Biochemical piezoresistive sensors based on pH- and glucose-sensitive hydrogels for medical applications

Ulrike Schmidt; Margarita Guenther; Gerald Gerlach

doi:10.1515/cdbme-2016-0029

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Biochemical piezoresistive sensors based on pH- and glucose-sensitive hydrogels for medical applications

Ulrike Schmidt , Margarita Guenther and Gerald Gerlach

Published/Copyright: September 30, 2016

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From the journal Current Directions in Biomedical Engineering Volume 2 Issue 1

Abstract

Many conventional analysis techniques to detect chemical or biological species are able to achieve a high detection sensitivity, however, they are equipment- or time-expensive due to a multi-step procedure. In this work we describe sensor concepts using piezoresistive pressure sensor chips with integrated analyte-sensitive hydrogels, that enable inexpensive and robust biochemical sensors which are miniaturizable and in-line capable. Biocompatible hydrogels were developed and tested for pH- and glucose-monitoring during the chemical and biochemical processes. For that, monomer mixtures based on hydroxypropyl methacrylate HPMA, 2-(dimethylamino)ethyl methacrylate DMAEMA, tetraethylene glycol dimethacrylate TEGDMA and ethylene glycol EG were photo-polymerized. By means of carbodiimide chemistry, glucose oxidase was bound to the pH-sensitive HPMA/DMAEMA/TEGDMA/EG hydrogel squares causing the glucose-sensitivity. The crosslinked hydrogels were integrated in piezoresistive pressure sensors of different designs. pH- and glucose-depending reversible gel swelling processes were observed by means of the output voltage of dip sensors and of a novel implantable flexible sensor set-up. Due to its biocompatible components, the latter could be used inside the human body monitoring physiological blood values, for example glucose.

Keywords: biochemical sensor; glucose-sensitive; hydrogel; medical applications; pH-sensitive; piezoresistive

1 Introduction

Hydrogel-based biochemical piezoresistive dip sensors provide the opportunity for a continuous monitoring and on-line control of analyte concentrations in ambient aqueous solutions. These sensors reduce the volume needed for medical analysis from the ml to the nl range and are useful for the diagnosis and monitoring of diabetes [1]. Inside the hydrogel, a chemical reaction takes place due to the analyte concentration change resulting in a pH change, leading to a swelling or shrinking of the pH-sensitive hydrogel. The swelling itself can be detected by means of several methods.

This work focuses on piezoresistive silicon pressure sensor chips which measure the corresponding swelling pressure. The sensor chip incorporates the pH-sensitive, biocompatible, three-dimensional HPMA/DMAEMA/TEGDMA/EG hydrogels conjugated with glucose oxidase using EDC/NHS (1-Ethyl-3-(3-dimethylaminopropyl)-carbodiimide/N-hydroxysuccinimide). In the presence of oxygen and water, glucose oxidase catalyses the enzymatic reaction of glucose into gluconic acid and hydrogen peroxide (see eq. 1).

(1)Glucose+O2+H2⁢O→GluOxGluconic⁢acid+H2⁢O2.

Here, an increase of the glucose concentration causes a lowering in pH value [2]. Hence, the glucose oxidase loaded pH-sensitive hydrogel changes its swelling state depend on the surrounding glucose concentration. The corresponding volume change of the hydrogel integrated in the cavity of a piezoresistive pressure sensor is monitored by the electrical output voltage V_out.

Hydrogel-based piezoresistive dip sensors are inexpensive, robust, miniaturizable, reliable, reproducible, and long-term-stable. These advantages could be useful also for medical applications of sensors inside the human body. In this work, an implantable sensor set-up based on biocompatible polyimide foils and on a medical grade silicone encapsulation was demonstrated.

