Hypoxia commonly occurs within tumours and is a major cause of radiotherapy resistance. Clinical outcomes could be improved by locating and selectively increasing the dose delivered to hypoxic regions. Here we describe a miniature implantable sensor for real-time monitoring of tissue oxygenation that could enable this novel treatment approach to be implemented. The sensor uses a solid-state electrochemical cell that was microfabricated at wafer level on a silicon substrate, and includes an integrated reference electrode and electrolyte membrane. It gave a linear response to oxygen concentration, and was unaffected by sterilisation and irradiation, but showed susceptibility to biofouling. Oxygen selectivity was also evaluated against various clinically relevant electroactive compounds. We investigated its robustness and functionality under realistic clinical conditions using a sheep model of lung cancer. The sensor remained functional following CT-guided tumour implantation, and was sufficiently sensitive to track acute changes in oxygenation within tumour tissue. Keywords Oxygen sensor; Tumour hypoxia; Implantable; Microfabrication Page 3 of 36 1. Introduction 1 Hypoxia commonly occurs within solid tumours due to disorganised growth of blood vessels. 2 It is often spatially heterogeneous and can change with time [1, 2]. Unfortunately hypoxia 3 also causes resistance to radiotherapy, with cancer cells in low oxygen environments able to 4 withstand radiation doses 2 – 3 times higher than aerobic cells [3]. Hypoxia is also associated 5 with malignant progression and poor clinical outcome [4]. Knowledge of the location and 6 degree of hypoxia within a tumour could therefore be used to personalise and improve 7 radiotherapy by matching the dose distribution to treatment sensitivity – a concept known as 8 “dose painting” [5]. This has the potential to improve cancer treatment by optimising its 9 effectiveness while minimising side effects. Advanced techniques for achieving the necessary 10 precise dose delivery such as Intensity Modulated Radiotherapy are already in widespread 11 use [6], but there is no clinically applicable technique for continuous real-time measurement 12 of tumour hypoxia. 13 14 Existing medical imaging techniques that can measure tumour hypoxia, such as positron 15 emission tomography (PET) and magnetic resonance imaging, provide only qualitative 16 measurements, have limited spatial resolution, and give only a single time-point, limiting 17 their usefulness [7]. In addition, PET scans result in radiation exposure, limiting the 18 frequency with which they can be safely used. Quantitative imaging of hypoxia using 19 electron paramagnetic resonance imaging is under development, but will require advances in 20 readout technologies and the use of probe compounds that are not currently approved for use 21 in humans [8]. Direct measurement of oxygen within tumours is possible using the Eppendorf 22 polarographic needle electrode. This is a transcutaneous electrochemical sensor that can be 23 mechanically stepped through tissue along a linear track to measure hypoxia. It has been used 24 extensively in research settings [9], but is not clinically applicable as it requires an invasive 25 procedure for each measurement and again only gives single time-point measurements. 26 27 To overcome these limitations and deliver continuous real-time sensing for personalised 28 radiotherapy treatment, we have developed and tested a miniature implantable sensor that 29 could be inserted into multiple regions of a tumour. This would enable the construction of a 30 real-time hypoxia map, allowing personalised adaptations to radiotherapy dose distribution 31 during a course of treatment (Fig. 1) [10]. This type of implantable sensor was seen as 32 acceptable by the majority cancer patients we recently surveyed [11]. 33 Page 4 of 36 The sensor is based on electrochemical oxygen detection. This is the simplest method of 34 measuring oxygen concentration and was first described in 1897 by Danneel and Nernst [12]. 35 Following their discovery, a series of innovations culminated in the “Clark electrode” oxygen 36 sensor in 1956, which can reliably operate in biological media [13]. The Clark electrode is an 37 amperometric sensor, detecting oxygen at a platinum electrode via a reduction reaction that 38 produces a current proportional to the oxygen concentration [14]. A key development was 39 separation of its bare electrodes and their liquid electrolyte from external biological media 40 using a gas permeable membrane. This prevents dissolved proteins and metabolites from 41 interfering with the electrode and impairing its ability to sense oxygen [15]. However, 42 conventional Clark electrode sensors are still typically unsuitable for implantation as they are 43 bulky and contain a liquid electrolyte. To overcome these problems, we have developed a 44 miniature Clark-type oxygen sensor that is specifically designed for implantation. Silicon 45 based microfabrication techniques were used throughout the sensor fabrication process, as 46 they allow straightforward miniaturisation, excellent reproducibility, and low manufacturing 47 cost in volume. Crucially, they also open the possibility of future integration with 48 complementary metal-oxide-semiconductor (CMOS) instrumentation electronics. 