**1. Introduction**

In comparison with other methods of detection such as optical, spectroscopic, and chromatographic, electrochemical sensors possess advantages such as simplicity, rapid response times, and high sensitivity [1]. Electrochemical sensors can be easily adapted for the detection of a wide range of analytes and can be incorporated into robust, portable, low cost, minituarized devices that can be tailored for particular applications [2]. Taking advantage of these attributes and the incorporation of highly specific biological recognition elements (enzymes, nucleic acids, cells, tissues, and so on), electrochemical biosensors are capable of selectively detecting a broad range of target analytes. As defined by IUPAC, electrochemical biosensors are "self-contained integrated devices, which are capable of providing specific quantitative or semi-quantitative analytical information using a biological recognition element (biochemical receptor), which is retained in direct spatial contact with an electrochemical transduction element" [3]. Bioelectrochemical sensors are used in environmental monitoring, healthcare, and biological analysis, among others. Depending on the recognition process, biosensors can be subdivided into two main categories: affinity and biocatalytic sensors. Affinity sensors operate via selective binding between the analyte and the biological component (i.e., antibody and nucleic acid) [4]. In contrast, biocatalytic devices incorporate enzymes, whole cells, or tissue slices that recognize the target analyte, and subsequently produce an electroactive species [5].

The first biosensor was described by Clark and Lyons in 1962 [6]. This biosensor was composed of an oxygen electrode, an inner oxygen semipermeable membrane, and a thin layer of glucose oxidase (GOx, EC 1.1.3.4) entrapped by a dialysis membrane. The decrease in the level of oxygen was proportional to the concentration of glucose resulting from the enzyme catalysed oxidation of β-D-glucose to β-D-glucono-δ-lactone [7]. Since this pioneering work, extensive efforts have been made to develop electrochemical biosensors for a wide range of analytes. Bioelectrochemical sensing devices have been effectively transferred from the laboratory to the point-of-care (POC) with global sales growing from less than \$5 million per annum [8] thirty years ago to over \$18 billion in 2018. Although commercial systems are available for a range of small molecules (lactose, uric acid, cholesterol, lactate, ketone, and so on), the market is dominated by glucose sensors, with approximately 90% of the market associated with glucose monitoring for diabetes [9]. Diabetes mellitus is one of the leading causes of death and disability in the world [10]. It is a metabolic disorder that causes insulin deficiency and hyperglycemia, resulting in blood glucose concentration deviating from the normal range of 3.9–6.2 mM [11]. According to the International Diabetes Federation (IDF), the number of diabetic patients increased from 151 million in 2000 to 415 in 2015. The IDF also predicted that the number of diabetic patients would increase to 642 million in 2040, with diabetes becoming the seventh-leading cause of mortality [12]. Commercial home use blood glucose sensors generally detect glucose in the concentration range of 1.1–33.3 mM with test times of less than 30 s [4]. GOx is widely employed as the recognition element in glucose biosensors owing its relatively low cost, high selectivity, and stability [13]. First and second generation sensors rely on the immobilization of the enzyme onto an electrode surface. As the redox active centers of GOx are at least 13–18 Å from the electrode surface, mediators are employed to shuttle electrons between the electrode surface and enzyme's active site [14]. The direct oxidation of GOx occurs in third generation sensors, where the enzyme is specifically wired to minimize the distance between the active site of the enzyme and the electrode surface. Despite the considerable progress that has been made, the majority of commercial glucose sensors are based on second generation glucose sensors. The vast majority of commercial devices utilise blood samples obtained via a finger prick. The development of glucose biosensors based on the detection of glucose in fluids such as tears [15], saliva [16], and sweat [17] has been described. Such systems face challenges, in particular the poor correlation between glucose levels in blood and in other fluids and also significantly lower concentrations of glucose in fluids such as tears. Individually optimized designs must be developed [18] for commercially viable sensors, where challenges such as low cost, ease of manufacture, robustness, and portability are additional factors for consideration [19].