2 Sensor set-ups

For the design of the piezoresistive chemical sensors, commercially available pressure sensor chips (C41-Series, Epcos, Munich, Germany) were used. The sensors operate by monitoring analyte-induced volume changes of a thin polymeric hydrogel disc used as chemo-mechanical transducer. The hydrogel deflects a flexible thin silicon bending plate (see Figure 1B). A piezoresistive Wheatstone bridge integrated at the plate surface works as a mechano-electrical transducer and transforms the plate deflection w into an electrical output signal V_out. A crosslinked and conditioned hydrogel square was placed into the cavity at the back side of the silicon chip. The cavity is closed with a porous, biocompatible and hydrophilic Al₂O₃ membrane (pore size of 0.2 μm, Anopore^™, Structure Probe, SPI Supplies, West Chester, USA). This membrane keeps the hydrogel inside the cavity and provides a low protein binding for implanted sensors. Solution and analyte molecules diffuse through the membrane into the chip cavity and induce a swelling or shrinking of the hydrogel. This swelling pressure change leads to a deflection change of the bending plate and causes a change of the stress state inside the plate, which in turn effects a change of the resistivity of the piezoresistors affecting proportionally the output voltage V_out of the sensor. A hydrogel swelling results in an increasing, a deswelling in a decreasing value of V_out [2–3].

Figure 1:

(A) Dip sensor and (B) cross section of the sensor element: 1, bending plate; 2, mechano-electrical transducer (piezoresistive bridge); 3, cavity with hydrogel; 4, Si chip; 5, porous membrane; 6, substrate with opening; 7, interconnect; 8, cap; 9, analyte solution.

2.1 Dip sensor

Custom-made circuit boards (Beta Layout, Aarbergen, Germany) were used as substrates for dip sensors (see Figure 1). The pressure sensor die with hydrogel and Al₂O₃ membrane was positioned above an opening in the circuit board and fixed with glue. Bond wires connected electrically the sensor chip with the circuit board. The soldered wires ensure a connection to the read-out device. A silicone cap protects the electrical components from aqueous solutions. Additionally, silicone rubber Scrintec^® 901 (Carl Roth, Karlsruhe, Germany) was coated around the pressure sensor die and the silicone cap in order to seal the dip sensor (not shown in Figure 1).

Motivated by a demand for inexpensive, robust, miniaturizable and reliable biochemical sensors with high signal reproducibility and long-term-stable sensitivity, especially for medical applications, which can be used also inside the human body, the concept of the hydrogel-based piezoresistive dip sensor has been converted into an implantable and biocompatible sensor set-up (see Figure 2).

Figure 2:

Test arrangement of an implantable flexible sensor: 1, plug with electrical connections; 2, Kapton foil with gold conductor lines and opening above Si chip; 3, piezoresistive Si chip with hydrogel and Al₂O₃ membrane.

2.2 Implantable flexible sensor

The test arrangement of an implantable flexible sensor is shown in Figure 2. Biocompatible polyimide foils (Kapton HN, coloprint tech-films, Frankenthal, Germany) were used as substrates. The vapour deposited individually designed gold conductor lines at the Kapton foil electrically connect the flip chip bonded pressure sensor chip to the read-out device by means of a plug. An encapsulation with medical grade silicone (MED-4211, NuSil Technology Europe, Mougins, France) protects the electrical components from fluids and ensures the biocompatibility of the sensor set-up. An area of approx. 3 × 3 mm² on the top of the porous membrane was left uncoated in order to allow the diffusion of solution and analytes into the chip cavity.

The implantable piezoresistive sensor set-up is useful for the medical diagnostics and for monitoring of sugar level, pH or other physiological blood parameters in the human body.

3 Hydrogel material preparation

The following monomers were purchased from Sigma Aldrich and used as received: Hydroxypropyl methacrylate HPMA, 2-(dimethylamino)ethyl methacrylate DMAEMA and tetraethylene glycol dimethacrylate TEGDMA. Furthermore, ethylene glycol EG, 2-hydroxy-4′-(2-hydroxyethoxy)-2-methylpropio-phenone, ß-D-glucose oxidase, 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide EDC and N-hydroxysuccinimide NHS from Sigma Aldrich were used. The monomer DMAEMA contains pH-sensitive tertiary amines. HPMA monomers serve the adjustment of the volume phase transition of the hydrogel to the physiological pH range [2]. TEGDMA acts as a crosslinker, whereas EG is the solvent. Monomer mixtures from HPMA, DMAEMA, TEGDMA and EG were produced in molar ratios of 60/40/02/20 and stirred for 30 min under a continuous nitrogen purge. After adding the photo initiator 2-hydroxy-4′-(2-hydroxyethoxy)-2-methylpropiophenone and further stirring with nitrogen purge under light exclusion, the pregel solution was cast in a 250 μm deep PTFE mold. The mold was covered with a photo mask with square structures (2 × 2 mm²) between two glass plates. The ultraviolet (UV) photo polymerization occurred for 2.5 min.

