaller reduction peak. This suggested that a quasi-reversible redox reaction of DA occurred at the CoP-sensor, which was associated with its oxidation product associating/dissociating from the CoP film.

Electrochemical Characterization of CoP Film
Cyclic voltammetry (CV) offers an insight into the overall characteristics of a sensor material. Figure 2 illustrates the voltammograms of the modified Au electrode in NaPB, pH 8.0, at a scan rate of 50 mV·s −1 . It can be said that responses of the modified electrodes CoP and CoP-Tyr are bigger compared to bare gold electrode, which may be attributed to the increase in surface area of the modified electrodes. Voltammograms obtained with the CoP-Tyr-biosensor in presence of two different DA concentrations showed a considerable increase in current intensity for the peaks of the DA redox reaction. This suggested an increase in the electron transfer rate. Figure 2 reveals a very well defined anodic peak assigned to the oxidation of DA, which was paired with the corresponding smaller reduction peak. This suggested that a quasi-reversible redox reaction of DA occurred at the CoP-sensor, which was associated with its oxidation product associating/dissociating from the CoP film. In order to characterize the mechanism of the electrode reaction at the CoP film, the relationship between scan rate and peak current during CVs was studied, as shown in Figure 3A. There was a considerable increase in the current intensity for the oxidation peak of DA, with an increasing scan rate. This suggested an increase in the electron transfer rate. However, the oxidation peak potential (0.15 V) was almost unchanged, exhibiting a linear relationship with the square root of the scan rates over the range 10-110 mV s −1 . The same linear range was observed for the reduction peak at 0.08 V. The linear relationship of the square root of the scan rate with the values of the anodic peak current ( ) and cathodic peak current ( ) can be seen in Figure 3B. Taking into account the linear regression equations below, the influence of scan rate explained the electrode process in terms of a diffusion controlled reaction (mass transport).   In order to characterize the mechanism of the electrode reaction at the CoP film, the relationship between scan rate and peak current during CVs was studied, as shown in Figure 3A. There was a considerable increase in the current intensity for the oxidation peak of DA, with an increasing scan rate. This suggested an increase in the electron transfer rate. However, the oxidation peak potential (0.15 V) was almost unchanged, exhibiting a linear relationship with the square root of the scan rates over the range 10-110 mV s −1 . The same linear range was observed for the reduction peak at 0.08 V. The linear relationship of the square root of the scan rate with the values of the anodic peak current (I pa ) and cathodic peak current (I pc ) can be seen in Figure 3B. Taking into account the linear regression equations below, the influence of scan rate explained the electrode process in terms of a diffusion controlled reaction (mass transport). In order to characterize the mechanism of the electrode reaction at the CoP film, the relationship between scan rate and peak current during CVs was studied, as shown in Figure 3A. There was a considerable increase in the current intensity for the oxidation peak of DA, with an increasing scan rate. This suggested an increase in the electron transfer rate. However, the oxidation peak potential (0.15 V) was almost unchanged, exhibiting a linear relationship with the square root of the scan rates over the range 10-110 mV s −1 . The same linear range was observed for the reduction peak at 0.08 V. The linear relationship of the square root of the scan rate with the values of the anodic peak current ( ) and cathodic peak current ( ) can be seen in Figure 3B. Taking into account the linear regression equations below, the influence of scan rate explained the electrode process in terms of a diffusion controlled reaction (mass transport).

Electrochemical Impedance Spectroscopy
EIS was used to characterize the bulk and electrode/electrolyte interface phenomena, providing information about electron transfer and charge polarization. It was employed to identify interfacial changes after CoP film deposition and Tyr entrapment, similar to a layer-by-layer (LbL) structure formation on the electrode surface. Figure 4 shows the complex plane representation of the fitted impedance spectra acquired for Au, Au/CoP and Au/CoP-Tyr at a potential applied during measurements of −0.2 V vs. Ag/AgCl. The potential value was chosen to highlight the modifications of the electrode-electrolyte interface phenomena with each layer deposition, taking into consideration that the electrode material was Au.

