Development of a new type of biosensors based on ATR‐FTIR Spectroscopy

(1)

Université Libre de Bruxelles Service de Structure et Fonction des Faculté des Sciences Membranes Biologiques

Development of a new type of biosensors based on ATR‐FTIR

Spectroscopy

Thesis submitted as part of the requirements for the degree of Doctor in Science

September 2012

Andréa Goldsztein

Thesis supervisor : Erik Goormaghtigh

(2)

(3)

The multidisciplinary research described in this Thesis, combines technology, innovation, chemistry, biology and medicine, reflecting quite well my rather unusual academic curriculum, and in some way probably my personality as well.

(4)

(5)

Acknowledgements

My deepest thanks go to:

- Erik Goormaghtigh and Jean‐Marie Ruysschaert for welcoming me in the laboratory and giving me the opportunity to work with the SFMB team. Thank you both for having taken turns in guiding my research and for trusting my abilities despite my rather unusual academic curriculum. Special thanks also to Fabrice Homblé, for always being there to help me with many practical problems I have too often encountered.

- I am deeply indebted to Erik Goormaghtigh for his wise and supportive mentorship all through my Thesis work. I am particularly grateful for your patience and the understanding you showed along the moments during which I was sidetracked, focusing on my future family. You always showed confidence in my work, without ever pressuring me, giving me the freedom I needed and bolstering my self confidence. I am infinitely grateful to you for this.

- Everyone in the BIA‐ATR and MED‐ATR consortium, for these interesting and superbly organized meetings, which provided me with the opportunity to transcend the everyday research routine and see the bigger picture. I immensely appreciated the team work, learning more about organic chemistry, medical applications and needs, industrial issues, patents law, but above all, it gave me the opportunity to learn about how to manage a large project.

- Everybody in the SFMB laboratory who directly or indirectly helped me during these years of laboratory work

- My family, my parents Marcel and Shoshana and my sister Sasha, who never put any pressure on my shoulders and who supported me throughout this work. More specifically, thanks to my mother who helped me more « concretely », giving me this last little push I needed to finalize my Thesis.

- Tom, who rightly put a little pressure, when it was necessary, to motivate me and help me focus on the goal. Thank you for being here with me, thank you for being who you are.

- My big little girls, for all this happiness, love and joy that I can never get enough of.

(6)

(7)

Summary

Biosensors are analytical devices used for molecular recognition. They convert a biochemical signal into an quantifiable electrical signal. The principle is based on the recognition of one or more molecules of interest in solution (the ligand) by a biological component (the receptor) closely linked to a transductor substrate. The sensor responds to receptor‐ligand interactions and produces a measurable signal generally proportional to the concentration of bound ligands. Biosensors are already used in many different fields, especially in the medical domain (diagnostic and therapeutic), for environmental control and analysis, and in monitoring biotechnological processes.

The research described in this Thesis concerns the development of a new type of versatile biosensors. These sensors use an optical transduction element whose surface has been functionalized to allow the selective detection of receptor‐ligand interactions as well as the dosage of molecules bound to the receptor. The technique used for detection and quantification is ATR‐FTIR spectroscopy (Fourier Transform Infrared Spectroscopy in Attenuated Total Reflection mode). The system allows the direct online detection and without labeling of the target molecules. ATR‐FTIR spectroscopy enables the analysis of ligand molecules on the basis of their infrared (IR) spectral fingerprint feature, which offers a wealth of information. The technology is sensitive to the conformation of biological macromolecules and may, for example, be used to determine the secondary structure and orientation of proteins. Attenuated total reflection (ATR) is based on the presence of an evanescent wave that interacts with the sample placed in contact with the surface of the sensor. Its characteristics allow ligands to be detected only when they are bound to receptors closely linked to the sensor surface. This new biosensor system, called BIA‐ATR (for Biospecific Interaction Analysis ‐ Attenuated Total Reflection) is original and presents major advantages over most of commercially available biosensors. It offers the user the entire IR spectrum of the studied molecule, allowing not only to quantify the latter, but also offering the possibility of identifying its intrinsic nature. Another advantage is that it can be used to detect small peptides binding and in some cases associated chemical reactions with high sensitivity. The potential of this new sensor technology is evaluated in this work by its application to several systems of biological and medical interest. Those comprise the detection of the blood coagulation factor VIII involved in hemophilia type A, monitoring the phosphorylation of the gastric ATPase, a large receptor protein responsible for acid secretion in the stomach, and the quantification of an antibiotic, vancomycin, used in intensive care hospital settings in cases of severe bacterial infections with Staphylococcus aureus.

