**2. Basis of LW sensors**

The Love wave physical effect was originally discovered by the mathematician Augustus Edward Hough Love. He observed an effect caused by earthquake waves far from the epi‐ center due to the lower acoustic wave velocity of waves propagating along the stratified geological layers [9]. The LW sensor is a layered structure formed, basically, by a piezoelec‐ tric substrate and a guiding layer (see Figure 1a). LW devices belong to the family of *surface acoustic wave* (SAW) devices in which the acoustic wave propagates along a single surface of the substrate. The piezoelectric substrate of a LW device primarily excites a *shear horizontal surface acoustic wave* (SH-SAW) or a *surface skimming bulk wave* (SSBW) depending on the ma‐ terial and excitation mode of the substrate. Both waves have shear horizontal particle dis‐ placements (perpendicular to the wave propagation direction and parallel to the waveguide surface). This type of acoustic wave operates efficiently in liquid media, since the radiation of compressional waves into the liquid is minimized.

**Figure 1.** a) Basic structure of a LW sensor. b) Five-layer model of a LW biosensor.

Traditionally, the most commonly used acoustic wave biosensors were based on QCM devi‐ ces. This was primarily due to the fact that the QCM has been studied in detail for over 50 years and has become a mature, commercially available, robust and affordable technology [3, 4]. LW acoustic sensors have attracted a great deal of attention in the scientific communi‐ ty during the last two decades, due to its reported high sensitivity in liquid media compared to traditional QCM-based sensors. Nevertheless, there are still some issues to be further un‐ derstood, clarified and/or improved about this technology; mostly for biosensor applica‐

LW devices are able to operate at higher frequencies than traditional QCMs [5]; typical oper‐ ation frequencies are between 80-300 MHz. Higher frequencies lead, in principle, to higher sensitivity because the acoustic wave penetration depth into the adjacent media is reduced [6]. However, the increase in the operation frequency also results in an increased noise level, thus restricting the LOD. The LOD determines the minimum surface mass that can be de‐ tected. In this sense, the optimization of the read out and characterization system for these

Another important aspect of LW technology is the optimization of the fluidics, specially the flow cell. This is of extreme importance for reducing the noise and increasing the biosensor

The analysis and interpretation of the results obtained with LW biosensors must be deeper understood, since the acoustic signal presents a mixed contribution of changes in the mass and the viscoelasticity of the adsorbed layers due to interactions of the biomolecules. A bet‐ ter understanding of the transduction mechanism in LW sensors is a first step to advance in

The fabrication process of the transducer, unlike in traditional QCM sensors, is another as‐ pect under investigation in LW technology, where features such as: substrate materials,

This chapter aims to provide an updated insight in the mentioned topics focused on biosen‐

The Love wave physical effect was originally discovered by the mathematician Augustus Edward Hough Love. He observed an effect caused by earthquake waves far from the epi‐ center due to the lower acoustic wave velocity of waves propagating along the stratified geological layers [9]. The LW sensor is a layered structure formed, basically, by a piezoelec‐ tric substrate and a guiding layer (see Figure 1a). LW devices belong to the family of *surface acoustic wave* (SAW) devices in which the acoustic wave propagates along a single surface of the substrate. The piezoelectric substrate of a LW device primarily excites a *shear horizontal surface acoustic wave* (SH-SAW) or a *surface skimming bulk wave* (SSBW) depending on the ma‐ terial and excitation mode of the substrate. Both waves have shear horizontal particle dis‐

this issue; however its inherent complexity leads, in many cases, to frustration [8].

high frequency devices is a key aspect for improving the LOD [7].

system stability; aspects that will contribute to improve the LOD.

sizes, structures and packaging must be still optimized.

tions.

278 State of the Art in Biosensors - General Aspects

sors applications.

**2. Basis of LW sensors**

LW sensors consist of a transducing area and a sensing area. The transducing area consists of the *interdigital transducers* (IDTs), which are metal electrodes, sandwiched between the piezoelectric substrate and the guiding layer. The input IDT is excited electrically (applying an rf signal) and launches a mechanical acoustic wave into the piezoelectric material which is guided through the guiding layer up to the output IDT, where it gets transformed back to a measurable electrical signal. The *sensing area* is the area of the sensor surface, located be‐ tween the input and output IDT, which is exposed to the analyte.