49 50 The miniature sensor is based on a three-electrode cell, with a platinum working electrode 51 (WE) at which oxygen reduction occurs, a Ag/AgCl reference electrode (RE) that provides a 52 fixed electrochemical potential against which the WE potential can be accurately set, and a 53 platinum counter electrode (CE). We selected a Ag/AgCl RE because it is well tested in 54 biosensing applications and is more straightforward to miniaturise than other standard REs 55 [16]. The sensor surface is coated with a thin-film of the porous ionomer Nafion that allows 56 transport of water and protons [17], combining the functions of both an electrolyte matrix and 57 membrane into a single layer. This enables the device to be fully solid-state. Several other 58 designs of microfabricated electrochemical oxygen sensors have previously been described 59 [18-31], highlighting the broad interest in this technology. However, to date no design has 60 been reported that is not only amenable to wafer-level CMOS post-processing, and integrated 61 with a stable on-chip RE and membrane, but also successfully validated in an implantable 62 application. Here we describe the microfabrication and testing of a sensor platform that meets 63 all these criteria, and show proof-of-principle operation within an ovine lung tumour. 64 65 66 Page 5 of 36 2. Material and methods 67 2.1. Electrode fabrication 68 Sensors were fabricated on 100 mm diameter, n-type, <100> orientation silicon wafers 69 (Si-Mat). A bottom insulator layer of 500 nm silicon dioxide was grown by thermal 70 oxidation. Electrodes and interconnect were fabricated from a layer of 50 nm Pt on top of a 71 10 nm Ti adhesion layer, both deposited by electron beam evaporation, patterned using 72 photolithography, and etched in Ar plasma. A 1000 nm layer of aluminium was sputtered 73 onto the surface and patterned using a lift-off process to form bond pads. The wafer surface 74 was passivated with 500 nm of silicon nitride, followed by 250 nm of silicon dioxide to 75 promote adhesion of the Nafion membrane [32]. Both passivation layers were deposited 76 using high-frequency plasma enhanced chemical vapour deposition. Openings to the 77 electrode and bond pad areas were then defined in the passivation by photolithography and 78 subsequent processing in a CHF3 / Ar plasma (to etch silicon dioxide) followed by CF4 / O2 79 plasma (to etch silicon nitride). For the RE, early testing showed that bare platinum did not 80 produce a sufficiently consistent potential, so instead we used an additional Ag/AgCl layer on 81 the RE surface to provide a fixed and stable electrochemical potential. A 500 nm layer of 82 silver was deposited by electron beam evaporation and patterned over the RE area using a 83 lift-off process. To chloridise the silver, a protective layer of photoresist was first deposited 84 over the aluminium bond pads to prevent their oxidation, and then the wafer was dipped in 85 50 mM aqueous FeCl3 for 60 s at room temperature, which deposited a solid layer of AgCl on 86 the RE surface. After chloridisation the protective photoresist was removed by wet stripping. 87 88 2.2. Nafion membrane stack fabrication 89 90 The Nafion membrane was designed to cover all three electrodes, as it has previously been 91 shown to generally reduce electrode biofouling [33, 34] and to extend Ag/AgCl RE lifetime 92 [35]. We found that microfabrication of Nafion structures was challenging, as the layer was 93 easily damaged during wafer processing and packaging. We therefore developed a novel 94 method to deposit and protect it during these steps. The wafer surface was first treated in a 95 solution of 2% Silane A174 (3-(Trimethoxysilyl)propyl methacrylate) in ethanol for 5 min at 96 room temperature to promote Nafion adhesion [23, 28], then dried at 60 ˚C for 30 min. A 97 solution of 5% Nafion in lower aliphatic alcohols and water (274704, Sigma-Aldrich) was 98 Page 6 of 36 then spin coated onto the wafer at 150 rpm for 10 s, followed by 500 rpm for a further 30 s. 99 The layer was air dried at room temperature and annealed at 120 ˚C for 5 min on a hotplate. 100 SPR350 photoresist (Dow) was deposited by spin coating at 1000 rpm for 60 s, soft-baked at 101 90 ˚C for 90 s, patterned using photolithography to define the membrane area, and then 102 developed for 1 min. An MF-26A developer solution (Dow) diluted to 2:1 (developer : water) 103 was used to prevent attack of the Nafion surface [32]. The Nafion was etched in O2 plasma, 104 and the photoresist was left in place following etching as an over-etch buffer for later steps 105 (Fig. S1a). Finally, to provide solvent resistance, a temporary protective layer of 500 nm 106 Parylene-C was deposited by chemical vapour deposition, patterned using photolithography 107 to cover the Nafion layer, and etched in O2 plasma (Fig. S1b). Together these layers protected 108 the Nafion from vigorous cleaning processes during packaging. The final membrane was 109 characterised using a DektakXT surface profiler on test dies from which the protective layers 110 had been stripped. 111 112 2.3. Sensor packaging 113 114 Sensors were diced into 2 mm × 3 mm chips and mounted on the end of a custom six-track 115 flexible PCB (Merlin Flex Ltd) using Loctite 4014 biocompatible adhesive. The flexible PCB 116 had dimensions of 1.7 mm × 200 mm (for bench testing) or 1.7 mm × 500 mm (for sheep 117 implantation). Connections from the sensor bond pads to the flexible PCB were made using 118 gold wire bonds. Bonds were made from each pad to two adjacent tracks on the PCB, 119 allowing verification of the pad connection after implantation by testing electrical continuity 120 between the tracks. The sensor assembly was encapsulated in OG116-31 photocurable 121 biocompatible epoxy (Epoxy Technology, Inc.) that was selectively cured by exposure to 122 ultraviolet (UV) light to form a hermetic package around the die and bonds. The external 123 dimensions were approximately 2.8 mm × 5.1 mm × 1.4 mm (W × L × H), and an unexposed 124 region approximately 1.4 mm × 1.8 mm was left over the sensor active area. The packaging 125 dimensions were designed to be compatible with delivery through an 8G Jamshidi needle. 