(8)

(9)

Abbreviations

Å: Angstrom (Å = 10^‐10m) Ac: acetyl

ADP: adenosine diphosphate Ala: alanine

ATP : adenosine triphosphate ATR: attenuated total relfection a.u.: absorbance unit

BIA: biospecific interaction analysis C: Celsius

CS: covalent sensor

DNA: desoxyribo nucleic acid D‐Ala: D‐Alanine

DTGS: deuterated tryglycine sulfate Eq: equation

FVIII: coagulation factor VIII FTIR: Fourier transform infrared Ge: germanium

GSH: generic hydophobic sensor H⁺: hydropgen ion

HEPES : 4‐(2‐hydroxyethyl)‐1‐piperazineethanesulfonic acid HSA: human serum albumine

IR: infrared

IRE: internal reflection element IU: interntional unit

K: potassium

KCl: potassium chloride kDa: kilo Dalton

Kd: dissociation constant Lys: lysine

MCT: mercury cadmium telluride

MED‐ATR: medicine atenuated total reflection

MIC: minimal inhibitory concentration: the lowest concentration of an antimicrobial that will inhibit the visible growth of a microorganism after overnight incubation

Mg: magnesium

MilliQ: deionisated water min: minute

mM: millimolar (10^‐3Molar)

(14)

mw: molecular weight NHS: N‐hydroxysuccinimide nm: nano meter (10^‐9 meter) OTS: octadecyltricholrosilane PBS: Phosphate‐buffered saline PC: phosphatidylcholine PDB: protein data bank

PE: phosphatidylethanolamine PEG: polyethylenglycol

pH: ‐log [H⁺]: measure the hydrogen ion concentration PS: phosphatidylserine

RI: refractive index SA: straptavidin

SAM: self assembled monolayer SD: standard deviation

SDS: sodium dodecyl sulfate

SDS‐PAGE : Sodium Dodecyl Sulfate – Polyacrylamide Electrophoresis Si: silicon

SPR: surface plasmon resonance µg: microgramme (10^‐6g)

µl: microlitre (10^‐6l) µM: micromolar (10^‐6M) UV: ultra violet

W: Watt

(15)

1

GENERAL INTRODUCTION

(16)

2

(17)

3 1) Introduction to Biosensors

Biosensors are analytical composite devices used for molecular recognition detection. They consist in an immobilized biological material in intimate contact with a suitable transducer that converts a biochemical signal into a quantifiable electrical signal (1). Their principle is based on the recognition of a biological analyte of interest free in solution (the ligand), by another biocomponent (the receptor) closely linked to the transducer sensor substrate (2, 3). The sensor reacts to the receptor‐ligand interactions and produces a measurable signal usually proportional to the concentration of the bound ligand.

Different transduction mechanisms can be applied to detect these molecular interactions based on the nature of the sensor support. Among those are electrochemical, which include potentiometric (4), amperometric (5)), calorimetric (6) and piezoelectric devices (7‐10), and surface acoustic wave technology (11). The most commonly reported class of biosensors uses optical systems (12‐l4, see below). The biological component in the sensors may be enzymes (15, 16), receptors (17, 18), whole cells (19‐21), organelles (22, 23), tissues (24), antibodies (25, 26), nucleic acids (27, 28) etc. In addition, various immobilization process can be used, such as entrapment or encapsulation, covalent binding, cross‐linking and adsorption (29‐30). The choice of a particular process generally depends on the nature of the biological element, the type of transducer used, the physicochemical properties of the component and the operating conditions in which the biosensor has to function.

The concept of biosensors was pioneered by Clark Jr and Lyons who proposed that enzymes could be immobilized at electrical detectors to form enzyme electrodes (31). However general interest in biosensors grew considerably since the work of Updike and Kicks (32) who described the first functional enzyme electrode based on glucose oxidase deposited on an oxygen sensor. This work marked the beginning of a major research effort into biotechnological and environmental applications, and offered the first commercial opportunities in biomedical science and healthcare with the introduction on the market, at the end of the 70s, of the glucose detector (glucometer).

In the early 80s, the growing interest in interactions between key macromolecules, such as proteins or DNA, and other molecules generated a new wave of inventions with a wide range of applications. Monitoring such interactions became essential for pharmaceutical, biomedical, agriculture and food industries (33, 34, 35). Biosensors were subsequently developed to detect the presence of pesticides, toxins, antibiotics, vitamins, or for ligands of various receptors for drug discovery (36‐40). Presently, biosensors are mostly used for measuring affinities between biomolecules, kinetics properties, as well as for molecular diagnostic, and quality control (41‐46).

(18)

4 1.1. Optical Biosensors

Optical biosensor devices consist of a biological sensing element integrated or connected to an optical transducer system. Their basic objective is to produce an electronic signal which is proportional in magnitude or frequency to the concentration of a specific analyte or group of analytes to which the biosensing element binds (47). Most of the optical biosensors exploit the evanescent wave phenomenon to characterize interactions between receptors that are attached to the biosensor surface and the ligands that are in solution above the surface (48‐51). The reflection of light at the interface of two media with different wave motion properties, forms an evanescent wave, whose intensity exponentially decays with the distance from the interface (52). Molecules interacting within the evanescent field contribute to the detected signal and can hence be sensed. As the maximum penetration depth of the evanescent field into the surrounding medium is of the order of 100 nm, it allows detection of molecules that are bound to immobilized recognition elements (53).