In contrast, detection of larger biomolecules such as nucleic acids and proteins faces significant additional challenges that include electrode fouling, non-specific adsorption of biological components at the electrode surface, lack of sensitivity in the appropriate concentration range, and in particular at low concentrations (femtomolar to attomolar) [8]. Commercial systems for the detection of larger biomolecules are dominated by pregnancy tests that rely on the detection of human chorionic gonadotrophin (hCG), a glycoprotein hormone secreted during pregnancy [20].

In this review, we describe recent advances on the use of materials as supports in electrochemical biosensors, and in particular the use of materials such as nanoporous metals, graphene, carbon nanotubes, and mesoporous carbon. Examples of the detection of clinically relevant molecules are provided, with a focus on the detection of glucose. An overview of invasive and non-invasive glucose monitoring with case studies is given. In addition, we discuss the use of additive manufacturing for electrochemical sensing applications.

#### **2. Electrode Materials**

Owing to their intrinsic conductivity, biocompatibility, and ease of manufacture, high surface area materials such as nanoporous gold, graphene, carbon nanotubes, and mesoporous carbon have been used for the preparation of electrodes for bioelectrochemical applications.

#### *2.1. Nanoporous Metal Electrodes*

Nanoporous metals are 3D bicontinuous structures with tuneable pore diameters and lengths that possess large surface areas, mechanical resistance, and high conductivity [21,22]. Although nanoporous electrodes have been prepared using a range of metals such as copper, silver, and palladium, the majority of research has focused on nanoporous gold (NPG) owing to its ease of manufacture, chemical stability, and biocompatibility [23]. NPG is a 3D nanostructured material with pore sizes

that can be tuned over the range 5 nm to greater than 2 μm [14]. The morphology of NPG is generally characterised using atomic force microscope (AFM) and scanning or transmission electron microscopy (SEM/TEM). The electrochemically addressable accessible surface area is evaluated by measuring the roughness factor calculated from the charge associated with reduction of gold oxide in 0.5 M H2SO4 solution and applying a conversion factor of 390 μC/cm<sup>2</sup> [24]. NPG electrodes possess good electrical conductivity, catalytic activity, high surface-to-volume ratio, permeability, chemical, and thermal and mechanical stability [25,26], as well as properties of interest for a range of applications including biocatalysis [27], nucleic acid sensors [28], enzymatic sensors [29], non-enzymatic sensors [30], immunosensors [31], supercapacitors [32], enzymatic fuel cells [33], and so on.

Different methods have been studied for the controlled manufacture of nanoporous gold [33]. For example, using anodization methods, the 3D structure is generally formed by the anodization of gold in oxalic acid at different applied potentials, which enables the formation of specific nanoporous structures [34]. Recently, a NPG microelectrode was fabricated via electrochemical anodization-reduction steps in 0.5 M H2SO4, exhibiting pore sizes in the range of 30–50 nm [35]. Although anodization of gold avoids the use of corrosive chemicals, the pore diameters are typically ca. 20 nm in size [36], making it potentially difficult to achieve high loading of biomolecules. Another route entails using hydrogen bubbles formed via the electrochemical reduction of H<sup>+</sup> as the template [37]. Gabriella Sanzo et al. synthesized a gold nanocorals porous structure with an electroactive area 500 times higher than a gold screen printed electrode that was used as the base substrate [38]. The nanocoral electrode was modified with glucose oxidase for the development of an enzymatic biosensor based on the detection of H2O2. The nanocoral electrode showed a sensitivity of 48.3 μA/mMcm2, two times higher than that of the bare gold electrode. The hydrogen template produces materials with pore sizes in the micrometer region. In order to overcome the limitation on pore size, other template routes can be used. The hard template route usually involves two steps: assembly of monodisperse spheres, then electrodeposition of the metal followed by removal of the hard template, where the diameter and thickness of the porous structure are controlled in the range of 100–2000 nm [39]. The spatial arrangemen<sup>t</sup> and size of the pores can be controlled using colloidal crystals as a template. For instance, Szamocki et al. fabricated macroporous gold electrodes of different sizes for the electrochemical oxidation of glucose with glucose dehydrogenase (GDH), with an enhanced electrochemical response by more than one order of magnitude compared with planar gold electrodes [39]. Gamero et al. immobilised lactate oxidase (LOx) on NPG with a pore size of 500 nm, with a linear response observed up to a concentration of 1.3 mM [40].