Before putting the UV-polymerized pH-sensitive hydrogel squares into the sensor chip cavity, an initial gel conditioning procedure was performed. An overnight washing in ethanol preceded 5-7 swelling/deswelling cycles in pH4/pH8 phosphate-buffered saline solutions PBS.

In order to prepare the glucose-sensitive hydrogels according to [4], EDC/NHS was used to conjugate the carboxyl groups of the enzyme glucose oxidase to the amine groups of the polymerized HPMA/DMAEMA/TEGDMA/EG (60/40/02/20) hydrogel squares. The conditioned pH-sensitive hydrogel discs were washed for 24 h in PBS pH3. Then, the discs were put into the conjunction solution with EDC/NHS concentrations of 200 mM/120 mM in PBS pH3 for 24 h. Afterwards, the hydrogel squares were incubated in PBS pH3 solution with 1000 u/ml glucose oxidase for 24 h. Subsequently, the hydrogel discs were washed with 0.1 M sodium phosphate for 2 h, with 2 M NaCl for 12 h, placed in NaCl overnight and stored in PBS pH7.4 for 24 h.

4 Results and discussion

The pH-sensitive HPMA/DMAEMA/TEGDMA hydrogel used in the pH dip sensor was prepared as described in the Section 3. The sensing mechanism is based on the protonation of tertiary amines on the DMAEMA backbone. In solutions with acidic pH values lower than pH8, an elevated backbone protonation temporarily increases the osmotic swelling pressure within the hydrogel, which subsequently leads to the gel swelling [3]. Figure 3 demonstrates a reversible gel swelling/shrinking at cycling between pH4 and pH8 monitored by the output voltage V_out of the dip sensor.

Figure 3:

Output voltage V_out of a dip sensor with UV-polymerized HPMA/DMAEMA/TEGDMA/EG (60/40/02/20) hydrogel at alternating pH values in PBS solution (pH8 and pH4).

A pH-sensitive HPMA/DMAEMA/TEGDMA hydrogel with conjugated glucose oxidase used in the dip sensor for glucose was prepared as described in the Section 3. The gel sensitivity with regard to glucose was investigated in vitro within a glucose concentration range from 0 mM to 100 mM in PBS solutions with the physiological pH value of 7.4. The swelling/deswelling process of the glucose-sensitive hydrogel due to the change of glucose concentration c_Glucose in PBS was monitored by the corresponding change of the output voltage V_out of the dip sensor (see Figure 4). The DMAEMA- and glucose oxidase containing hydrogel showed a volume increase with increasing glucose concentration. This increase occurred due to an enzymatic reaction causing an acidification inside the hydrogel. Hence, the tertiary amines on the DMAEMA backbone were protonated. Thereby, the hydrogel network swelled. Because of the smaller pH changes due to the enzyme reaction compared to the environmental pH changes depicted in Figure 3, a volume change of the glucose-sensitive hydrogel was accordingly lower. Consequently, a reduced change of V_out was observed. Nevertheless, Figure 4 demonstrates rapid V_out changes with alternating c_Glucose changes between 0 mM and 100 mM. Also, smaller glucose concentration changes from 20 mM to 0 mM could be detected.

Figure 4:

Output voltage V_out of a dip sensor with glucose oxidase conjugated HPMA/DMAEMA/TEGDMA/EG (60/40/02/20) hydrogel in PBS pH7.4 solutions with different glucose concentrations c_Glucoe (0 mM, 20 mM and 100 mM).

In the novel developed implantable sensor set-up onto flexible polyimide foil (see Figure 2), a pH-sensitive HPMA/DMAEMA/TEGDMA hydrogel was used. Figure 5 shows the pH depending increase/decrease of the V_out–value of the implantable flexible sensor at cycling between pH4 and pH8.