LW sensors can be used for the characterization of processes involving several layers de‐ posited over the sensing area; such is the case of biosensors. A LW biosensor can be descri‐ bed as a layered compound formed by the LW sensor in contact with a finite viscoelastic layer, the so-called coating, contacting a semi-infinite viscoelastic liquid as indicated in Fig‐ ure 1b. Each layer has its material properties given by: the shear modulus *µ*, density *ρ* and viscosity *η*. Hence, the subscripts S, L, SA, C and F denotes the substrate, guiding layer, sensing area, coating and fluid layers, respectively. Biochemical interactions at the sensing area cause changes in the properties of the propagating acoustic wave which can be detect‐ ed at the output IDT.

The difference between the mechanical properties of the guiding layer and the substrate cre‐ ates an entrapment of the acoustic energy in the guiding layer keeping the wave energy near the surface and slowing down the wave propagation velocity. This guiding layer phenom‐ enon makes LW devices very sensitive towards any changes occurring on the sensor surface, such as those related to mass loading, viscosity and conductivity [5]. The higher the confine‐ ment of the wave in the guiding layer, the higher the sensitivity [10].

The proper design of a LW device for biosensor applications must consider the advances made on these basic elements. Updated information about each one of these elements is then required and can be found in the following sections.

#### **2.1. Piezoelectric substrate**

Thanks to *piezoelectricity* electrical charges can be generated by the imposition of mechanical stress. The phenomenon is reciprocal; applying an appropriate electrical field to a piezoelec‐ tric material generates a mechanical stress [11]. In LW sensors an oscillating electric field (rf signal) is applied in the input IDT which, due to the piezoelectric properties of the substrate, launches an acoustic guided wave. The guided wave propagates through the guiding layer up to the output IDT where, again due to the piezoelectric properties of the substrate, is con‐ verted back to an electric field for measurement. A remarkable parameter of the piezoelec‐ tric substrate is the *electromechanical coupling coefficient* (K2 ), which indicates the conversion efficiency from electric energy to mechanical energy; its value depends on the material prop‐ erties. This is an important design parameter in LW sensors, since higher K2 lead to low loss LW devices and, therefore, more sensitive LW sensors [12].

When choosing a material for the substrate of LW devices, apart from the desired low losses, other requirements, such as low *temperature coefficient of frequency* (TCF) have to be consid‐ ered as well. Special crystal cuts of the piezoelectric substrate material1 can yield an intrinsi‐ cally temperature-compensated device which minimizes the influence of temperature on the sensor response, thus improving the LOD [13,14].

The shear horizontal polarization required for operation of the LW sensor in liquid media is another aspect to be considered when choosing the substrate material. In this sense, quartz is the only common substrate material that can be used to obtain a *purely* shear polarized wave [13]. The crystal cut and the wave propagation direction, which depends on the IDTs orientation, define the elastic, dielectric and piezoelectric constants of the crystal, and there‐ fore the wave polarization. Possible cuts which generate a purely shear polarized wave are the AT-cut quartz and the ST-cut quartz. AT-cut quartz and ST-cut quartz are both Y-cuts, rotated 35°15' and 42°45° about the original crystallographic X-axis, respectively.

Initially, LW devices were made in ST-cut quartz [15], however, ST-cut quartz is very sensi‐ tive to temperature (its TCF is around 40 ppm/°C) [16]. This was a restrictive factor in terms of sensor LOD and, thus, temperature-compensated systems based on different quartz cuts and different materials for the substrate such as lithium tantalate (LT), LiTaO3, and lithium niobate (LN), LiNbO3, were investigated [17-19]. In particular, AT-cut quartz, 36° YX LT and 36° YX LN were proposed, the last two corresponds to specific cuts of LT and LN materials [10]. Table 1 contains the values of some characteristic parameters of the previously men‐

<sup>1</sup> The substrate crystal cuts (or plates) are obtained by cutting slices of a single-crystal starting material with an arbitra‐ ry orientation relative to the three orthogonal crystallographic axes.

tioned substrate materials. In column 2, the *substrate shear velocity vS,* is defined by the sub‐ strate material properties (*vS* = (*µS/ρS*) 1/2).


**Table 1.** Most commonly employed crystal cuts for LW devices (modified from [18]).\*Approximate value.

LN substrates have higher coupling factor and low propagation loss than LT and quartz substrates. However, these substrates are extremely vulnerable to abrupt thermal shocks.