Optical biosensors are powerful alternatives to conventional analytical techniques given their particularly high specificity, sensitivity, small size and cost effectiveness (54, 55). Their accuracy and fast response time make them also very attractive compare to the traditional in‐vitro sensing techniques like ELISA, High Performance Liquid Chromatography (HPLC) or Fast Performance Liquid Chromatography (FPLC) (54, 56). Research and technological developments have focused on optical biosensors during the last decade because these sensors have a great potential for direct, real‐time and label‐free detection of many chemical and biological substances (8). Label‐free detection has the advantage of leaving the target molecule unaltered allowing it to be detected in its natural form.

There exist different optical detection methods including optical absorption detection and Raman spectroscopic detection (33, 57, 58). However, many of the best‐known optical biosensors are based on RI (refractive index) detection, and use SPR (Surface Plasmon Resonance) technology (59). In SPR the binding of molecules in solution to surface‐

immobilized receptors alters the refractive index of the medium near the sensor surface.

The SPR biosensor from Biacore is the most widely available biosensors of this type. A typical experimental set‐up for this sensor is shown in Fig. 1. In this set‐up the change in the refractive index is proportional to the mass loading on the sensor surface and can be monitored in real time to measure accurately the amount of bound analyte, its affinity for the receptor and the association and dissociation kinetics of the interaction (49, 60‐64).

(19)

5

Fig. 1: Typical set‐up for an SPR biosensor. Surface plasmon resonance (SPR) detects changes in the refractive index in the immediate vicinity of the surface layer of a sensor chip. SPR is observed as a sharp shadow in the reflected light from the surface at an angle that is dependent on the mass of material at the surface. The SPR angle shifts (from I to II in the lower left‐hand diagram) when biomolecules bind to the surface and change the mass of the surface layer. This change in resonant angle can be monitored non‐invasively in real time as a plot of resonance signal (proportional to mass change) versus time. Shematic drawing borrowed from http://www.path.ox.ac.uk/Facilities/sprfolder/Principles

In optical biosensors, the interface between the sensor surface and the chemical or biological systems to be studied is a key component. Receptors must be attached to some form of solid support, while retaining their native conformation and binding activity. This attachment must be stable over the course of a binding assay, and in addition, sufficient binding sites should be presented to the solution phase to interact with the ligand. Most crucially, the support should be resistant to non‐specific binding of the sample, which can mask the specific binding signal. To achieve these ends, many coupling strategies use a chemical linker layer between the sensor base and the biological component. Biacore uses, for example, a thin gold film evaporated on a glass sensor substrate and the biological receptors are immobilized through interactions with an amorphous dextran matrix (61) or via an alkanethiol‐based self assembled monolayer (SAM) (65‐67). However, SPR detection techniques yield only a mass‐response but do not provide chemical information about the ligand. Alternative techniques are therefore needed to establish the nature of the molecules that specifically bind to the receptor.

The present thesis addresses this gap. It describes the development of a new type of optical biosensors based on Attenuated Total Reflection Fourier Transform Infrared (ATR‐FTIR) spectroscopy. ATR‐FTIR spectroscopy allows the detection of molecules on the basis of their

(20)

6

characteristic IR spectral fingerprints, which offers a wealth of information for identifying them. The technology is sensitive to the conformation of biological macromolecules (68‐70) and is particularly well adapted to the characterization of organic molecules and biological systems (71, 72). It may, for example, be used to determine the protein secondary structure and orientation of proteins (70, 73).

2) BIA‐ATR sensors

The BIA (Biospecific Interaction Analysis) – ATR (Attenuated Total Reflection) sensors are the new designed biosensor devices developed in this work. They are based on ATR‐FTIR spectroscopy and are specifically geared to the investigation of ligand‐receptor interactions since the ATR‐FTIR technique allows gathering direct physico‐chemical information on the molecular interactions involved from the IR spectrum (74, 75).

The implementation of a new biosensor requires a multidisciplinary approach. Such approach involves ensuring the presence of biomaterial (immobilization of molecules on the sensor substrate), stabilizing the receptors, developing the actual sensor device (sensitivity, reproducibility issues) and integrating the fluidic and electronic components of the system.

The development of the ATR‐FTIR‐based biosensors described in this Thesis, was therefore carried out as a collaboration of four research groups with complementary expertise: the WOW engineering company (Namur, Belgium), providing engineering guidance and devices, the Catholic University of Louvain (UCL), responsible for the synthesis of the organic molecules (hydrophilic, hydrophobic) and for sensor surface grafting, the University of Mons‐Hainaut (UMH), focusing on the analyses of the molecular interfaces, and our laboratory at the Free University of Brussels (ULB) specializing in applying ATR‐FTIR spectroscopy to the investigation of biomolecules.