Electrochemical Impedance Spectroscopy
EIS was used to characterize the bulk and electrode/electrolyte interface phenomena, providing information about electron transfer and charge polarization. It was employed to identify interfacial changes after CoP film deposition and Tyr entrapment, similar to a layer-by-layer (LbL) structure formation on the electrode surface. Figure 4 shows the complex plane representation of the fitted impedance spectra acquired for Au, Au/CoP and Au/CoP-Tyr at a potential applied during measurements of −0.2 V vs. Ag/AgCl. The potential value was chosen to highlight the modifications of the electrode-electrolyte interface phenomena with each layer deposition, taking into consideration that the electrode material was Au. The spectra were fitted using an equivalent electrical circuit shown in the inset of Figure 4, and consisted of a cell resistance, representing the electrical resistance of the cell and electrolyte solution (RΩ), in series with two parallel combinations. The first parallel combination was associated, in all situations, with a CoP or CoP-Tyr film modified with an Au electrode/electrolyte solution interface and consisted of a charge transfer resistance (Rct) and a double layer non-ideal capacitance (CPEdl). The Au electrode/film interface (CoP or CoP-Tyr films) introduced in series a second parallel combination of the film charge-transfer resistance (Rf) and a non-ideal capacitance (CPEf). Both non-ideal capacitances were represented by constant phase elements (CPE), according to the equation: where C is the ideal capacitance, ω the radial frequency and the exponent α, which reflects the surface uniformity. Table 1 shows the values of the circuit components obtained by fitting the experimental spectra to the electrical equivalent circuit. The value of the electrical resistance of the electrolyte solution and electrical contacts remained almost constant throughout each deposition step, RΩ ≅ 18 Ω·cm 2 . The CoP film and the enzyme layer contribution were shown by the high CPEf values, ranging between 956.3 and 901.6 µF·cm −2 ·s α−1 . This indicated a charge accumulation at the electrode/film interface. The film resistance value increased from 0.78 kΩ·cm 2 for CoP (conductive), up to 1.42 kΩ·cm 2 after enzyme immobilization (less conductive layer), which was in concordance with CPEf values. The high average value of αf ≅ 0.97 reflected the uniformity and smoothness of both enzyme and CoP The spectra were fitted using an equivalent electrical circuit shown in the inset of Figure 4, and consisted of a cell resistance, representing the electrical resistance of the cell and electrolyte solution (R Ω ), in series with two parallel combinations. The first parallel combination was associated, in all situations, with a CoP or CoP-Tyr film modified with an Au electrode/electrolyte solution interface and consisted of a charge transfer resistance (R ct ) and a double layer non-ideal capacitance (CPE dl ). The Au electrode/film interface (CoP or CoP-Tyr films) introduced in series a second parallel combination of the film charge-transfer resistance (R f ) and a non-ideal capacitance (CPE f ). Both non-ideal capacitances were represented by constant phase elements (CPE), according to the equation: where C is the ideal capacitance, ω the radial frequency and the exponent α, which reflects the surface uniformity. Table 1 shows the values of the circuit components obtained by fitting the experimental spectra to the electrical equivalent circuit. The value of the electrical resistance of the electrolyte solution and electrical contacts remained almost constant throughout each deposition step, R Ω ∼ = 18 Ω·cm 2 . The CoP film and the enzyme layer contribution were shown by the high CPE f values, ranging between 956.3 and 901.6 µF·cm −2 ·s α−1 . This indicated a charge accumulation at the electrode/film interface. The film resistance value increased from 0.78 kΩ·cm 2 for CoP (conductive), up to 1.42 kΩ·cm 2 after enzyme immobilization (less conductive layer), which was in concordance with CPE f values. The high average value of α f ∼ = 0.97 reflected the uniformity and smoothness of both enzyme and CoP films, in accordance with the AFM images. At the electrolyte interface, R ct values were higher for the bare electrode (4.64 kΩ·cm 2 ) and started decreasing with each deposited layer: from 4.64 kΩ cm 2 for bare Au, to 2.09 kΩ·cm 2 for Au-CoP and 2 kΩ·cm 2 after enzyme entrapment, which was attributed to a higher electron transfer through interface. The value of CPE dl increased from 29.4 µF·cm −2 ·s α−1 for bare Au, to 30 µF·cm −2 ·s α−1 for CoP film, and doubled for the enzyme layer. This suggested that the adsorption of both materials led to changes in space charge polarization. The values of α dl ranging from 0.80 to 0.87 suggested that the interface changed after each adsorption step. Table 1. Equivalent circuit element values obtained by fitting the impedance spectra from Figure 5 to the corresponding circuit. EIS was also used to characterize the electrode/electrolyte interface phenomena, in the presence of dopamine. Figure 