The low insertion loss, very large electromechanical coupling factor K2 and low propagation loss which characterize 36° YX LT substrates [20] provide advantages over other substrates such as quartz cuts, where exquisite care in the fluidic packaging is required to prevent ex‐ cessive wave damping [21]. For this reason, LT seems to be the substrate material of choice in high-loss applications due to its high coupling factor K2 , while in low-loss applications quartz may exhibit better wave characteristic [22]. The main shortcomings of 36° YX LT sub‐ strates are: they do not generate a pure shear wave, which increases the damping when is liquid loaded; and they have a poor thermal stability due to their high TFC (-30 to -40 ppm/°C [19]) if compared with AT-quartz.

From the point of view of thermal stability, AT-cut quartz seems to be the most appropriate due to its very low TCF [14]. Although the coupling coefficient of the AT-cut quartz is the lowest compared to other cuts, in our opinion, AT-cut quartz is currently the most suitable substrate for LW biosensing applications among the mentioned substrates, for several rea‐ sons: 1) it is capable of generating pure shear waves, diminishing the damping when is liq‐ uid loaded; 2) its thermal stability is the highest one, which improves the LOD; 3) the mass sensitivity of quartz substrates is significantly high than that of LT substrates [17,23]; and 4) LT and LN substrates are extremely fragile and must be handled with great care during the device fabrication to prevent them from breaking in pieces.

#### **2.2. Interdigital transducers**

such as those related to mass loading, viscosity and conductivity [5]. The higher the confine‐

The proper design of a LW device for biosensor applications must consider the advances made on these basic elements. Updated information about each one of these elements is then

Thanks to *piezoelectricity* electrical charges can be generated by the imposition of mechanical stress. The phenomenon is reciprocal; applying an appropriate electrical field to a piezoelec‐ tric material generates a mechanical stress [11]. In LW sensors an oscillating electric field (rf signal) is applied in the input IDT which, due to the piezoelectric properties of the substrate, launches an acoustic guided wave. The guided wave propagates through the guiding layer up to the output IDT where, again due to the piezoelectric properties of the substrate, is con‐ verted back to an electric field for measurement. A remarkable parameter of the piezoelec‐

efficiency from electric energy to mechanical energy; its value depends on the material prop‐ erties. This is an important design parameter in LW sensors, since higher K2 lead to low loss

When choosing a material for the substrate of LW devices, apart from the desired low losses, other requirements, such as low *temperature coefficient of frequency* (TCF) have to be consid‐ ered as well. Special crystal cuts of the piezoelectric substrate material1 can yield an intrinsi‐ cally temperature-compensated device which minimizes the influence of temperature on the

The shear horizontal polarization required for operation of the LW sensor in liquid media is another aspect to be considered when choosing the substrate material. In this sense, quartz is the only common substrate material that can be used to obtain a *purely* shear polarized wave [13]. The crystal cut and the wave propagation direction, which depends on the IDTs orientation, define the elastic, dielectric and piezoelectric constants of the crystal, and there‐ fore the wave polarization. Possible cuts which generate a purely shear polarized wave are the AT-cut quartz and the ST-cut quartz. AT-cut quartz and ST-cut quartz are both Y-cuts,

Initially, LW devices were made in ST-cut quartz [15], however, ST-cut quartz is very sensi‐ tive to temperature (its TCF is around 40 ppm/°C) [16]. This was a restrictive factor in terms of sensor LOD and, thus, temperature-compensated systems based on different quartz cuts and different materials for the substrate such as lithium tantalate (LT), LiTaO3, and lithium niobate (LN), LiNbO3, were investigated [17-19]. In particular, AT-cut quartz, 36° YX LT and 36° YX LN were proposed, the last two corresponds to specific cuts of LT and LN materials [10]. Table 1 contains the values of some characteristic parameters of the previously men‐

1 The substrate crystal cuts (or plates) are obtained by cutting slices of a single-crystal starting material with an arbitra‐

rotated 35°15' and 42°45° about the original crystallographic X-axis, respectively.

), which indicates the conversion

ment of the wave in the guiding layer, the higher the sensitivity [10].

required and can be found in the following sections.

tric substrate is the *electromechanical coupling coefficient* (K2

LW devices and, therefore, more sensitive LW sensors [12].

sensor response, thus improving the LOD [13,14].

ry orientation relative to the three orthogonal crystallographic axes.