Figure 5:

Output voltage V_out of an implantable flexible sensor (see Figure 2) with HPMA/DMAEMA/TEGDMA/EG (60/40/02/20) hydrogel during cycling between pH8 and pH4 in PBS solutions.

By using the encapsulation with medical grade silicone, the pressure sensor chip and the gold conducting paths were protected from the surrounding fluids, salts, ions and other molecules. The life time of the piezoresistive biochemical sensors studied here under harsh physiological conditions amounted up to several weeks. Besides, the elasticity of the silicone encapsulation enables the sufficient bending plate displacements due to the hydrogel volume change leading to the changes of the output voltage, which have been measured also at small changes of the analyte concentration. It was found, that the implantable flexible sensor showed a lower V_out value at pH4 than the dip sensor (see Figures 3 and 5). The silicone seems to restrict the deflection of the bending plate lowering the maximal value of the output voltage. On the other hand, the silicone encapsulation seems to generate an initial prestress in the bending plate resulting in a higher V_out at pH8, when the gel is deswollen. Moreover, the response time t50% from the deswollen state in PBS pH8 to 50% swollen state in PBS pH4 was approx. 5 h for the dip sensor (see Figure 3) and ca. 3 h for the implantable flexible sensor (see Figure 5). This means, that the silicon coating leads to a shortened response time of the sensor.

5 Summary

UV-polymerized, conditioned, pH- and glucose-sensitive HPMA/DMAEMA/TEGDA/EG (60/40/02/20) hydrogel squares shaped like the chip cavity were incorporated in piezoresistive pressure sensor set-ups. Motivated by a demand for inexpensive, robust, miniaturizable and reliable biochemical sensors with high signal reproducibility and long-term-stable sensitivity, especially for implantable medical applications, the pH-sensitive hydrogel-based piezoresistive dip sensor concept was converted into an implantable sensor set-up based on biocompatible polyimide foil (see Figure 2). Making the sensor biocompatible and protecting electrical components from fluid, the medical grade silicone encapsulation of the implantable flexible sensor led to a shortened sensor response time.

The response of a glucose dip sensor with DMAEMA-based hydrogel conjugated with glucose oxidase using EDC/NHS was investigated in vitro in simulated physiological solutions with various glucose concentrations.

In the future, the response time of the sensors should be improved by using thinner hydrogel discs or hydrogel particles to accelerate the solution diffusion.

Author’s Statement

Research funding: The authors thank the Deutsche Forschungsgemeinschaft (DFG) for the financial support (Research Training Group 1865). Conflict of interest: Authors state no conflict of interest. Material and methods: Informed consent: Informed consent has been obtained from all individuals included in this study. Ethical approval: The research related to human use complies with all the relevant national regulations, institutional policies and was performed in accordance with the tenets of the Helsinki Declaration, and has been approved by the authors’ institutional review board or equivalent committee.

References

[1] Guenther M, Gerlach G, Wallmersperger T. Piezoresistive biochemical sensors based on hydrogels. Microsyst Technol. 2010;16:703–15.10.1007/s00542-009-0978-zSearch in Google Scholar

[2] Jung D-Y, Magda JJ, Han IS. Catalase effects on glucose-sensitive hydrogels. Macromolecules. 2000;33:3332–6.10.1021/ma992098bSearch in Google Scholar

[3] Guenther M, Gerlach G, Wallmersperger T, Avula M, Cho SH, Xie X, et al. Smart hydrogel-based biochemical microsensor array for medical diagnostics. Adv. Sci. Technol. 2013;85:47–52.10.4028/www.scientific.net/AST.85.47Search in Google Scholar

[4] Ghanem A, Ghaly A. Immobilization of glucose oxidase in chitosan gel beads. J Appl Polym Sci. 2004;91:861–6.10.1002/app.13221Search in Google Scholar

Published Online: 2016-9-30

Published in Print: 2016-9-1

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Articles in the same Issue

https://doi.org/10.1515/cdbme-2016-0029

Keywords for this article

biochemical sensor; glucose-sensitive; hydrogel; medical applications; pH-sensitive; piezoresistive

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