126 Unexposed epoxy was manually removed from the outside of the package and internal sensor 127 area using acetone. Epoxy was then deposited on the back of the sensor and cured by flood 128 UV exposure. Following encapsulation, the protective Parylene-C layer on the sensor die 129 surface was removed using O2 plasma etching, and the photoresist layer on the Nafion was 130 removed by dipping in MF-26A developer for 10 - 15 s, followed by immersion in water 131 (Fig. S1c). Finally, the epoxy encapsulation was heat cured at 80 ˚C for 2 h to ensure 132 Page 7 of 36 biocompatibility and maximum mechanical strength. The flexible PCB was connected to 133 potentiostat instrumentation through a miniature edge connector (XF3M-0615-1B-R100, 134 Omron). 135 136 2.4. Scanning electron microscopy and focused ion beam milling 137 138 The RE surface was inspected during fabrication using a Tescan Vega3 XMU scanning 139 electron microscope. Surface composition was characterised using an energy dispersive X-140 ray spectroscopy detector at 12.5 kV beam energy over a 10 µm × 7.5 µm field. For each 141 sampled area, 1×105 counts were analysed using Bruker EDS software. Die cross-sections 142 were milled using a FEI Strata 200XP focussed ion beam system with a 30 kV gallium beam. 143 Milled die samples were coated in a thin layer of evaporated gold to protect the polymer 144 layers prior to imaging. 145 146 2.5. Electrochemical characterisation 147 148 An Autolab PGSTAT12 bench potentiostat (Metrohm AG), controlled using Nova software, 149 was used for sensor characterisation. For stability and lifetime testing a MUX.MULTI4 150 multiplexer (Metrohm AG) was also used to allow sequential measurement of the sensor RE 151 potential and WE current. CV measurements were performed using a 2 mV step size and 152 100 mV/s scan rate. All reagents used for electrochemical characterisation of sensors were 153 obtained from Sigma-Aldrich, and were used as received. Chemical permeability testing of 154 the Nafion membrane was performed in a solution of 1 mM K3[Fe(CN)6], 1 mM 155 K4[Fe(CN)6], and 100 mM KCl. Sensor performance was measured in phosphate buffered 156 saline (PBS) containing 154 mM NaCl and 10 mM phosphate buffer at pH 7.4. The RE 157 response to changing pH was characterised in PBS at the same constant NaCl concentration, 158 but with different pH values set by adjusting the ratio of monobasic to dibasic sodium 159 phosphate. To characterise the RE response to chloride, aqueous NaCl was varied in 160 concentration between 10 mM and 300 mM, while holding the total ionic strength constant at 161 300 mM using NaNO3. On-chip RE potentials were measured relative to a commercial 162 Ag|AgCl|KCl (3 M) reference electrode (Sigma-Aldrich). Linearity and sensitivity of the 163 oxygen response were measured at atmospheric pressure in PBS that was bubbled with 164 different mixtures of oxygen and nitrogen. The gas composition was set using a custom gas 165 mixing rig that was manually adjusted to achieve specific dissolved oxygen concentrations, 166 Page 8 of 36 which were directly measured in solution using a Seven2Go S9 meter with an InLab OptiOx 167 probe (Mettler Toledo). Biofouling experiments were performed in air-saturated PBS 168 containing 35 mg/mL BSA, and interference experiments were performed in air-saturated 169 PBS containing either 3 µM hydrogen peroxide, 375 µM uric acid, 45 µM ascorbic acid, or 170 100 µM paracetamol. 171 172 2.6. Sensor sterilisation and irradiation 173 174 Sensors were sterilised in ethylene oxide gas produced using an Anprolene AN-73 EtO gas 175 ampoule (Andersen Products) at room temperature and pressure for 12 h, then left to degas 176 for at least 2 h before use. For irradiation testing, sensors were immersed in 1.5 cm of PBS to 177 provide a build-up region, and then exposed to a total radiotherapy dose of 24 Gy, delivered 178 in four fractions of 6 Gy at 3 min intervals, under a 6 MV beam from a Varian Clinac 2100 179 C/D linear accelerator (Varian Medical Systems). 180 2.7. Lung tumour implantation 181 Sheep studies were performed under a UK Home Office Project Licence in accordance with 182 the Animals (Scientific Procedures) Act 1986 and approval from the University of Edinburgh 183 Animal Welfare and Ethical Review Boards. Three adult female sheep (Highlander, n = 1; 184 Scottish blackface, n = 2), weighing 52 - 65 kg and diagnosed with naturally occurring pre-185 clinical ovine pulmonary adenocarcinoma, were obtained through an on-farm ultrasound 186 eradication programme [36, 37]. Sheep were housed in groups of at least two animals, bedded 187 on straw, with ad libitum access to food and water. Anaesthesia and analgesia were managed 188 by a team of specialist veterinary anaesthetists. The clinical management and surgery of these 189 cases has previously been described [38]. All animals received a central venous cannula 190 placed in the jugular vein which was used for administering drugs and crystalloid fluids. An 191 arterial cannula was also placed in the central auricular artery for monitoring arterial blood 192 pressure and to obtain arterial blood samples. Intermittent blood-gas, biochemical, and 193 haematological analysis was performed with an Epoc portable blood gas electrolyte and 194 critical care analyser (Woodley Equipment Company). Pulse rate and blood pressure along 195 with pulse oximetry, capnography, temperature, spirometry, electrocardiography and inspired 196 and expired gases (O2, CO2, and inhalant anaesthetic agent) were continuously monitored 197 using a Datex-Ohmeda S/5 multiparameter patient monitoring device (SOMA Technology). 198 Page 9 of 36 All animals were euthanised with intravenous sodium pentobarbitone (Animalcare). CT 199 imaging procedures were performed using a single-section SOMATOM Definition AS 64 200 slice helical CT machine (Siemens Healthcare) [38]. 