**2.1. Piezoelectric substrate**

280 State of the Art in Biosensors - General Aspects

*Interdigital transducers* (IDTs) were firstly reported in 1965 by White and Voltmer [24] for generating SAWs in a piezoelectric substrate. An IDT, in its most simple version, is formed by two identical combs-like metal electrodes whose strips are located in a periodic alternat‐ ing pattern located on top of the piezoelectric substrate surface. Figure 2a shows the struc‐ ture of a *single-electrode* IDT which consists of two strips per *period p* and *acoustic aperture W*. The strip width is equal to the space between strips (*p*/4). One comb is connected to ground and the other to the center conductor of a coaxial cable where a rf signal is provided. A pair of two strips with different potential is called a *finger pair*.

The IDT electric equivalent circuit is explained in reference [25]. Figure 2b shows the IDT frequency response, where *A(f)* is the electrical amplitude of the rf signal. The maximum in *A(f)* occurs when the *wavelength λ* of the generated acoustic wave is equal to the period *p* and this arises at the so called *synchronous frequency fs*. In relation to the *bandwidth B* of an IDT frequency response, this will be narrower when increasing the *number of finger pairs N*. However, there is a limitation in the maximum *N* recommended, due to the fact that, in practice, when *N* exceeds 100, the losses associated with mass loading and the scattering from the electrodes increase. This neutralizes any additional advantage associated with the increase of the number of the finger pairs.

Due to symmetry of the IDT in the direction of propagation, the LW energy is emitted in equal amounts in opposite directions, giving an inherent loss of 3 dB. In a two-port device this factor contributes 6 dB to the total insertion loss [25,26].

Aluminum has been widely used as IDTs material and has been extensibility demonstrated in literature as suitable for SAW generation. Aluminum has an ability to resist corrosion and is the third most abundant element on Earth (after oxygen and silicon). It also has a low cost compared to other metals. The metallic layer of the electrodes must be thick enough to present a low electric resistance, but sufficiently thin to avoid an excessive mechanic charge for the acoustic wave (acoustic impedance breaking) [27]. Generally, a thickness between 100 and 200 nm of aluminum is employed.

There are a number of second-order effects, which are often significant in practice, that af‐ fect the transducer frequency response. The effect for which the transducer strips reflect sur‐ face waves causing mechanical and electrical perturbations of the surface is called *electrode interaction* [30]*.* Usually, these unwanted reflections cancel each other over wide frequency bands and become negligible. However, in a certain frequency band, the scattered waves are in phase, adding them constructively and causing very strong reflection (*Bragg reflection*) which distorts the transducer frequency response. For a *single-electrode* IDT (see Figure 2a), this situation occurs at the resonance condition *λ* = *p*. Thus, *double-electrode* (or *double finger pair* or *split-electrode*) IDTs are used to avoid this unwanted effect. In double-electrode IDTs there are four strips per period (see Figure 2c) and thus, the Bragg reflection can be sup‐ pressed at the LW resonance frequency [28]. One disadvantage of the double-electrode is the increased lithographic resolution required for fabricating the IDTs [29].

Another significant second-order effect is the generation of the *triple-transit signal*. In a de‐ vice using two IDTs, which is the case of a LW device, the output IDT will in general pro‐ duce a reflected wave, which is then reflected a second time by the input IDT. Thus, a reflected wave reaches the output IDT after traversing the substrate three times, giving an unwanted output signal known as the *triple-transit signal* [26]. This effect is reduced by mak‐ ing the input and output IDT separation large enough.

Some authors use *reflectors* to improve the frequency response of the LW device. Reflectors are composed of metal gratings placed in the same configuration than IDTs and are located

at the ends of the IDTs (see Figure 2d). These components have generally less finger pairs than the IDTs. The space periodicity of the reflectors is equal than in the IDTs [30]. Very nar‐ row low-loss pass band can be realized in a two-port device, when the device is designed so that the reflectors resonate at the IDT resonance frequency, since the transfer admittance be‐ comes very large [28]. 10 with great care during the device fabrication to prevent them from breaking in pieces. 11 **2.2. Interdigital transducers**  12 *Interdigital transducers* (IDTs) were firstly reported in 1965 by White and Voltmer [24] for 13 generating SAWs in a piezoelectric substrate. An IDT, in its most simple version, is formed by two