2.1. Principle of the BIAATR technology

BIA‐ATR sensors consist of an ATR element, usually a germanium or silicon crystal, which is transparent in the infrared spectral domain, and whose surface has been functionalized by covalently linking it to receptor molecules. Unlike in SPR, these molecules are grafted by wet chemistry directly on the sensor substrate at a distance determined by a spacer molecule and without an intermediary metal layer. A solution containing the potential ligand is flown over the crystal and detection of the binding event is achieved by recording ATR‐FTIR spectra.

The free ligand remains essentially undetected up to mM concentrations, but upon binding to the receptor, its concentration close to the sensor surface increases dramatically, rendering it readily detectable in the IR spectrum. The BIA‐ATR technology thus enables to

(21)

7

covalently link proteins to the sensor surface and investigate their response to modifications in their environment (76). A schematic representation of the experimental device is shown on Fig. 2.

Fig. 2: Schematic representation of the experimental device used for BIA‐ATR biosensors (A1: internal total reflection element transparent in the infrared spectral domain, A2: incident infrared beam, A3: ligand‐receptor interactions at the crystal surface). Inset 1: molecular recognition (B1: ATR element surface, B2:

functionalization layer, B3: free ligands and ligands bound to receptors). Inset 2: Schematic representation of the molecular construction (C1: anchoring molecule, C2: spacer molecule, C3: receptor). Shematic drawing borrowed from ref 76.

Thus, BIA‐ATR sensors use ATR‐FTIR spectroscopy as mechanism of signal transduction.

Fourier transform infrared spectroscopy involves sending electromagnetic radiation of mid‐

IR spectral region (4000 to 400 cm^‐1), whose wavelengths excite specific vibrational modes in organic molecules (proteins, lipids, etc…) bound or absorbed on the surface of an internal reflection element (IRE). In the present case, the IRE consists of a germanium or silicon crystal. Biological molecules have characteristic vibration modes, hence the radiation is absorbed at different frequencies depending on the molecule. Attenuated total reflection mode (ATR) is based on the presence of the evanescent wave produced at the reflection interface, which interacts with the sample placed in contact with the surface of the sensor chip (71, 77). As the amplitude of the evanescent wave decreases exponentially with the distance from the interface in the surrounding medium, only molecules close to the chip surface can significantly contribute to the infrared spectrum. Analytes present in water containing media can then be detected only when they are in close contact with the sensor surface.

A 1

B 1

A 2 A 3

L iq u id m o bile ph a s e : s o lv e n t a n d lig a n d s in s o lu tio n

I n se t 1

I n s e t 2 B 2

B 3

C 1 C 2 C 3

(22)

8

2.1.1. General principle of the ATR‐FTIR spectroscopy

It is essential to understand the basic principles that govern the absorption of the IR light at the reflecting interface of an internal reflection element (IRE). These principles have a profound impact on: (1) the spectrum intensity, (2) the band shape, (3) the intensity ratio between bands located at different wavelengths, (4) the ratio between the contributions of the bulk of the solvent and the sample, (5) the quantitative evaluation of surface concentrations, and (6) the impact of the polymer (or metallic layers) on the signal‐to‐noise ratio. The most common design for a attenuated total reflection set‐up is the trapezoidal plate (see section 2.3.4.). A schematic representation of an ATR set‐up is shown in Fig. 3.

The infrared beam is directed towards a high refractive index medium (germanium or silicon) transparent for the IR radiation of interest. Above a critical angle θc, which depends on the refractive index of the IRE, n1, and that of the external medium, n2,

θc = sin⁻¹n21 (1)

with n21 = n2/n1. The light beam is completely reflected when it impinges on the surface of the IRE. Several internal total reflections occur within the IRE until the beam reaches the end.

Fig. 3: Schematic representation of the internal reflection element (IRE) and the light path. The Cartesian components of the electric field are shown along the X, Y and Z axes. Two possible planes of polarization of the incident light are indicated by E// (polarization parallel to the incidence plane) and E⊥ (polarization perpendicular to the incidence plane). The incident beam makes an angle with respect to a normal to the IRE surface. The edges of the IRE are beveled so that the incident beam penetrates the IRE through a surface that is perpendicular to its propagation. Shematic drawing borrowed from ref 72.