201 All experiments were conducted on anaesthetised, non-recovery animals. Sensors were 202 inserted using a trans-thoracic percutaneous technique under CT guidance using an 8G × 203 15 cm Jamshidi biopsy needle (Carefusion). An initial CT scan was used to select an OPA 204 lesion and plan the needle path trajectory for sensor implantation. Non-metallic, radiopaque 205 grid lines (GuideLines, Oncology Imaging Systems) placed on the thoracic wall skin surface 206 and intercostal spaces aided lesion localisation and determined the site for percutaneous 207 sensor placement. Implantation was performed as previously described [38]. Briefly, a small 208 skin incision was made approximately 1 – 2 intercostal spaces caudal to the desired entry 209 point into the thoracic cavity through which the biopsy needle with its stylet in place was 210 inserted. The needle was advanced cranially through subcutaneous tissues, then pushed 211 through the chest wall at the desired intercostal space. A CT scan confirmed correct needle 212 positioning before it was seated into the OPA tumour. Serial CT scans and needle 213 advancements were made until the needle was positioned centrally within the tumour. Once 214 positioned the stylet was removed, and a sterilised sensor and lead were introduced down the 215 bore of the needle and pushed past the tip of the needle into OPA tissue. The implantation 216 needle was withdrawn, leaving the sensor and lead wire in situ which was then sutured in 217 place. A CT scan was performed to evaluate sensor positioning prior to testing, and a final 218 scan was performed post-mortem to ensure the sensor had remained in place. 219 To assess the ability of the implanted sensors to detect changes in intratumoral pO2 through 220 changes in blood oxygenation, a series of alterations in target fraction of inspired oxygen 221 (FiO2) was performed. The cycle of FiO2 alterations was: 1.00 (baseline), 0.50 (inducing mild 222 hypoxaemia), 0.21 (inducing moderate hypoxaemia) before finally returning to 1.00. Each 223 FiO2 step was performed for approximately 20 min (exact timings varied slightly according 224 to clinical need, and are shown in Fig. S5). Arterial blood samples were taken at the end of 225 each step. 226 2.8. Statistical analysis 227 Data were tested for normality using the Shapiro-Wilk test. Parametric testing on normally 228 distributed data was performed using two-tailed paired t-tests to compare between two 229 Page 10 of 36 groups, and a repeated measures one-way ANOVA followed by Dunnett’s post-hoc test to 230 compare between three or more groups. For non-normally distributed data, a two-tailed 231 Wilcoxon matched-pairs signed rank test was used to compare between two groups. Non-232 linear regression with a semi-log line was used to fit the RE response to chloride 233 concentration, and linear regression was used to fit the RE response to pH and the sensor 234 response to oxygen partial pressure. Fitted parameters were compared to theoretical values 235 using an extra sum-of-squares F-test. Statistical analysis was carried out using Prism 8 236 (GraphPad). Values in the text and figures are presented as mean ± standard deviation (SD). 237 3. Results 238 3.1. Sensor fabrication 239 Each sensor consists of a three-electrode electrochemical cell containing a 50 µm diameter 240 circular WE, surrounded by a RE and CE with a WE:RE:CE area ratio of 1:3:10. The sensors 241 were microfabricated on a silicon wafer, using conventional thin-film techniques to create 242 platinum electrodes and external aluminium bond pads, followed by production of a Ag/AgCl 243 layer on the RE by chemical oxidation of a thin-film of silver (Fig. 2a,b). Formation of AgCl 244 was monitored using scanning electron microscopy with energy dispersive X-ray 245 spectroscopy (SEM/EDX). X-ray peaks corresponding to silver were readily identified on the 246 RE area prior to chloridisation, and additional elemental chlorine peaks corresponding to 247 chloride were evident following chloridisation (Fig. 2c). Inspection of the surface using SEM 248 showed that the initially smooth evaporated silver surface was modified by chloridisation, 249 with granular structured AgCl growing on and into the electrode surface (Fig. 2d i-iii). The 250 Ag/AgCl layer was also examined in cross-section using focussed ion beam milling followed 251 by SEM. This showed a clearly visible layer of Ag/AgCl over the RE area, although there 252 was no distinct division between the two materials, suggesting either a diffuse interfacial 253 boundary or that little metallic silver remained (Fig. 2d iv). Separate investigation of the 254 chloridisation process using a series of microfabricated silver thickness test structures 255 confirmed that the latter was the case [39]. Since the Ag/AgCl layer was fabricated on an 256 underlying layer of platinum, an electrically conductive path to the RE was maintained 257 despite the highly efficient chloridisation. 258 259 The Nafion membrane was then fabricated over the electrodes. To integrate this layer with 260 subsequent wafer processing and packaging, we developed a novel series of processes in 261 Page 11 of 36 which Nafion was spin-coated on the wafer, thermally annealed, patterned using 262 photolithography and plasma etched at wafer level, and then temporarily protected using 263 sacrificial layers of photoresist and Parylene-C (Fig. 2a and Fig. S1a,b). Fabricated sensor 264 dies were wire bonded to a flexible printed circuit board (PCB) lead for connection to 265 external instrumentation, and photocurable epoxy was used to selectively encapsulate the die 266 edges, leaving an uncured window over the sensor active area (Fig. 2e,f) [40]. Finally, the 267 protective layers were removed from the Nafion membrane prior to use (Fig. S1c). The final 268 Nafion layer thickness was 518 ± 32 nm (n = 10 devices, from two wafers), measured by 269 profilometry. 