Running Title <sup>5</sup>

1 shear wave, which increases the damping when is liquid loaded; and they have a poor thermal

 From the point of view of thermal stability, AT-cut quartz seems to be the most appropriate due to its very low TCF [14]. Although the coupling coefficient of the AT-cut quartz is the lowest compared to other cuts, in our opinion, AT-cut quartz is currently the most suitable substrate for LW biosensing applications among the mentioned substrates, for several reasons: 1) it is capable of generating pure

2 stability due to their high TFC (-30 to -40 ppm/°C [19]) if compared with AT-quartz.

#### **2.3. Guiding layer** 14 identical combs-like metal electrodes whose strips are located in a periodic alternating pattern located 15 on top of the piezoelectric substrate surface. Figure 2a shows the structure of a *single-electrode* IDT

19 *finger pair*.

and the other to the center conductor of a coaxial cable where a rf signal is provided. A pair

The IDT electric equivalent circuit is explained in reference [25]. Figure 2b shows the IDT frequency response, where *A(f)* is the electrical amplitude of the rf signal. The maximum in *A(f)* occurs when the *wavelength λ* of the generated acoustic wave is equal to the period *p* and this arises at the so called *synchronous frequency fs*. In relation to the *bandwidth B* of an IDT frequency response, this will be narrower when increasing the *number of finger pairs N*. However, there is a limitation in the maximum *N* recommended, due to the fact that, in practice, when *N* exceeds 100, the losses associated with mass loading and the scattering from the electrodes increase. This neutralizes any additional advantage associated with the

Due to symmetry of the IDT in the direction of propagation, the LW energy is emitted in equal amounts in opposite directions, giving an inherent loss of 3 dB. In a two-port device

Aluminum has been widely used as IDTs material and has been extensibility demonstrated in literature as suitable for SAW generation. Aluminum has an ability to resist corrosion and is the third most abundant element on Earth (after oxygen and silicon). It also has a low cost compared to other metals. The metallic layer of the electrodes must be thick enough to present a low electric resistance, but sufficiently thin to avoid an excessive mechanic charge for the acoustic wave (acoustic impedance breaking) [27]. Generally, a thickness between

There are a number of second-order effects, which are often significant in practice, that af‐ fect the transducer frequency response. The effect for which the transducer strips reflect sur‐ face waves causing mechanical and electrical perturbations of the surface is called *electrode interaction* [30]*.* Usually, these unwanted reflections cancel each other over wide frequency bands and become negligible. However, in a certain frequency band, the scattered waves are in phase, adding them constructively and causing very strong reflection (*Bragg reflection*) which distorts the transducer frequency response. For a *single-electrode* IDT (see Figure 2a), this situation occurs at the resonance condition *λ* = *p*. Thus, *double-electrode* (or *double finger pair* or *split-electrode*) IDTs are used to avoid this unwanted effect. In double-electrode IDTs there are four strips per period (see Figure 2c) and thus, the Bragg reflection can be sup‐ pressed at the LW resonance frequency [28]. One disadvantage of the double-electrode is the

Another significant second-order effect is the generation of the *triple-transit signal*. In a de‐ vice using two IDTs, which is the case of a LW device, the output IDT will in general pro‐ duce a reflected wave, which is then reflected a second time by the input IDT. Thus, a reflected wave reaches the output IDT after traversing the substrate three times, giving an unwanted output signal known as the *triple-transit signal* [26]. This effect is reduced by mak‐

Some authors use *reflectors* to improve the frequency response of the LW device. Reflectors are composed of metal gratings placed in the same configuration than IDTs and are located

increased lithographic resolution required for fabricating the IDTs [29].

ing the input and output IDT separation large enough.

of two strips with different potential is called a *finger pair*.

this factor contributes 6 dB to the total insertion loss [25,26].

increase of the number of the finger pairs.

282 State of the Art in Biosensors - General Aspects

100 and 200 nm of aluminum is employed.