5 shows the complex plane representation, with a 0.1 V potential applied during measurements, of the impedance spectra acquired for Au/CoP-Tyr in 0.1 M NAPB, pH 8.0 containing 30 and 60 µM DA. The semicircle diameter of this Nyquist plot reflected the electron transfer resistance (R ct ), which refers to current flow produced by the reactions at the interface, and was found to be lower when the DA concentration increased, suggesting the bio-catalytic activity of tyrosinase at the surface of the biosensor towards the oxidation of DA. The spectra were fitted using an equivalent electrical circuit shown in the inset of Figure 5. This consisted of a cell resistance, R Ω , in series with a parallel combination of a charge transfer resistance, R ct , through the Au-CoP/Tyr film interface, and a double layer capacitance (CPE f ), represented as a constant phase element, which resulted from charge being stored in the double layer at the interface, in high and intermediate frequency regions. A further CPE element (CPE dl ), in series with the parallel combination, was used to monitor the capacitive behavior of the upper immobilized enzyme layer (Tyr cross-linked with glutaraldehyde), in contact with the electrolyte in the low frequency region, which significantly varied in the presence of DA. films, in accordance with the AFM images. At the electrolyte interface, Rct values were higher for the bare electrode (4.64 kΩ·cm 2 ) and started decreasing with each deposited layer: from 4.64 kΩ cm 2 for bare Au, to 2.09 kΩ·cm 2 for Au-CoP and 2 kΩ·cm 2 after enzyme entrapment, which was attributed to a higher electron transfer through interface. The value of CPEdl increased from 29.4 µF·cm −2 ·s α−1 for bare Au, to 30 µF·cm −2 ·s α−1 for CoP film, and doubled for the enzyme layer. This suggested that the adsorption of both materials led to changes in space charge polarization. The values of αdl ranging from 0.80 to 0.87 suggested that the interface changed after each adsorption step. Table 1. Equivalent circuit element values obtained by fitting the impedance spectra from Figure 5 to the corresponding circuit. EIS was also used to characterize the electrode/electrolyte interface phenomena, in the presence of dopamine. Figure 5 shows the complex plane representation, with a 0.1 V potential applied during measurements, of the impedance spectra acquired for Au/CoP-Tyr in 0.1 M NAPB, pH 8.0 containing 30 and 60 µM DA. The semicircle diameter of this Nyquist plot reflected the electron transfer resistance (Rct), which refers to current flow produced by the reactions at the interface, and was found to be lower when the DA concentration increased, suggesting the bio-catalytic activity of tyrosinase at the surface of the biosensor towards the oxidation of DA. The spectra were fitted using an equivalent electrical circuit shown in the inset of Figure 5. This consisted of a cell resistance, RΩ, in series with a parallel combination of a charge transfer resistance, Rct, through the Au-CoP/Tyr film interface, and a double layer capacitance (CPEf), represented as a constant phase element, which resulted from charge being stored in the double layer at the interface, in high and intermediate frequency regions. A further CPE element (CPEdl), in series with the parallel combination, was used to monitor the capacitive behavior of the upper immobilized enzyme layer (Tyr cross-linked with glutaraldehyde), in contact with the electrolyte in the low frequency region, which significantly varied in the presence of DA.  Table 2 shows the values of the circuit components obtained by fitting the experimental spectra to the equivalent electrical circuit for DA oxidation. The cell resistance kept an almost constant value around 5 Ω·cm 2 . As expected, the value of the charge transfer resistance Rct keeps decreasing for each  Table 2 shows the values of the circuit components obtained by fitting the experimental spectra to the equivalent electrical circuit for DA oxidation. The cell resistance kept an almost constant value around 5 Ω·cm 2 . As expected, the value of the charge transfer resistance R ct keeps decreasing for each DA addition, from 58.57 kΩ·cm 2 in the absence of DA, to 6.99 kΩ·cm 2 in the presence of DA, indicating the conducting properties of the Au and CoP film. The capacitance of the modified electrodes with electrocatalytically CoP film depends mainly on the surface area accessible to the electrolyte ions and redox species, which depends in turn on the specific surface area, pore-size distribution and shape. The increase in both the double layer capacitance CPE f (61.07 to 100.01 µF·cm −2 ·s α−1 ) and the α dl values ranging from 0.71 to 0.82 indicated a charge accumulation at the Au-CoP/Tyr layer interface, influenced by the oxidation of DA. At the upper Tyr layer /electrolyte interface, capacitance values CPE dl also increased with DA concentration, doubling in value from 300.63 (in the absence of DA) to 620.50, and 665.47 µF·cm −2 ·s α−1 respectively, for 30 and 60 µM DA. With DA and dopa-quinone molecules accumulating at the electrode/film interface, α f values decreased.