270 271 3.2. Sensor characterisation 272 273 Performance of the on-chip Ag/AgCl RE in packaged sensors was characterised in PBS 274 against a commercial Ag/AgCl RE with an internal 3 M KCl filling solution. The median on-275 chip RE potential was +74.0 mV (inter-quartile range: +71.7 to +76.4 mV, n = 81 sensors, 276 from two wafers), confirming that the fabrication process had produced a functional and 277 reproducible RE. We hypothesised that the small positive potential difference was due to the 278 lower chloride activity in PBS contacting the on-chip RE compared to the concentrated KCl 279 filling solution of the commercial RE, rather than a Donnan membrane potential [41]. To 280 confirm this, the concentration of sodium chloride in solution bathing the sensor was varied 281 while holding the ionic strength constant using sodium nitrate to maintain constant activity. 282 The on-chip RE potential changed with the logarithm of chloride concentration, giving a 283 slope of −58.6 mV/dec (Fig. S2a). This was not significantly different to an ideal Nernstian 284 response of −58.2 mV/dec, supporting the hypothesis. The response of the RE to changing 285 pH was also evaluated in PBS over a physiological range of pH 6.2 – 7.8, and showed no 286 significant variation, as expected (Fig. S2b). 287 To establish whether the Nafion layer was acting as a protective sensor membrane in the 288 packaged sensors, we performed cyclic voltammetry (CV) at the WE in a solution of 289 potassium ferricyanide and potassium ferrocyanide. The ferri/ferrocyanide anions should be 290 excluded from Nafion (due to their size and/or charge), and thus blocked from reacting at the 291 WE. Consistent with this, no redox processes were visible in the cyclic voltammograms from 292 Nafion coated sensors. In contrast, control devices lacking a Nafion membrane showed clear 293 peaks corresponding to the reduction and oxidation of the ferro/ferricyanide in solution, 294 Page 12 of 36 indicating their presence at the electrode surface (Fig. 3a). These results show that the 295 fabricated Nafion membrane fully covered the WE surface and that no defects were 296 introduced into the layer during packaging. In addition, they confirm that the epoxy 297 packaging was effective at hermetically sealing the bond pads and wire bonds from contact 298 with external solutions. 299 To characterise the electrochemical performance of the complete sensor system we 300 performed CV in PBS using the on-chip WE, RE and CE. The WE potential was swept 301 between −0.7 V and +0.9 V (vs the on-chip Ag/AgCl RE), remaining within the 302 electrochemical potential window of the system. An initial slight peak then pronounced wave 303 at approximately −0.5 V was visible, and was absent after purging oxygen from the PBS 304 solution using nitrogen (Fig. 3b). This indicates that the peak was caused by oxygen 305 reduction at the WE surface, and the wave by diffusion limited oxygen reduction through the 306 Nafion membrane. These results confirmed that the annealed Nafion membrane was capable 307 of permitting oxygen and small ion transport. Additional peaks were frequently evident in the 308 PBS CV at approximately +0.07 V and −0.15 V, which are characteristic of silver chloride 309 and may indicate some mechanical transfer from the RE onto the WE surface during 310 fabrication. However, these were well separated in potential from the oxygen reduction 311 reaction, and were not observed to interfere with its position or magnitude. 312 We next performed chronoamperometry (CA) at −0.5 V (vs the on-chip Ag/AgCl RE) for 313 20 s. In air-saturated PBS the WE current showed an expected initial transient related to 314 capacitive charging, reduction of oxygen, and formation of oxygen diffusion gradients at its 315 surface (consistent with the CV peak). This was followed by the development of an 316 approximately steady-state current caused by diffusion-controlled oxygen reduction 317 (consistent with the CV wave). In nitrogen purged PBS the steady-state current was markedly 318 decreased, again consistent with the CV results and identifying the current as oxygen 319 reduction (Fig. 3c). To investigate the linearity of the oxygen response, its partial pressure 320 (pO2) in solution was varied from 0.1 to 25.3 kPa (equivalent to 0.1 to 25.0% of atmospheric 321 pressure) by sparging with a range of oxygen/nitrogen mixtures. A greater sampling 322 resolution was used in the lower range relevant to biological tissue. The sensor response to 323 oxygen was quantified using the mean current over the final 5 s of a 20 s CA recording, as 324 the steady-state was well established by this point. Linear regression showed a proportional 325 relationship between the steady-state current and oxygen concentration (r2 = 0.991), with a 326 Page 13 of 36 fitted sensitivity of −0.595 ± 0.009 nA/kPa and a statistically insignificant offset of −0.143 ± 327 0.100 nA (Fig. 3d). Quantification of oxygen concentration as the mean steady-state CA 328 current at −0.5 V (vs on-chip Ag/AgCl RE), averaged over the final 5 s of a 20 s recording, 329 was therefore adopted for all further measurements. Together these results show that the 330 sensor output varies directly with oxygen concentration as expected, with insignificant offset 331 due to unrelated currents at the WE. 332 To investigate the performance of the packaged sensors over time, oxygen measurements 333 using CA were repeatedly performed in air-saturated PBS. The RE potential was also 334 recorded from the same devices against an external Ag/AgCl RE in the rest periods between 335 CA measurements. Over 4 h the sensors gave an average steady-state