The difference between the mechanical properties of the piezoelectric substrate and the guiding layer generates a confinement of the acoustic energy in the guiding layer, slowing down the wave propagation velocity, but maintaining the propagation loss [32]. In particu‐ lar, the condition for the existence of Love wave modes is that the shear velocity of the guid‐ ing layer material (*vL* = (*µL/ρL*) 1/2) is less than that of the substrate (*vS* = (*µS/ρS*) 1/2) [31]. When both materials, substrate and guiding layer, have similar density the ratio *µS*/ *µL* determine the dispersion of the Love mode; a large value of that ratio (higher *µS* and lower *µL*) leads to a stronger entrapment of the acoustic energy [32] and thus, greater sensitivity. Hence, the benefit of the guiding layer is that an enhanced sensitivity to mass deposition can be ob‐ tained [33], but also to viscoelastic interactions. 16 which consists of two strips per *period p* and *acoustic aperture W*. The strip width is equal to the 17 space between strips (*p*/4). One comb is connected to ground and the other to the center conductor of 18 a coaxial cable where a rf signal is provided. A pair of two strips with different potential is called a 20 The IDT electric equivalent circuit is explained in reference [25]. Figure 2b shows the IDT frequency 21 response, where *A(f)* is the electrical amplitude of the rf signal. The maximum in *A(f)* occurs when the 22 *wavelength λ* of the generated acoustic wave is equal to the period *p* and this arises at the so called *synchronous frequency fs* 23 . In relation to the *bandwidth B* of an IDT frequency response, this will be

Figure 2. a) Single-electrode Interdigital Transducer (IDT) with period *p*, electrode width equal to space between electrodes (p/4) and aperture W (modified from [25]). b) Frequency response of an IDT (positive frequencies), where *A(f)* is the electrical amplitude (modified from [25]). c) Double-electrode IDT with **Figure 2.** a) Single-electrode Interdigital Transducer (IDT) with period *p*, electrode width equal to space between elec‐ trodes (p/4) and aperture W (modified from [25]). b) Frequency response of an IDT (positive frequencies), where *A(f)* is the electrical amplitude (modified from [25]). c) Double-electrode IDT with period *p*, electrode width equal to space between electrodes and aperture W. d) Two grating reflectors are place at both ends of the IDTs (modified from [30]).

period *p*, electrode width equal to space between electrodes and aperture W. d) Two grating reflectors are

place at both ends of the IDTs (modified from [30]).

The effect of the guiding layer on Love modes influence the substrate coupling factor K2 , in‐ creasing it [14]. Also influence the temperature behavior, since modifies the TCF compared to their parent SSBWs device.

In relation to the materials used for the guiding layer, those with a low shear velocity and low insertion loss seem to be the most promising materials for developing sensitive biosen‐ sors [22,32,34]. Materials such as polymers [35], silicon dioxide (SiO2) [17], gold (Au) [36] and zinc oxide (ZnO) [37,38] have been used as guiding layers [21]. In Table 2 some proper‐ ties of these materials are presented2 . The use of polymers (like Novolac, polyimide, polydi‐ methylsiloxane (PDMS) and polymethylmethacrylate (PMMA)) is interesting from the point of view of the sensitivity, since they have low shear velocity. Additionally, some pol‐ ymers, like Novolac photoresist, are very resistant to chemical agents [39,40]. However, polymers have high acoustic damping (losses) [39] and this is a clear disadvantage for bio‐ sensing application.


**Table 2.** Employed materials for guiding layers of LW devices.

Guiding layer/substrate structures made with ZnO as guiding layer have some advantages over those with a different material. This is the case of ZnO/ST-quartz structure, for which significantly high sensitivity, small TCF and high K2 were reported [38]. ZnO/LT devices were also found to have higher mass sensitivity than SiO2/LT [23]. However, ZnO has sever‐ al disadvantages: it is CMOS contaminant, a semiconductor, and thus, it can deteriorate the efficiency of the transducers and make some artifacts. In addition, it gets easily rough when sputtered and it is very reactive with acids, liquids, or moisture, so it will dissolve if ex‐ posed to water or humid environment, which is a big problem in biosensors application. Re‐ garding Au guiding layers, they provide very strong wave guiding, since Au has a relatively

<sup>2</sup> These values are for guidance, since for deposited or grown materials these values depend on the desposition tech‐ nique and for polymers layers on the cure process.

low shear acoustic velocity and a high density. However, it couples the rf signal from input to output IDT.