The Role of CoP in Dopamine Oxidation
The rate of the electrochemical reactions was significantly influenced by the nature of the electrode surface. Porphyrins are less widely used as a surface modifier or electrochemical mediator in (bio)sensors and their interaction with different analytes are less well-studied. Metalloporphyrins (porphyrin systems with metallic ions) have low energy excitations in the visible spectral region, and they also accept or donate electrons easily [37]. The two-dimensional geometry of porphyrins and their electronic structure both promote very rapid and vectorial electron transfer, and thorough interaction of these macrorings with analytes [38]. It has been highlighted that two fundamental cooperative effects can take place in the sensing phenomenon, and are the main determinants of the performances of chemical sensors based on porphyrins: weak interactions (such as Van der Waals or London forces and hydrogen bonding) and the coordination of analytes [39].
The central metal of the metalloporphyrin affects sensing, as dopamine oxidation to dopaquinone can be performed using both transition metals, as well as catalysis by enzymes (e.g., tyrosinase). For CoP-sensors, the aromatic-stacking and electrostatic attraction between positively-charged dopamine (protonated amine group at physiological pH) and negatively-charged porphyrin can accelerate the electron transfer, while weakening AA oxidation (the main interferent on physiological samples) on the porphyrin-functionalized gold-modified electrode. For CoP-Tyr-biosensor, the CoP film acts as a mediator to enhance the direct electron transfer between the enzyme and Au electrodes, which is usually prohibited due to the shielding of redox active sites by the protein shells. Mediators are widely used to access the redox center of the enzyme, and thus act as electron shuttles.
The following characteristics recommend the use of metalloporphyrins: 1.
Electrocatalytic activity toward dopamine oxidation (enhancing the electronic conductivity and promoting electron transfer rate between the DA and electrode surface) in electrochemical sensor development, and 2.
Electrochemical mediator activity during enzyme-catalyzed oxidation of dopamine (enhancing electronic conductivity and acting as charge carriers) and 3. Support for enzyme immobilization for biosensor development.
With CoP-sensors, dopamine can be easily electrocatalytically oxidized at the CoP film to form dopamine quinone (DAQ) which can be reduced at the electrode surface when a potential is applied to the electrode, after the exchange of two electrons (and two protons) to produce a Faradaic current [40]. In the case of biosensors, during the DA oxidation steps, the oxidation states of the copper atoms of tyrosinase change to give different forms of the enzyme [23]. Native tyrosinase occurs mainly as met-tyrosinase (Met-Tyr) in which a hydroxyl ion is bound to the two copper ions and both copper ions are in the Cu(II) oxidation state; this form, in the presence of oxygen, catalyzes the oxidation of catechols like DA to DAQ with H 2 O production. During this process, Met-Tyr is reduced to deoxy-tyrosinase (Doxy-Tyr) in which both copper ions are in the Cu(I) oxidation state. Doxy-Tyr binds oxygen to generate Oxy-Tyr, which is reduced to Met-Tyr while it catalyzes the oxidation of DA to DAQ (Scheme 1A). DAQ is further reduced at the electrode surface (Scheme 1B). atoms of tyrosinase change to give different forms of the enzyme [23]. Native tyrosinase occurs mainly as met-tyrosinase (Met-Tyr) in which a hydroxyl ion is bound to the two copper ions and both copper ions are in the Cu(II) oxidation state; this form, in the presence of oxygen, catalyzes the oxidation of catechols like DA to DAQ with H2O production. During this process, Met-Tyr is reduced to deoxy-tyrosinase (Doxy-Tyr) in which both copper ions are in the Cu(I) oxidation state. Doxy-Tyr binds oxygen to generate Oxy-Tyr, which is reduced to Met-Tyr while it catalyzes the oxidation of DA to DAQ (Scheme 1A). DAQ is further reduced at the electrode surface (Scheme 1B).