current of 336 −11.4 ± 0.8 nA (n = 8 sensors) in air-saturated PBS (Fig. 3e). To quantify individual sensor 337 stability, the coefficient of variance (CoV, defined as [SD / Mean] × 100%) of the output 338 from each sensor was also calculated, and gave a mean CoV of 12.2% over the 4 h 339 measurement period (n = 8 sensors). The RE potential remained very stable over the same 340 period (Fig. 3f), with an average value of +74.8 ± 1.5 mV (n = 8 sensors), and a mean CoV of 341 only 0.33% (n = 8 sensors). The measurements were then continued until sensor failure to 342 investigate their lifetime. Ultimately the first element of the sensor to fail was the RE. A 343 change in RE potential typically occurred quite suddenly, leading in turn to unreliable CA 344 oxygen measurements as the WE bias could not be accurately set (Fig. S3). A difference in 345 RE potential of more than ±50 mV from its initial value was defined as the failure point, and 346 gave a mean time-to-failure of 23.3 h (range 5.1 – 64.5 h, n = 8 sensors). 347 We next systematically explored how sensor performance may be affected by exposure to the 348 in vivo environment. To test whether the sensor was susceptible to protein biofouling, we 349 recorded oxygen measurements from sensors in air-saturated PBS containing bovine serum 350 albumin (BSA). This is a well characterised, water soluble protein, commonly used as a 351 surrogate to model biofouling. Air-saturated PBS alone was used as a control. Sensors 352 exposed to BSA showed a significant decrease in steady-state WE current after 24 h (Fig. 4a), 353 suggesting that biofouling due to non-specific protein adsorption may lower the sensitivity of 354 the Nafion coated sensors. We then tested the interference effect of common electroactive 355 compounds (hydrogen peroxide, urate, ascorbate, and paracetamol) that are found in vivo and 356 known to affect electrochemical measurements. Compounds were dissolved in air-saturated 357 PBS, and oxygen measurements were sequentially performed in each solution in a 358 Page 14 of 36 randomised order. When compared to PBS alone, we found that physiological concentrations 359 of hydrogen peroxide (3 µM), ascorbate (45 µM), and paracetamol (100 µM) all had no 360 significant effect on the steady-state WE current, while urate (375 µM) caused a significant 361 decrease in the magnitude of the current (Fig. 4b). 362 As the sensor was designed to be surgically implanted and remain in place during a course of 363 radiotherapy, we finally tested whether it was adversely affected by sterilisation or 364 irradiation. Ethylene oxide gas was used to sterilise the sensors as it is routinely used to 365 sterilise medical equipment. Irradiation was performed under a clinical radiotherapy beam, 366 using a dose of four fractions of 6 Gy. This fraction size is within the upper range typically 367 delivered in human clinical use [42], and matches that used in previous work involving a 368 sheep lung cancer model [43]. Following these treatments, the on-chip RE potential was 369 unchanged (Fig. 5a), and the Nafion membrane was still impermeable to ferri/ferrocyanide 370 anions (Fig. 5b), indicating that its physical integrity and selectivity had not been 371 compromised. Using CV in PBS we observed that the expected oxygen reduction features 372 were still evident (Fig. 5c), and that the corresponding CA steady-state oxygen current at 373 −0.5 V (vs the on-chip Ag/AgCl RE) was also unchanged (Fig. 5d), showing that the 374 electrodes were undamaged. 375 3.3. In vivo validation 376 Bench characterisation showed that the sensor could report oxygen concentration in solution 377 effectively over many hours. We therefore next tested whether the sensor was sufficiently 378 robust to operate in vivo and detect dynamic changes in tissue oxygenation. We used a novel 379 clinically relevant ovine model of lung cancer [38], developed specifically for this project. 380 Tumours in three sheep with naturally occurring pre-clinical ovine pulmonary 381 adenocarcinoma were implanted with either one sensor (Cases F1 & F2) or two sensors (Case 382 F3). Each sensor was implanted under anaesthesia using a Jamshidi needle using CT 383 guidance (Fig. 6a,b). The sheep were initially maintained at a target FiO2 of 1.00, then 384 exposed to a protocol of mild graded tissue hypoxia by stepping their FiO2 down to 0.50 and 385 then 0.21, before being returned to an FiO2 of 1.00 (Fig. 6c). Arterial blood analysis showed 386 the expected changes in oxygenation (Fig. 6d and Fig. S4a), while other relevant 387 physiological variables did not change significantly (Fig. S4b-e). During the protocol, 388 measurements were repeatedly made from the sensors using CA and all gave measurable 389 real-time outputs (Fig. S5). Interestingly, there was substantial variability between cases and 390 Page 15 of 36 so readings from each sensor were analysed independently. Both sensors in Case F3 showed 391 low initial currents which did not respond to changes in FiO2, while the sensors in Cases F1 392 and F2 showed a relatively higher initial current with partially reversible decreases at lower 393 FiO2 values and a decrease following euthanasia (Fig. 6e). These differences reflect the 394 expected heterogeneity of tumour hypoxia and oxygen responsiveness. 395 To confirm that the sensors survived the tumour implantation procedure they were recovered 396 following euthanasia, gently rinsed in water, and tested. Their surfaces showed the presence 397 of some contamination (Fig. 7a), likely derived from lung tissue at the implant site. CV in air-398 saturated PBS showed the expected presence of oxygen reduction at negative potentials for 399 all sensors (Fig. 7b). The mean steady-state CA current showed greater variability between 400 sensors post-implantation but was not significantly different to its pre-implantation value 401 (Fig. 7c). Together these results demonstrate that the packaged sensor is sufficiently 402 physically robust to remain functional throughout a clinically realistic implantation 403 procedure. 