Silicon dioxide (SiO2) -also known as fused silica- is a standard material in semiconductor industry and offers low damping, sufficient low shear velocity and excellent chemical and mechanical resistance [41]. It is the only native oxide of a common semiconductor which is stable in water and at elevated temperatures, an excellent electrical insulator, a mask to com‐ mon diffusing species, and capable of forming a nearly perfect electrical interface with its substrate. When SiO2 is needed on materials other than silicon, it is obtained by chemical va‐ por deposition (CVD), either thermal CVD or Plasma enhanced CVD (PECVD) [42]. The main shortcoming for SiO2 is that the optimum thickness, at which the maximum sensitivity is reached, is very high (see Section 5), so this complicates the manufacturing process. Nev‐ ertheless, at the present, we consider that SiO2 is the most appropriate material for LW bio‐ sensors guiding layer, mainly due to its low damping and excellent chemical and mechanical properties [42].

#### **2.4. Sensing area**

The effect of the guiding layer on Love modes influence the substrate coupling factor K2

to their parent SSBWs device.

284 State of the Art in Biosensors - General Aspects

ties of these materials are presented2

**Table 2.** Employed materials for guiding layers of LW devices.

nique and for polymers layers on the cure process.

significantly high sensitivity, small TCF and high K2

sensing application.

creasing it [14]. Also influence the temperature behavior, since modifies the TCF compared

In relation to the materials used for the guiding layer, those with a low shear velocity and low insertion loss seem to be the most promising materials for developing sensitive biosen‐ sors [22,32,34]. Materials such as polymers [35], silicon dioxide (SiO2) [17], gold (Au) [36] and zinc oxide (ZnO) [37,38] have been used as guiding layers [21]. In Table 2 some proper‐

methylsiloxane (PDMS) and polymethylmethacrylate (PMMA)) is interesting from the point of view of the sensitivity, since they have low shear velocity. Additionally, some pol‐ ymers, like Novolac photoresist, are very resistant to chemical agents [39,40]. However, polymers have high acoustic damping (losses) [39] and this is a clear disadvantage for bio‐

**Guiding layer material μ***L (GPa)* **ρ***L (kg/m3) vL (m/s)*

SiO2 17.87 2200 2850.04

ZnO 40.17 5720 2650.00

Au 28.50 19300 1215.19

Polyimide 0.87 1420 780.48

PDMS 250×10-6 965 16.09

PMMA 1.70 1180 1200.28

Guiding layer/substrate structures made with ZnO as guiding layer have some advantages over those with a different material. This is the case of ZnO/ST-quartz structure, for which

were also found to have higher mass sensitivity than SiO2/LT [23]. However, ZnO has sever‐ al disadvantages: it is CMOS contaminant, a semiconductor, and thus, it can deteriorate the efficiency of the transducers and make some artifacts. In addition, it gets easily rough when sputtered and it is very reactive with acids, liquids, or moisture, so it will dissolve if ex‐ posed to water or humid environment, which is a big problem in biosensors application. Re‐ garding Au guiding layers, they provide very strong wave guiding, since Au has a relatively

2 These values are for guidance, since for deposited or grown materials these values depend on the desposition tech‐

. The use of polymers (like Novolac, polyimide, polydi‐

were reported [38]. ZnO/LT devices

, in‐

The sensing area can be made of different material than the guiding layer. Sensing layers have been reported composed of materials like PMMA [43] and SU-8 [44], but the most com‐ monly employed is gold (Au). Generally, the thickness of this layer varies from 50-100 nm and 2-10 nm of chrome (Cr) or titanium (Ti) is needed to promote adherence to the guiding layer. Au surfaces are very attractive candidates for self-assembly due to their metallic na‐ ture, great nobility, and particular affinity for sulphur. This aspect allows functionalization with thiols of various types and adhesion to diverse organic molecules, which are modified to contain a sulphur atom. These coatings, assembled onto the gold surfaces, can serve as biosensors [36]. Immobilization techniques on gold for biosensing are quite common and much utilized in the scientific community. However, immobilization techniques on different materials, like SiO2, could greatly simplify the LW biosensors fabrication.