An alternative approach relies on chemical dealloying of the less noble metal of an alloy, which can be prepared by sputtering a gold-metal alloy onto a support or by using commercially available gold alloys, for example "white gold". During the dealloying process, atoms of the less noble metal are detached from the surface and subsequently dissolved under the etchant conditions, forming nanoporous structures. Different alloy systems including Au-Zn [41], Au-Ni [42], Au-Al [43], Au-Si [44], and Au-Ag [45], have been used for the formation of nanoporous gold by dealloying the least noble metal component. Au-Ag is the most commonly used owing to the ease of removal of silver, which is generally removed under corrosive conditions (usually 70% nitric acid). In a systematic study, different alloy compositions Ag70Au30, Ag50Au50, and Ag35Au65 were prepared (Figure 1 A–D) [46]. The silver content in the alloy Ag35/Au65 was too low to enable nanoporous structures to be formed. A homogeneous distribution of nanopores was formed using the Ag70Au30 alloy. The thickness and composition of the layer were controlled by the sputtering conditions, while the pore sizes were controlled by factors such as the time period and the temperature of the process. For instance, by varying the temperature and time of dealloying of a 100 nm thick Ag70Au30 alloy, the pore size of the dealloyed sheets ranged from 4 to 78 nm, with a maximum surface area 44 times greater than the geometric area [46]. NPG prepared using this approach exhibits a controllable pore size range from 5 to 700 nm [47], a range sufficiently large to accommodate biomolecules. As with planar gold electrodes, the surfaces of NPG can subsequently be modified. For example, carboxylic acid terminated

diazonium salts were covalently attached onto NPG and the immobilization of fructose dehydrogenase (FDH) was subsequently accomplished via crosslinking with CMC [48]. The sensor showed a linear range of 0.05–0.3 mM, with a sensitivity of 3.7 μA/cm<sup>2</sup> mM and a limit of detection (LOD) of 1.2 μM with a fast response of less than 5 s. The linear range encompasses that observed in juices and the sensor displayed excellent selectivity.

**Figure 1.** Schematic representation of the manufacture of nanoporous gold (NPG) electrodes with (**A**) different layers and thicknesses, (**B**) sputtered glass sheet prior to etching, (**C**) formation of nanopores after etching, and (**D**) the completed NPG electrode. Adapted from [46].

Wearable sensors have the potential to play a major role in the development of continuous monitoring for glucose and other biomarkers in different fluids such as tears, saliva, interstitial fluids, and sweat. Flexible NPG was prepared using an electrochemical dealloying approach. NPG electrodes were further modified with lactate oxidase and bilirubin oxidase to develop a lactate/O2 enzymatic fuel cell, which was successfully tested in artificial lachrymal fluids [33].

Matharu et al. described the fabrication of NPG with different pore sizes via dealloying of a 600 nm thick Au-Ag alloy to investigate DNA hybridization in the presence of biofouling species [49]. The thiolated capture probe DNA and its target DNA were used to investigate hybridization using methylene blue as intercalator. In the absence of biofouling conditions, the accessibility of target DNA increased with larger pores, resulting in higher signal suppression with maximum values of ∼70% of that for a pore size of about 45 nm. However, in the presence of biofouling conditions, electrodes with average pore sizes of 25–30 nm maximized the accessibility of target DNA as the pores were sufficiently small to block the entrance of biofouling molecules. In contrast, larger pores were susceptible to electrode blockage by biofouling, decreasing the biosensor performance.

Owing to the expensive nature of gold, electrodes have been manufactured using lower cost non-noble metals such as Cu, Ni, Ti, or Fe [23]. However, the reproducible preparation of nanoporous structures from such alloys needs to be addressed [50].

NPG is the most widely used metal support used for biosensing, with reviews on the preparation and application of nanoporous gold published recently [14,51]. The high surface area per volume, biocompatibility, and the ability to prepare flexible electrodes make NPG an attractive material for use with biological systems. However, the high cost of gold and the complexity of the manufacturing process currently limit the applicability of NPG to research applications [43].

#### *2.2. Carbon Based Materials*