404 4. Discussion 405 Together, this work describes a microfabricated electrochemical oxygen sensor that is 406 suitable for implantation, and demonstrates its functionality and relevance for real-time 407 measurement in a realistic clinical environment. The wafer-level microfabrication approach 408 we adopted for producing the sensor has two advantages. Firstly, it enables the sensor to be 409 mass manufacturable with high repeatability through use of standard industrial tools and tight 410 process control, and minimises costly and potentially inconsistent die-level processing. These 411 features will be very important for gaining medical regulatory approval, as sensor accuracy 412 and repeatability are essential for making safe clinical decisions. Secondly, it will enable our 413 future aim of post-processing the oxygen sensor architecture on CMOS wafers to produce a 414 device that contains integrated potentiostat instrumentation and wireless power and 415 communication electronics. This will also allow the volume of the device to be minimised by 416 eliminating dependence on the existing assembly and packaging elements required for 417 connection to external instrumentation. 418 The sensor membrane is a defining component of the Clark electrode, separating its bare 419 electrodes from the chemically complex surrounding fluid. Our membrane was fabricated 420 from Nafion, a perfluorinated cation exchange resin that is both oxygen permeable and can 421 Page 16 of 36 conduct protons between electrodes to support the electrochemical reactions [17]. Wafer-422 level spin coating was used to deposit the Nafion membrane, followed by heat annealing. 423 Annealing increases Nafion solvent resistance [44], improving integration with standard wet 424 microfabrication processes and improving its stability in aqueous environments. Furthermore, 425 it also improves Nafion proton conductivity and electrochemical surface area [45, 46], tensile 426 strength [45], and the lifetime of implanted Nafion coated Ag/AgCl reference electrodes [35]. 427 Photolithographic patterning and dry etching were then used to accurately define the 428 membrane areas across the wafer. Together this novel combination of processes significantly 429 improved manufacturability compared to previously described microfabricated oxygen 430 sensors that used die-level drop cast Nafion membranes that were not annealed [21, 23, 28]. 431 Selectivity of the Clark sensor membrane is important for sensor accuracy. The polymer 432 backbone of Nafion carries negatively charged sulphonate groups that can electrostatically 433 exclude anionic species, and the nanometer scale of its hydrophillic pores causes size 434 exclusion of large ions [17]. Consistent with these effects, we observed that our membrane 435 blocked ferricyanide ([Fe(CN)6]3−) and ferrocynanide ([Fe(CN)6]4−) ions from reacting at the 436 WE. In addition, no significant interference was observed from physiological concentrations 437 of ascorbate, hydrogen peroxide, or paracetamol. We found that urate decreased the measured 438 current from the sensor, indicating that it may interfere with oxygen measurements. There is 439 conflicting literature regarding Nafion permeability to urate, with several reports showing 440 that Nafion improves selectivity for catecholamine neurotransmitter measurements in the 441 presence of urate [47], while others have shown that it can be used to block ascorbate and 442 allow selective measurement of urate [48-50]. Our data suggest that the Nafion sensor 443 membrane is permeable to urate, likely in its protonated uric acid form as it would otherwise 444 be electrostatically excluded, and may become deposited within the hydrated pores limiting 445 mass transport. 446 An additional function of the membrane was to prevent electrode biofouling. Nafion is 447 generally thought to reduce biofouling [33, 34], however we observed that exposure of the 448 sensor to a solution of BSA caused a significant decrease in its oxygen sensitivity over 24 h. 449 This is likely to be due to surface adsorption of protein, since BSA will be size excluded from 450 the Nafion membrane pores. Biofouling may partly account for the measured variability in 451 the performance of sensors that were recovered at post-mortem following tumour 452 implantation, and it will need to be addressed before translation to human clinical trials. 453 Page 17 of 36 An essential component of any electrochemical sensor is the RE, necessary to provide the 454 potential against which the WE potential is measured or set. Key parameters of RE 455 performance are potential stability, sensitivity to environmental variables, and lifetime. Our 456 Nafion coated Ag/AgCl RE successfully provided a sufficiently stable potential to enable 457 accurate measurements to be made both in vitro and in vivo. Its potential was set by the 458 surrounding chloride concentration in a typical Nerstian response, as others have previously 459 observed for reference electrodes without a defined internal electrolyte [16]. This is not 460 expected to cause problems in an implanted setting, as chloride concentration is under tight 461 homeostatic control, with a typical range of 97 – 107 mM in human plasma [51]. During 462 long-term testing, we typically saw failure (defined by a sudden large potential change) of the 463 RE in less than a day, leading to unreliable measurements. This may reflect mechanical 464 instability of the AgCl layer, or exhaustion of either Ag or AgCl. The failure origin will be 465 investigated, since the short lifetime of our sensor remains a limitation that will need to be 466 improved for clinical applications. 