Scheme 1. (A) Tyrosinase forms during enzyme catalysis of dopamine at CoP-Tyr-biosensor and (B)
detection principle at the CoP-Tyr biosensor surface.

Influence of pH
In order to enhance the selective DA detection in the presence of AA at the CoP-film surface, the influence of the pH of the supporting electrolyte on the electrochemical activity of the CoP-sensors was studied. The current response obtained from the enzyme-catalyzed reaction presented a maximum value and maximum peaks separation at pH 8.0, which was chosen for all measurements.

Influence of CoP Concentration
The effect of cobalt (II)-porphyrin compound concentrations in the solution used for CoP film deposition on an Au surface was studied, and a concentration of 0.5% in chloroform provided better results, thus this concentration value was selected.

Parameters of the Detection Method
Since differential pulse voltammetry (DPV) is more sensitive for cyclic voltammetry, it was used to detect DA (in the absence and presence of AA). In order to increase the sensor response, pretreatment steps were employed, immediately followed by the measurement step. All steps were performed in the electrolyte solution containing the analytes (DA and AA). The effect of modifying the electrode surface with CoP film was noticed in DPV since the separated DA and AA peaks appeared at lower potentials: at 0.08 V for DA and at −0.03 V for AA (compared with 0.15 V for both AA and DA with bare electrodes, results not shown). Optimized parameters were determined for both selective and sensitive DA detection ( Figure 6A), as well as to avoid electrode surface passivation by the phenoxy radicals (dopachrome). Thus, the optimized pretreatment settings of the CoP-sensors towards DA detection were: a conditioning potential of −1.2 V for a period of 20 s, followed by 40 s of a preconcentration step at −0.5 V. The pretreatment influence on the CoP-Tyr-biosensor response was studied, to better discriminate between the AA and DA peaks, since the peaks of AA and DA were slightly overlapping without any pretreatments. In the presence of the enzyme, the conditioning potential was no longer necessary, since it did not influence the peak potentials (data not shown). The variation of the preconcentration potential was studied, and a value of −1.3 V for a fixed time of 40 s was chosen, since it clearly assisted in discriminating the AA and DA

Influence of pH
In order to enhance the selective DA detection in the presence of AA at the CoP-film surface, the influence of the pH of the supporting electrolyte on the electrochemical activity of the CoP-sensors was studied. The current response obtained from the enzyme-catalyzed reaction presented a maximum value and maximum peaks separation at pH 8.0, which was chosen for all measurements.

Influence of CoP Concentration
The effect of cobalt (II)-porphyrin compound concentrations in the solution used for CoP film deposition on an Au surface was studied, and a concentration of 0.5% in chloroform provided better results, thus this concentration value was selected.

Parameters of the Detection Method
Since differential pulse voltammetry (DPV) is more sensitive for cyclic voltammetry, it was used to detect DA (in the absence and presence of AA). In order to increase the sensor response, pretreatment steps were employed, immediately followed by the measurement step. All steps were performed in the electrolyte solution containing the analytes (DA and AA). The effect of modifying the electrode surface with CoP film was noticed in DPV since the separated DA and AA peaks appeared at lower potentials: at 0.08 V for DA and at −0.03 V for AA (compared with 0.15 V for both AA and DA with bare electrodes, results not shown). Optimized parameters were determined for both selective and sensitive DA detection ( Figure 6A), as well as to avoid electrode surface passivation by the phenoxy radicals (dopachrome). Thus, the optimized pretreatment settings of the CoP-sensors towards DA detection were: a conditioning potential of −1.2 V for a period of 20 s, followed by 40 s of a preconcentration step at −0.5 V. The pretreatment influence on the CoP-Tyr-biosensor response was studied, to better discriminate between the AA and DA peaks, since the peaks of AA and DA were slightly overlapping without any pretreatments. In the presence of the enzyme, the conditioning potential was no longer necessary, since it did not influence the peak potentials (data not shown). The variation of the preconcentration potential was studied, and a value of −1.3 V for a fixed time of 40 s was chosen, since it clearly assisted in discriminating the AA and DA peaks ( Figure 6B). Since AA (which is not a substrate of tyrosinase) can diffuse to the electrode surface, we can concluded that the use of negative potentials in the pretreatment steps, was necessary to better discriminate between the AA and DA peaks for both the CoP-sensor and the CoP-Tyr-biosensor. Positive potentials were also tested, but no discrimination between the two peaks was observed. surface, we can concluded that the use of negative potentials in the pretreatment steps, was necessary to better discriminate between the AA and DA peaks for both the CoP-sensor and the CoP-Tyr-biosensor. Positive potentials were also tested, but no discrimination between the two peaks was observed.