(23)

9

It can be shown from Maxwell’s equations that superimposition of the incoming and reflected waves yields a standing wave within the IRE, established normal to the totally reflecting surface. Importantly, an electromagnetic disturbance also exists in the rarer medium beyond the reflecting interface. This so‐called evanescent wave is characterized by its amplitude, E(z), representing the time averaged field intensity at a distance z from the interface of the rarer medium, which decays exponentially with this distance:

E(z) = E0 e^−z/dp (2)

where E0 is the time averaged electric field intensity at the interface (z = 0) and dp is the penetration depth of the evanescent field, given by:

^/ (3)

where λ is the wavelength, λ1 = λ/n1, n21 = n2/n1 (78) and θ is the beam incidence angle of the IRE. Larger λ or smaller θ values correspond to a larger penetration depth. Fig. 4 illustrates a typical decay of the evanescent field intensity for total reflection in two different IREs. It is the presence of the evanescent field that makes possible the interaction between infrared light and the sample present on the surface of the IRE, within approximately the penetration depth of the field. An obvious conclusion that can be drawn is that the sample has to be in close contact with the IRE. Furthermore, the molecules from the bulk of the solvent are usually not sensed at all because they are too diluted and too far away from the reflecting interface. It is also apparent from Eq(3) that band intensity will depend on the wavelength since the penetration depth, and thereby the interaction with the sample, increases with λ. The principles described above yield spectral features, which are specific to ATR spectroscopy.

Fig. 4: Schematic representation of the evanescent wave on a 50 mm × 20 mm (A) KRS‐5 (Thalium Bromoiodide) and (B) germanium IRE, at 45° incidence, for a beam width of 3 mm. The grey density decreases as the function of the evanescent field intensity. For clarity, the evanescent field is represented only on the upper side of the IRE. Shematic drawing borrowed from ref 72.

(24)

10 2.1.2. Spectral specificity

Each biological molecule has its characteristic IR absorbance spectrum, which depends on its vibration modes. Indeed, chemical bonds vibrate at different energy levels and thus absorb at different wave numbers. For example in proteins, bands around 1650 cm⁻¹ (amide I) and 1544 cm⁻¹ (amide II) arise mainly from the absorption of the peptide groups. Amide I is the most intense absorption band of polypeptides. It is positioned between 1700 and 1600 cm⁻¹, but its exact frequency is determined by the geometry of the polypeptide chain and the hydrogen bonding patterns. ν(C=O) has a predominant role in amide I, accounting for 70–85% of the potential energy (79). Amide II occurs in the 1580–1510 cm⁻¹ region and derives mainly from the in‐plane N–H bending. Between 1800 and 1700 cm^‐1 one can find absorption corrsponding to C=O lipids vibration, allowing therefore to analyse the lipid/protein ratio in the 1800–1500 cm⁻¹ spectral region. Proteins adopting different structures and different secondary structure contents can be readily distinguished, as α helical regions absorb at a different wavenumber as β sheet regions (Fig. 5). This is of major importance when investigating conformational changes of peptides implicated in protein aggregation phenomena associated with diseases, such as Alzheimer, where the amyloid‐

beta peptide, responsible of pathogenesis, has been described as having different toxicities depending on the conformations it adopts (80).

Fig. 5: IR spectrum of a set of typical proteins with different secondary structure contents. The percentage of the two major secondary structure types (α‐helix and β‐sheet structure) is indicated in parentheses for each protein. Figure borrowed from ref 143.

Human albumin (α: 70 + β: 0)

Apolipoprotein E (α: 65 + β: 0) Dihydropteridine reductase (α: 37 + β 24)

Lysozyme (α: 31 + β: 6) Subtilisin (α: 31 + β: 18) Ovalbumin (α: 30 + β: 31)

Metallothionein (α: 0 + β: 0) Streptavidin (α: 0 + β: 48)

(25)

11 2.2. BIAATR versus SPR technology

Both ATR‐FTIR and SPR technologies use an optical element whose surface has been functionalized with a bound receptor. While SPR measures the variation in the resonance angle at the sensor‐substrate interface due to mass changes induced by ligand binding (81), ATR‐FTIR detects the difference in the IR reflected beam due to ligand attachment at the sensor surface. The SPR signal is recorded on a sensorgramme expressed in resonance unit (RU), which depends on the mass loaded on the sensor surface. The ATR‐FTIR data are represented as IR spectra characteristic of each type of biomolecule, whose intensity is related to the concentration of the bound ligand. A schematic comparison of both technologies is presented on Fig. 6.

Fig. 6: schematic comparison of the SPR and ATR‐FTIR technologies.

The SPR detection technique yields only a quantitative response from the sensor. In contrast, ATR‐FTIR, which can monitor the binding of a ligand to a receptor as other detection methods, gives in addition qualitative information about the chemical nature and structure of the ligand. Two or more different types of compounds that bind to the sensor can be detected simultaneously. Chemical reactions or conformational changes induced by the interaction can also be detected, thereby providing valuable information, e.g. on the mechanism of action of a drug or the activity of a protein (82, 83). Another interesting feature of the IR detection is that it allows the concentrations to be determined from the integrated molar extinction coefficient (84).