#### **3. Measurement techniques**

Figure 3a shows a configuration of a two-port LW device where it behaves as a delay line. *D* is the distance between input and output IDT and *L* is the center-to-center dis‐ tance between the IDTs. Thus, the sensor is a transmission line which transmits a mechan‐ ical signal (acoustic wave) launched by the input port (input IDT) due to the applied rf electrical signal. After a time delay the traveling mechanical wave is converted back to an electric signal in the output port (output IDT). In general, changes in the coating layer and/or in the semi-infinite fluid medium (see Figure 1b) produce variations in the acous‐ tic wave properties (wave propagation velocity, amplitude or resonant frequency). These variations can be measured comparing the input and output electrical signal, since pha‐ sor *Vin* remains unchanged, while phasor *Vout* changes. Thus, from an electric point of view, a LW delay line can be defined by its transfer function *H(f)* = *Vout/Vin*, which repre‐ sents the relationship between input and output electrical signal. *H(f)* is a complex num‐ ber which can be defined as *H(f)* = *Aejφ*, being *A* = |*Vout/Vin|* the amplitude and *φ* the *phase shift* between *Vout* and *Vin*. In terms of voltage, the *insertion loss* (*IL*) in dB is given by 20 log10(*A*). Figure 3b, presents the frequency response of an AT-cut Z' propagating/SiO2 LW device designed a fabricated by the authors of this chapter.

**Figure 3.** a) Scheme of a LW delay line. It consists of two ports. In the input IDT an rf signal is applied which launches an acoustic propagating wave. The output signal is recorded at the output IDT. *D* is the distance between input and output IDT and *L* is the center-to-center distance between the IDTs. b) Frequency response of a LW device designed a fabricated by the authors of this chapter. The phase shift (dotted line) and IL (solid line) were measured using a net‐ work analyzer.

In biosensors, biochemical interactions at the sensing area will modify the thickness and properties of the coating, and therefore variations in the amplitude and phase of the electri‐ cal transfer function can be measured. These variations can be monitored in real time, which provides valuable information about the interaction process.

The LW delay line can be used as frequency determining element of an oscillator circuit (*closed loop* configuration). Effectively, in an oscillator circuit the LW device is placed as a de‐ lay line in the feedback loop of an rf amplifier in a closed loop configuration [10,45]. There‐ fore, a change in the wave velocity, due to a sensing effect, produces a time delay in the signal through the LW device which appears as phase-shift; this phase-shift is transferred in terms of frequency-shift in an oscillator configuration. The oscillator is, apparently, the sim‐ plest electronic setup: the low cost of their circuitry as well as the integration capability and continuous monitoring are some features which make the oscillators an attractive configura‐ tion for the monitoring of the determining parameter of the resonator sensor, which in the case of the LW device is the phase-shift of the signal at resonance [46-49]. However, due to the following drawbacks, in our opinion, the oscillators are not the best option for acoustic wave sensor characterization: 1) they do not provide direct information about signal ampli‐ tude; 2) they, eventually, can stop oscillation if insertion losses exceed the amplifier gain during an experiment; and 3) despite of the apparent simple configuration, a very good de‐ sign is necessary to guarantee that a LW resonator will operate at a specific frequency, and this is not a simple task. In effect, in the same way than in QCM oscillators it is required to assure that the sensor resonates on one defined resonance mode and does not "jump" be‐ tween spurious resonances [7], in LW oscillators one must assure that the sensor will oper‐ ate at one phase ramp in the sensor response band-pass, and does not jump from one to another which are almost of identical characteristics (see Figure 3b). Moreover, when the resonator dimensions get smaller and the frequency increases this becomes more difficult to achieve, since when increasing frequency there is a decrease of the resonator quality factor, a decrease in frequency stability [50] and in LW the ramps become nearer to each other.

In an *open loop* configuration, the input transducer is excited at a fixed frequency while the phase shift between *Vout* and *Vin*, *φ,* is recorded [32]. In this configuration, in the absence of interferences, phase variations measured experimentally can be related to changes in the physical properties of the layers deposited over the sensing area.

Network analyzers are the most commonly used instrumentation for characterizing LW de‐ lay lines in open loop configurations. Nevertheless, recently, some authors successfully vali‐ dated a new characterization technique based on the open loop configuration [51]. A read out circuit based on this technique for high frequency liquid loaded QCM devices has been developed by the same authors [52], and tested with LW devices with very satisfactory re‐ sults [53]. The main advantages of this read out circuit are its low cost, high integration, small size, calibration facility and the possibility of being used as an interface for multi-anal‐ ysis detection.