Analytical Parameters of the CoP-Tyr-Biosensor vs. the CoP-Sensor
DPV of DA at the CoP-sensor showed that the DA peak increased with increasing concentration. The determination of DA performed under optimized conditions exhibited a linear response from 10 to 50 µM, a sensitivity of 0.577 ± 0,041 µA·cm -2 ·µM -1 , and a detection limit of 0.98 µM [limit of detection (LOD) = 3 × SD/S, where SD was the standard deviation of the blank and Swas the sensor sensibility). All of the calibration curves were obtained in triplicate. Simultaneous increasing concentrations for AA and DA (5-50 µM DA and 25-250 µM AA) were also monitored, and the data is presented in the Figure 6A. In this case, the CoP-sensor performances towards DA were slightly decreased to a sensitivity of 0.443 ± 0.01 µA·cm −2 ·µM −1 , and an increased detection limit of 3.13 µM (R 2 = 0.996). The CoP modified sensor reached a sensitivity of 0.561 ± 0.01 µA·cm −2 ·µM −1 and a detection limit of 2.46 µM (R 2 = 0.997) for DA, when 200 µM AA concentration was kept constant ( Figure 7B). Since in physiological conditions (the central nervous system), the concentrations of AA can be 100-1000 times higher than those of DA [41], all experimental data was carried out with a higher ratio of AA to DA.
The sensor performances were also tested in the presence of dopamine hydrochloride (5 mg/mL) medication, which is used in intravenous infusions. The data is presented in comparison to the CoP-Tyr-biosensor in Chapter 3.4.
DPV was also used as a sensitive and selective detection method for DA in the presence of AA using the CoP-Tyr-biosensor. The peak potential for AA was in the negative region at −0.07 V while that of DA was at 0.06 V, maintaining a difference of 130 mV between the two peaks. Table 3 summarizes the analytical performances of both CoP and CoP-Tyr (bio)sensors.

Analytical Parameters of the CoP-Tyr-Biosensor vs. the CoP-Sensor
DPV of DA at the CoP-sensor showed that the DA peak increased with increasing concentration. The determination of DA performed under optimized conditions exhibited a linear response from 10 to 50 µM, a sensitivity of 0.577 ± 0,041 µA·cm -2 ·µM -1 , and a detection limit of 0.98 µM [limit of detection (LOD) = 3 × SD/S, where SD was the standard deviation of the blank and Swas the sensor sensibility). All of the calibration curves were obtained in triplicate. Simultaneous increasing concentrations for AA and DA (5-50 µM DA and 25-250 µM AA) were also monitored, and the data is presented in the Figure 6A. In this case, the CoP-sensor performances towards DA were slightly decreased to a sensitivity of 0.443 ± 0.01 µA·cm −2 ·µM −1 , and an increased detection limit of 3.13 µM (R 2 = 0.996). The CoP modified sensor reached a sensitivity of 0.561 ± 0.01 µA·cm −2 ·µM −1 and a detection limit of 2.46 µM (R 2 = 0.997) for DA, when 200 µM AA concentration was kept constant ( Figure 7B). Since in physiological conditions (the central nervous system), the concentrations of AA can be 100-1000 times higher than those of DA [41], all experimental data was carried out with a higher ratio of AA to DA.
The sensor performances were also tested in the presence of dopamine hydrochloride (5 mg/mL) medication, which is used in intravenous infusions. The data is presented in comparison to the CoP-Tyr-biosensor in Chapter 3.4.
DPV was also used as a sensitive and selective detection method for DA in the presence of AA using the CoP-Tyr-biosensor. The peak potential for AA was in the negative region at −0.07 V while that of DA was at 0.06 V, maintaining a difference of 130 mV between the two peaks. Table 3 summarizes the analytical performances of both CoP and CoP-Tyr (bio)sensors.    The biosensor performance was tested under the established conditions. Figure 8 shows a linear response of the CoP-Tyr-biosensor for DA for the range 2 to 30 µM, with a sensitivity of 1.22 ± 0.02 µA cm −2 µM −1 and a detection limit of 0.43 µM. The average sensitivity of the biosensor was considerably higher in comparison to the CoP-sensor (R 2 = 0.997). All the calibration curves were obtained in triplicate. Thus, for the CoP-Tyr-biosensor, the peaks of DA were much higher than that at the CoP-sensor. Figure 6B presents the CoP-Tyr-biosensor response for simultaneous increasing concentrations of AA and DA. The concentration of DA was varied as previously, while the concentration of AA was varied from 25 to 250 µM (50 µM per injection). It can be seen that the peak currents for the two analytes increased linearly with their concentrations. The biosensor    The biosensor performance was tested under the established conditions. Figure 8 shows a linear response of the CoP-Tyr-biosensor for DA for the range 2 to 30 µM, with a sensitivity of 1.22 ± 0.02 µA cm −2 µM −1 and a detection limit of 0.43 µM. The average sensitivity of the biosensor was considerably higher in comparison to the CoP-sensor (R 2 = 0.997). All the calibration curves were obtained in triplicate. Thus, for the CoP-Tyr-biosensor, the peaks of DA were much higher than that at the CoP-sensor. Figure 6B presents the CoP-Tyr-biosensor response for simultaneous increasing concentrations of AA and DA. The concentration of DA was varied as previously, while the concentration of AA was varied from 25 to 250 µM (50 µM per injection). It can be seen that the peak currents for the two analytes increased linearly with their concentrations. The biosensor The biosensor performance was tested under the established conditions. Figure 8 shows a linear response of the CoP-Tyr-biosensor for DA for the range 2 to 30 µM, with a sensitivity of 1.22 ± 0.02 µA cm −2 µM −1 and a detection limit of 0.43 µM. The average sensitivity of the biosensor was considerably higher in comparison to the CoP-sensor (R 2 = 0.997). All the calibration curves were obtained in triplicate. Thus, for the CoP-Tyr-biosensor, the peaks of DA were much higher than that at the CoP-sensor. Figure 6B presents the CoP-Tyr-biosensor response for simultaneous increasing concentrations of AA and DA. The concentration of DA was varied as previously, while the concentration of AA was varied from 25 to 250 µM (50 µM per injection). It can be seen that the peak currents for the two analytes increased linearly with their concentrations. The biosensor performances for DA remain at a sensitivity of 1.21 ± 0.03 µA cm −2 µM −1 with a detection limit of 0.55 µM. The change of AA concentration did not have a significant influence on the peak current and the peak potential of the DA. Figure 7A shows the response for DA in the presence of a fixed concentration of 200 µM AA, with a slightly lower sensitivity of 1.21 ± 0.02 µA cm −2 µM −1 with a detection limit of 0.52 µM. The changes of AA peaks were at the relative standard deviation (RSD) level of individual voltammetric measurements (here about 2.1%). The peaks were well separated on both situations when the CoP-Tyr-biosensor was used; around 130 mV, which is enough for avoiding undesired interferences. By comparison with the CoP-sensor ( Figure 7B), the separation between peaks is bigger and DA peaks are much higher. These results were reflected in the higher sensitivity and lower LOD for CoP-Tyr-biosensor towards DA in the presence of AA.
The values of the apparent Michaelis-Menten constant K M and the maximum current density J max , (corresponding to tyrosinase saturated with dopamine) were estimated by a Lineweaver-Burk type plot. Values of 155.52 µA cm −2 for the maximum current and 92.68 µM for K M were obtained, which are in accordance with values obtained in the literature for immobilized tyrosinase (60-200 µM). Generally, the Michaelis-Menten approach applied to the immobilized enzymes typically yields increased values of K M . The effect of surface confinement on the kinetics of bound enzymes is mainly related to the diffusion limitation of the substrate to the reaction centers, by virtue of the immobilization itself (spatial limitation). Smaller K M values suggest a higher affinity or binding strength between the immobilized enzyme and its substrate dopamine, that can overcome the negative effect of immobilization and possible conformational changes of the enzyme.