ATR-FTIR SPR

Functionalized surface

Metal layer Optical element

Glass (SPR) Si/Ge crystal (ATR‐FTIR)

light

detection Mobile phase solvents + ligands

(26)

12 2.3. Construction of the BIAATR sensors

One of the challenges encountered in the design of ATR‐FTIR‐ based sensors is the binding of receptor molecules of interest to the sensor substrate. This binding must maintain the integrity of the protein structure and activity, and the modified surface must prevent nonspecific binding. Several ways of surface modification have been investigated in previous work (72, 76). In the present Thesis, two types of BIA‐ATR sensors are developed: a hydrophobic sensor (GHS) for grafting lipids or membranes and a sensor called “covalent sensor” (CS) for attaching proteins and other receptor molecules. In both cases, the sensor substrate is a germanium or silicon crystal whose surface is first activated by oxidation to allow the attachment of hydrophobic compounds.

2.3.1. Crystal surface activation

The specificity of the sensor is related to the formation of specific ligand‐receptor associations. But this specificity can only be fully exploited by the sensor if the receptor (lipids, membranes, peptides, proteins…) can be covalently linked to the sensor substrate.

This covalent linkage crucially depends on the chemical nature of the ATR element. The most commonly used ATR elements are silicon (Si) and germanium (Ge) crystals. The chemical binding of molecules on Si substrates was studied in the past via the activation of a SiO2 layer (85‐87). This binding is generally improved by using an oxidation process to increase the density of Si‐OH groups on the surface. The modification of surface properties using such reactions is more complicated when considering Ge crystals because, unlike the SiO2, GeO2 is water soluble and the surface of germanium is less resistant to oxidation than silicon. However, Ge has a lager spectral window in the mid‐IR wavelengths suitable for studying biological molecules. Vogel and co‐workers eliminated the drawbacks for Ge surface activation (88, 89) by using self‐assembled monolayers (SAMs) of thiol‐terminated molecules on gold‐coated Ge elements to study ion complexes and proteins conformation with FTIR. However, they report adverse effects of the metal layer thickness (between 5 and 10 nm) deposited at the surface of an ATR element, on the infrared detection of biomolecules. The presence of both an anchoring oxide layer and a gold film significantly reduces the efficiency of the ATR‐FTIR method due to the attenuation of the evanescent wave in the thin metallic layer.

To avoid these problems, the BIA‐ATR technology developed another system for Ge crystal activation by successive immersions in HNO3 and H2O2/ethanedioic acid solutions (90). The activated Ge crystals can then be grafted with octadecyltrichlorosilane (OTS) molecules for lipids and membranes immobilization (hydrophobic sensor) or with PEG (polyethylenglycol) chains and specific spacer molecules to allow the covalent binding of receptors (covalent sensor).

(27)

13 2.3.2. The generic hydrophobic sensor (GHS)

In a sensor designed to work with lipids or membranes, the sensor surface is rendered hydrophobic by grafting a self‐assembled monolayer of octadecyltrichlorosilane (OTS) onto the Si or Ge surface (91, 92). The process involves the immersion of the activated ATR crystals in a solution containing OTS and hexadecane (see general Materials and Methods).

The corresponding chemical reactions are schematically represented on Fig. 7.

Fig. 7: Schematic representation of the chemical reactions used for OTS (octadecyltrichlorosilane) grafting onto the germanium surface.

The presence of the OTS layer is controlled by recording an ATR‐FTIR spectrum and its hydrophobicity is checked by the water droplet contact angle method (see general Materials and Methods). OTS‐grafted surfaces have a contact angle typically around 105° ± 5°. The amount of OTS bound on the crystal can be evaluated by comparison of the recorded FTIR spectra with the signal of DPPC (dipalmitoylphosphatidylcholine) monitored on Langmuir Blodgett monomolecular films (93, 94).

Once lipids or biological membranes with specific receptors are attached to the prepared OTS surface, the generic hydrophobic sensor is ready for use. Solutions containing the potential ligand can then be flown over the sensor and the ligand binding detection achieved by recording ATR‐FTIR.

2.3.3. The covalent sensor (CS)

Covalent grafting of the receptor is required in view of detecting ligand‐protein or ligand‐

biomolecule interactions. The BIA‐ATR technology uses an original organic coverage, without a metallic layer that allows the binding of a receptor and avoids nonspecific adsorptions. The preparation of the ATR crystal for the covalent anchoring of the biomolecule of interest requires several steps (90).

After activating the crystal surface, a homemade PEG (polyethyleneglycol) molecule (Fig. 8) is grafted by silanisation with a trietoxysilane ((EtO)3Si‐) foot anchor that reacts with the

()

¹⁵

OH

Ge Ge

O O

O

Ge

Activation OTS grafting

CH3 Si

Cl3‐Si ‐C18H37

(28)

14

oxidized surface, in order to reduce non specific absorption (95, 96, 97). The PEG used is a composite mixture containing an (EG)7 chains, and a small alkyl chain with a carbamate group. The alkyl chain and the carbamate are used to facilitate self packaging and protect the Si‐O bonds with the crystal from water.