467 The sensor was designed from the start for implantation, and so all the outward facing 468 materials were selected for their established biocompatibility, including silicon dioxide [52-469 54], Nafion [55-57], and Epotek OG116-31 epoxy resin. We have recently demonstrated that 470 implantation of these sensor materials does not cause changes in tumour pathology [58], 471 which will be essential for clinical deployment. Others have previously shown that a large 472 implanted Ag/AgCl RE can cause local toxicity. To minimise this risk, we used a Nafion RE 473 coating which is known to reduce the toxicity of implanted Ag/AgCl electrodes [35]. The 474 PCB insulation was manufactured from polyimide, which is already commonly used for 475 encapsulating and insulating implantable medical devices [59]. 476 To validate the sensor in tumour tissue we used an ovine pulmonary adenocarcinoma model. 477 This model has high translational value, as it closely reflects both human lung cancer biology 478 [60], and the surgical methods used for tumour management such as trans-thoracic needle 479 biopsy [38]. The experiments described here represents the first use of this novel preclinical 480 model for validation of an implantable sensor in tumour tissue. Oxygen sensor readings 481 obtained from the implanted sensors showed substantial variation between cases, consistent 482 with a poorly perfused hypoxic tumour environment at both sensor locations in Case F3, but 483 better perfusion and responsiveness to changes in blood oxygenation at the sensor locations 484 in Cases F1 and F2. This underlines the known spatial heterogeneity of hypoxia that is found 485 Page 18 of 36 in solid tumours [1]. Interestingly, our results are also consistent with a previous human lung 486 cancer study which showed highly variable responsiveness of tumour hypoxia to breathing 487 95% oxygen as a potential radiotherapy sensitising technique [61]. 488 Hypoxia is also a feature of many other pathological states. These include chronic conditions 489 such as non-healing wounds [62], while acute disruption occurs in life-threatening conditions 490 such as hypovolaemic, cardiogenic and septic shock [63], and following traumatic brain 491 injury [64]. Our minimally invasive technique for continuously measuring tissue oxygenation 492 could therefore also be a valuable tool for clinicians involved in the treatment and 493 management of a broad range of other medical conditions. We are currently engaged in 494 further implantation studies to determine whether the sensor is suitable for post-surgical 495 monitoring of intestinal tissue oxygenation following resection and anastomosis [65]. 496 497 5. Conclusion 498 Hypoxia is distributed heterogeneously within solid tumours, and can vary over time. It 499 inhibits the effectiveness of radiotherapy treatment, arguing for tumour oxygenation status to 500 be included in radiotherapy planning. Here, we describe a miniature implantable sensor for 501 continuous real-time monitoring of tumour hypoxia. The fabrication process comprises only 502 CMOS compatible wafer-level processing techniques, and produces a novel layer stack 503 suitable for integration with biocompatible epoxy encapsulation. The sensor gave a linear 504 response to oxygen, although it showed some susceptibility to biofouling and interference by 505 urate, and its lifetime was limited to between hours and days by failure of the on-chip RE. 506 We showed that the packaged sensor was sufficiently robust to function over this timescale 507 within a tumour using a translational model of lung cancer, and demonstrated measurement 508 of dynamic changes in tissue oxygenation in vivo. Our future work will focus on addressing 509 the limitations of sensor lifetime and selectivity, alleviating the effects of biofouling, and 510 ultimately achieving full integration with on-chip instrumentation and wireless power and 511 communications. 512 Page 19 of 36 Acknowledgements 513 This work was supported by funding from the UK Engineering and Physical Sciences 514 Research Council through the Implantable Microsystems for Personalised Anti-Cancer 515 Therapy (IMPACT) programme grant (EP/K034510/1), from a project grant from Bowel and 516 Cancer Research UK, and from a Wellcome Trust Biomedical Resource Grant to the 517 Wellcome Trust Critical Care Laboratory for Large Animals (104972/Z/14/Z). We are 518 grateful to S. Ramsay and R. Blair for their technical assistance during sensor fabrication, E. 519 MacDonald for constructing the O2/N2 gas mixing apparatus used for sensor characterisation, 520 I. Schmueser for advice on electrochemistry, E. Kay and H. Porter for assistance and advice 521 during sensor irradiation testing, J. Nixon, P. Tennant and A. Ritchie for sheep husbandry and 522 sensor sterilisation, and L. Grant for CT imaging. Sheep OPA cases were obtained by P. Scott 523 (Capital Veterinary Services). 524 Author contributions 525 AFM, DJA, AJW, ARM & IHK conceived the approach. CD, JRKM, EOB, AT & EG-F 526 designed and fabricated the sensors. JRKM, CD, EOB & MEG performed the sensor 527 characterisation experiments. MMP performed the sensor irradiation. MEG, AD, MS & MAP 528 developed methodology for in vivo recording. MEG, PS, SNG, RG, REC & DJA developed 529 methodology for the lung tumour implantation experiments and performed the sheep in vivo 530 work. JRKM & ARM analysed the data. AFM, DJA, AJW, IU, ARM, SS, JGT, SM, MAP, 531 MS & REC supervised the experiments. JRKM & MEG wrote the manuscript, and it was 532 reviewed and edited by all the authors. 533 Competing interests 534 The authors declare no competing interests. 535 Page 20 of 36