Biosensor Stability
The long term stability of our biosensors was evaluated by repeated measurements on different days over a timeframe of two weeks, determining its sensitivity for DA detection, as shown in Figure 9. The biosensor kept its activity up to 95% in the first week, dropping to less than 50% on the 9th day, and decreasing subsequently. In comparison to the CoP sensor, whose sensitivity dropped after two days (data not shown), the CoP-Tyr biosensor had better stability, with the sensitivity keeping initial values for almost a week. In between measurements, the biosensor was kept in buffer solution in the refrigerator for~4 • C. performances for DA remain at a sensitivity of 1.21 ± 0.03 µA cm −2 µM −1 with a detection limit of 0.55 µM. The change of AA concentration did not have a significant influence on the peak current and the peak potential of the DA. Figure 7A shows the response for DA in the presence of a fixed concentration of 200 µM AA, with a slightly lower sensitivity of 1.21 ± 0.02 µA cm −2 µM −1 with a detection limit of 0.52 µM. The changes of AA peaks were at the relative standard deviation (RSD) level of individual voltammetric measurements (here about 2.1%). The peaks were well separated on both situations when the CoP-Tyr-biosensor was used; around 130 mV, which is enough for avoiding undesired interferences. By comparison with the CoP-sensor ( Figure 7B), the separation between peaks is bigger and DA peaks are much higher. These results were reflected in the higher sensitivity and lower LOD for CoP-Tyr-biosensor towards DA in the presence of AA. The values of the apparent Michaelis-Menten constant KM and the maximum current density Jmax, (corresponding to tyrosinase saturated with dopamine) were estimated by a Lineweaver-Burk type plot. Values of 155.52 µA cm −2 for the maximum current and 92.68 µM for KM were obtained, which are in accordance with values obtained in the literature for immobilized tyrosinase (60-200 µM). Generally, the Michaelis-Menten approach applied to the immobilized enzymes typically yields increased values of KM. The effect of surface confinement on the kinetics of bound enzymes is mainly related to the diffusion limitation of the substrate to the reaction centers, by virtue of the immobilization itself (spatial limitation). Smaller KM values suggest a higher affinity or binding strength between the immobilized enzyme and its substrate dopamine, that can overcome the negative effect of immobilization and possible conformational changes of the enzyme.