O Si

O O

HN O

O

O O

Fig. 8: Silanisation reagent: home made compound: (EtO)3Si‐, C3H8, NH2COOH, (EG)7

The following step is the grafting of a spacer molecule (4 ‐ (p‐azidophenyl) butyrate, N‐

succinimidyl) by photoactivation (254 nm), which binds the specific PEG molecule on one extremity and any receptor ending with a primary amine on the other. The insertion of the molecule occurs randomly in the CH bonds of the PEG coverage. This relay molecule contains an ester/azide group capable of binding receptor molecules following photo activation. Biochemical reactions for the construction of the CS on a germanium crystal are schematically presented in Fig. 9.

Fig. 9: Schematic representation of the chemical reaction for the CS construction

O

(

)

⁷

(

^O

OH OH OH

Ge Ge Ge

Activation PEG grafting 2‐(methoxy (polylenoxy) propyl‐triethoxysilane)

PEG

(EtO)3Si OMe

N ₃

)

3 O O

O

Ge PEG Ge

O

HN

( )

3 O O

Spacer arms grafting Photoactivable ester/azide 4‐(p‐azidophenyl)butyrate de N‐

succinimidyl hυ

O O

O

O O

O

Si

Si Si

N

alkyl chain

carbamate PEG chain

triethoxysilane

(29)

15

Control of PEG and spacer molecule grafting is also achieved using the contact angle water drop technique (see general Materials and Methods): the acceptable contact angle for Ge‐

PEG surfaces is between 40° and 50°, and the required value for the Ge‐PEG‐NHS surface rises 60°‐70°. Structural analysis is performed by recording ATR‐FTIR spectra

After receptor grafting and stabilization, the covalent sensor is functional. Specific ligand can then be brought over the sensor and ligand‐receptor interactions detected by recording ATR‐FTIR.

2.3.4. Specific design of the ATR crystal

Two types of ATR elements are used for OTS grafting and binding of membrane phospholipid fragments: the trapezoidal standard ATR crystal, commercially available (50 x 20 mm²) and an original triangular prism (45 x 6.8 mm²) specifically designed as part of this work, and named “Toblerone crystal” in reference to the famous Toblerone chocolate bar.

While the first one is crossed by an IR beam with an angle of 45 ° which undergoes 25 internal reflections, the second has only one single internal reflection with an incidence angle of 45° (Fig. 10). Thus, the standard crystal has on its surface an evanescent field due to multiple reflections that allow the interaction of the IR beam with the sample along the entire length, whereas the triangular crystal has one reflection. However, the loss in the number of reflections is compensated by the improvement of the optics quality in the FTIR cell: the Toblerone is accommodated on a Golden Gate Micro‐ATR beam condenser (see general Materials and Methods). The latter is a more elaborate optical system (mirrors and lenses), which offers increased sensitivity.

Fig.10: Comparison of the standard and the newly designed Toblerone crystal

50 mm

20 mm

2 mm

45 mm

6.8mm

1 internal reflection 25 internal reflections

Standard crystal Toblerone crystal

(30)

16

Reduction of the sensor area provides a better control on the OTS surface functionalization and the use of smaller reagent volumes as well as smaller liquid flows in the experimental cell. The new geometry also allows to conduct several experiments on the same crystal with at least 11 lanes available.

Fig. 11: the 11 sensors channels of the Toblerone crystal.

The Toblerone crystals are built of germanium (Ge). The problem of Si, not encountered by Ge, is its opacity below 1400 cm^‐1. Yet the fingerprint region of organic molecules below 1400 cm^‐1 is essential for molecule identification: between 1230 and 1000 cm^‐1 one will find the absorption bands due to phosphate and sugar vibrations. To quantitatively detect these chemical compounds, the use of Ge crystal is highly suitable. Its grafting properties and hardness (wear resistance) make it also the material of choice.

2mm

(31)

17

Purpose of this work

This work describes the development of the BIA‐ATR technology and more particularly the control of the chemistry for grafting the surface of the optical element, as well as establishing the protocol for the quantitative detection of the molecules of interest. An important goal of the study is to demonstrate the applicability of the BIA‐ATR methodology to online detection of ligands binding to their specific receptors. Different applications of biological and medical interest are considered:

- In the first part, the potential of this new sensors technology is evaluated for detecting proteins in complex media, using as example the binding of the blood coagulation Factor VIII to phosphatidylserine membranes grafted onto the sensor crystal surface. The coagulation FVIII is involved in hemophilia A and the solutions used are prepared from commercial injectable samples containing excipients in which FVIII represents a very small fraction of the solutes.

- Next we demonstrate that the small size of a ligand relative to that of the receptor is not an obstacle for quantitative detection provided the ligand has a characteristic spectral fingerprint. This is illustrated by the work on the detection of a single phosphate group (phosphorylation) binding to the H⁺, K⁺ gastric ATPase, a large protein responsible for acid secretion in the stomach.

- Finally, the ability of BIA‐ATR sensors to detect and quantify an antibiotic is investigated using vancomycin‐peptide system. Detection of this antibiotic, used in hospital intensive cares in case of sever bacterial infections, is considered as a proof of concept for developing a sensor technology allowing ex‐vivo online monitoring of drug concentration in the serum of hospitalized patients.

These three investigations illustrate the versatility and advantages of the BIA‐ATR sensor technology, while also uncovering several aspects that require improvements and areas where further developments will be beneficial for establishing the BIA‐ATR technology as an attractive alternative to existing sensors.

(32)

18

(33)

19

GENERAL MATERIALS AND METHODS

(34)

20

(35)

21 1) ATR Spectrometer

Attenuated total reflection Fourier transformed infrared (ATR‐FTIR) spectra were obtained on a Bruker IFS55 FTIR spectrophotometer (Ettlingen, Germany) equipped with a MCT (Mercury Cadmium Telluride) detector (broad band 4000–800 cm^‐1) cooled with liquid N2 24 h hold time. The spectrometer was continuously purged with dry air (Whatman 75‐62, Haverhill, MA, USA). Measurements were recorded at room temperature.

The spectrophotometer was also equipped with a DTGS (Deuterated Triglycine Sulfate) detector, not cooled with liquid N2, which was used for some experiments in the context of the biosensor development for vancomycin detection.

ATR elements were installed on a beam condenser: a standard vertical ATR cell, customized with a lift system (WOW Company, Belgium) was used for trapezoïdal crystals (Fig. M1). The Toblerone crystals were accommodated on a Golden Gate Micro‐ATR beam condenser from Specac with a specific top plate containing a groove fitting the crystal (WOW Company) in replacement of the diamond‐bearing plate (Fig. M2)

Fig. M1: Standard vertical ATR cell, customized with a lift system (WOW Company, Belgium)

(36)

22

Fig. M2: Triangular‐shaped germanium crystal accommodated on a Golden Gate Micro‐ATR beam condenser from Specac with a specific top plate containing a groove fitting the crystal.

1.1. Spectrometer settings

When using the MCT detector, a stabilization time of 15 minutes after N2 cooling was required before use. Light transmission of the whole system was measured as the energy of the IR light reaching the detector. For optimal quality this energy should reach at least 4000 (arbitrary units); i.e. 15% of the energy measured in the absence of any accessory in the beam. In the case of the DTGS detector, the maximal energy obtained was of 2500 when using a Toblerone crystal.

The resulting absorbance spectrum comprises 1030 scans when recorded with the MCT detector compare to 175 scans when using the DTGS detector.

Resolution was of 8 cm^‐1and data were saved for wavelengths between 4000‐800 cm^‐1.

2) ATR elements

The internal reflection element (IRE) used for ATR‐FTIR infrared spectroscopy is a material transparent to infrared radiation and of high refractive index on which the receptor is grafted. Usually germanium is used, with a refractive index n = 4.0, alternatives are also zinc sulfide (n = 2.2), diamond (n = 2.35) and silicon (n = 3.4). The refractive index dictates the critical angle of incidence θc. The critical angle of the various materials in water (n = 1.5) is respectively of 22 °, 43 °, 40 ° and 26 °. The angle of incidence used on all devices is 45°.

(37)

23

In this work, germanium (Ge) was mainly used because of its higher refractive index and thus its lower critical angle of incidence. But other parameters have also been taken into account:

- The range of wavelength transparent in the mid‐IR: the useful window for Ge is broad enough to include the spectral region where receptor and ligand absorb (approximately 1800 to 1000 cm^‐1). It is from 4000 cm^‐1 to respectively 830, 950, 400 and 1400 cm^‐1 for Ge, ZnS, diamond and Si.

- Hardness: Ge hardness (about 780 kg/mm²) is the hardest material used for ATR. It is therefore particularly resistant to manipulation.

- The chemical reactivity: chemical reactivity of the Ge surfaces allows grafting receptor molecules.

Silicon was also suitable for chemical grafting, but its range of wavelengths was too limited.

It was however used in some experiments in the case of the standard trapezoidal crystal.

2.1. Standard trapezoidal crystals

The original commercially available ATR crystals were trapezoidal silicon (Si) or germanium (Ge) crystals (50 x 20 x 2 mm³). They were purchased from ACM, 78640 Villiers Saint Frédéric, France. They had an aperture angle of 45° yielding to 25 internal reflections (Fig.

M3).

Fig M3: Standard ATR germanium crystal

25 internal reflections 50mm

20mm

2mm

Development of a new type of biosensors based on ATR‐FTIR Spectroscopy