Biosensor Stability
The long term stability of our biosensors was evaluated by repeated measurements on different days over a timeframe of two weeks, determining its sensitivity for DA detection, as shown in figure  9. The biosensor kept its activity up to 95% in the first week, dropping to less than 50% on the 9th day, and decreasing subsequently. In comparison to the CoP sensor, whose sensitivity dropped after two days (data not shown), the CoP-Tyr biosensor had better stability, with the sensitivity keeping initial values for almost a week. In between measurements, the biosensor was kept in buffer solution in the refrigerator for ~4 °C.

Real Sample Detection
Sensor and biosensor performances were tested in the presence of dopamine hydrochloride medication (5 mg/mL) with an equivalent of 26.36 mM DA, which is usually used in blood pressure treatments. Aqueous dilutions up to 1000-10,000 times (depending on the working electrode type CoP, respectively CoP-Tyr) were prepared. The concentration of dopamine from the vial (Cx) was determined using standard addition method (SAM) and its recovery was calculated. After adding the sample, three known concentrations of DA were added, in order to extrapolate the value for Cx.

Real Sample Detection
Sensor and biosensor performances were tested in the presence of dopamine hydrochloride medication (5 mg/mL) with an equivalent of 26.36 mM DA, which is usually used in blood pressure treatments. Aqueous dilutions up to 1000-10,000 times (depending on the working electrode type CoP, respectively CoP-Tyr) were prepared. The concentration of dopamine from the vial (C x ) was determined using standard addition method (SAM) and its recovery was calculated. After adding the sample, three known concentrations of DA were added, in order to extrapolate the value for C x . All of the measurements were done in triplicate. The results are presented in Table 4. Since the relative standard deviation (RSD) values obtained for the biosensor had a value up to 4.7%, this indicated a potential use for the biosensor in real samples, such as blood or urine. The average recoveries were 96% for the CoP-Tyr biosensor and 99.6% for the CoP sensor. The lower value for the CoP-Tyr biosensor recovery can be explained by conformational changes induced by the cross-linking immobilization method, affecting the active site of the enzyme. Also, dopamine and dopamine-quinone diffusion towards the electrode surface can be hindered. Milder enzyme immobilization methods should be considered for future approaches. Calibration curves of both the CoP-sensor and the CoP-Tyr-biosensor have been also used to determine the sample concentration. The recovery for each found value is presented in Table 5. Table 5. Determination of DA in dopamine hydrochloride vial (n = 3) using the calibration curve from Figure 8. The performances of our sensors were also compared with other sensors reported in the literature for dopamine detection, and the results are shown in Table 6. TCPP-meso-tetra (4-carboxyphenyl)porphine; CCG-chemically reduced graphene; GCE-glassy carbon electrode; PPyox-overoxidized polypyrrole; Tyr-tyrosinase; CHIT-chitosan; GNS-graphene nanosheets; Bi-bismuth nanoparticles; RGO-reduced graphene oxide; Pd-NPs-palladium nanoparticles.

Conclusions
The novel Tyr-based biosensor cross-linked on cobalt (II)-porphyrin (CoP) film was fabricated for the sensitive and selective detection of dopamine. AFM images confirmed the presence of a thin film of CoP on the Au sensor surface, offering accessible binding sites for the direct immobilization of the enzyme layer. CoP has been proven to act as an electrochemical mediator during enzyme-catalyzed reaction, CV and EIS measurements showing improved electron transfer. CoP-Tyr biosensor presents high sensitivity and good stability up to seven days. A sensitivity of 1.22 ± 0.02 µA cm −2 µM −1 and a detection limit of 0.43 µM, with a linear range up to 30 µM were found, comparable with results found in the literature for nanoparticle-based sensors. In the presence of AA as an interferent, there was slight decrease in sensitivity. However, there was a separation of 130 mV between the AA-DA peaks in the DPV plot, which was more than enough to clearly discriminate between the two substances. By comparison with the CoP-sensor, the separation between peaks for the CoP-Tyr-biosensor was bigger, and the DA peaks were much higher. These facts are reflected in the higher sensitivity and lower LOD for the CoP-Tyr-biosensor towards DA in the presence of AA. Both sensors were analyzed in the presence of commercially available dopamine medication, showing good recovery and RSD, but in different linear ranges. These results suggested potential applicability of the biosensors for real samples such as human urine and blood serum.