**2. Polyelectrolyte multilayer film**

#### **2.1. Biomaterials: Generality and interest**

During a consensus conference in 1986, a definition was given for biomaterials. Indeed, a bi‐ omaterial is «a non-living material used and designed to be integrated with biological sys‐ tems». Biomaterials are defined according to their domain of use and regroup metals and alloys, ceramics, polymers (i.e. collagen)[3].

Biomaterials were used since the pharaoh's time to replace injured and affected organs. Pharaoh had used pure natural materials but presenting integration's problems. Since that, researches had grown up rapidly in this field in order to design the "ideal" material which will be more accepted by the human body. The designed material was referred to as "bioma‐ terial" afterwards and will recover a lot of biomedical applications for implants and tissues injuries covering.

Biomaterials' design must take into account the purpose and the place of its use. This bioma‐ terial must have a well defined shape depending on his position within the body. Indeed, for orthopedic usage, a biomaterial must conform to some criteria and regulations such as: a good mechanical structure, a good resistance to corrosion and metal fatigue. For vascular surgery, a biomaterial must not induce thrombosis, in odontology a biomaterial must with‐ stand changes that can occur to temperature (coffee, cool drinks), to pH (alcohol, lemon…) and to the buccal cavity [4].

Making reliable and cheap biomaterials is being a new challenge for researchers and indus‐ tries. In fact, the infallibility of every biomaterial depends on the materials from which it's made of. Consequently, there's a great demand in developing new suitable biomaterials (or making the existing ones better) used in multidisciplinary fields and involving physics, chemistry and biology.

In this study, the biomaterials used for fibroblasts adhesion are made of polyelectrolytes us‐ ing the layer-by-layer technique based on alternating oppositely charged polyelectrolytes on glass probes (more details are shown in paragraph III.2).

#### **2.2. Polyelectrolytes**

Besides keeping a multicellular organism together, cell adhesion is also a source of specific signals to adherent cells; their phenotype can thus be regulated by their adhesive interac‐ tions. In fact, most of the cell adhesion receptors were found to be involved in signal trans‐ duction. By interacting with growth factor receptors they are able to modulate their signaling efficiency. Therefore, gene expression, cytoskeletal dynamics and growth regula‐

In this chapter, I tried to find a possible correlation between polyelectrolyte multilayer films and human gingival fibroblasts to test these biomaterials biocompatibility. This represents a fundamental step needed to know about a possible use in a biological field (i.e. as implant). For that purpose, I characterized each solid surface used as a surface on which fibroblasts were cultured; by calculating their surface free energy and evaluating their chemical hetero‐ geneity, roughness and wettability using contact angle measurement. Thereafter, I followed

During a consensus conference in 1986, a definition was given for biomaterials. Indeed, a bi‐ omaterial is «a non-living material used and designed to be integrated with biological sys‐ tems». Biomaterials are defined according to their domain of use and regroup metals and

Biomaterials were used since the pharaoh's time to replace injured and affected organs. Pharaoh had used pure natural materials but presenting integration's problems. Since that, researches had grown up rapidly in this field in order to design the "ideal" material which will be more accepted by the human body. The designed material was referred to as "bioma‐ terial" afterwards and will recover a lot of biomedical applications for implants and tissues

Biomaterials' design must take into account the purpose and the place of its use. This bioma‐ terial must have a well defined shape depending on his position within the body. Indeed, for orthopedic usage, a biomaterial must conform to some criteria and regulations such as: a good mechanical structure, a good resistance to corrosion and metal fatigue. For vascular surgery, a biomaterial must not induce thrombosis, in odontology a biomaterial must with‐ stand changes that can occur to temperature (coffee, cool drinks), to pH (alcohol, lemon…)

Making reliable and cheap biomaterials is being a new challenge for researchers and indus‐ tries. In fact, the infallibility of every biomaterial depends on the materials from which it's made of. Consequently, there's a great demand in developing new suitable biomaterials (or making the existing ones better) used in multidisciplinary fields and involving physics,

tion all depend, at least partially, on cell adhesive interactions [2].

the adhesion of fibroblasts, their proliferation and their morphology.

**2. Polyelectrolyte multilayer film**

208 Advances in Biomaterials Science and Biomedical Applications

**2.1. Biomaterials: Generality and interest**

alloys, ceramics, polymers (i.e. collagen)[3].

injuries covering.

and to the buccal cavity [4].

chemistry and biology.

Polyelectrolytes are highly charged nanoscopic objects or macromolecules. Their electric charge density appears as more or less continuous, when it is seen from distances to the macromolecule equal to several times to the intercharge distance, giving them the polyelec‐ trolytic character. Obviously, their properties will be extremely different according to their geometry. Massive spherical objects will behave like colloids, whereas linear flexible objects will keep some of the macromolecular polymeric character [5]. They are defined as materials for which the solution's properties in dissolvent presenting a high permittivity are governed by electrostatic interactions for distances superior to the molecular dimensions [6]. Polyelec‐ trolytes are by no way a mere superposition of electrolytes and polymers properties. New and rather unexpected behaviours are observed:


These materials are widely used in industries as dispersive substances in aqueous medium, flocculants to aggregate sludge and industrial waste. Recently, they were used to make films by alternating thin layers of polymers of medical use such as dental prosthesis, fabrication of transplantable organs etc…

Polymers differ by their structure, their surface composition and their biological properties:

#### *2.2.1. Biological properties*

The biological properties reflect the origin of polymers. Indeed, one can distinguish three different origins for polymers [7]:


#### *2.2.2. Physico-chemical properties*

According to Oudet [7], polymers have different physical properties. The most important are their thermal conductivity reflecting polymers' behaviour under temperature changes. The second interesting physical property is their optical reactions towards light (refraction, reflection angle, polarization, absorption…). Moreover, polymers are characterized by their ability as electrical conductors or insulators.

From the chemical point of view, Fowkes [8] presumed the existence of different polymers surface structure: polymers with polar surfaces (polyethylene), polymers with acid (polyi‐ mide) or basic (polystyrene) sites dominance and others are regrouping both acid and basic characters (polyamide). These surfaces are governed by specific (dispersive forces attraction) and non-specific interactions (acid-base interaction).

Polymers properties are strongly influenced by molecular interactions such as Van der Waals interactions (low energy bonds), hydrogen interactions (low energy bonds having an electrostatic origin) and ionic interactions due to electrostatic attractions and repulsions be‐ tween ions or ionized groups.

#### **2.3. Polyelectrolyte multilayer film**

#### *2.3.1. Generality*

In recent years, polyelectrolyte multilayer film has been widely developed in different fields and for a variety of purposes. This kind of ultrathin film can be fabricated from oppositely charged polyelectrolytes using a method called self-assembly discovered by Decher and coworkers in 1992 and allows surface modification and therefore controlling their properties at the molecular (or even the atomic) level.

These films are of a great interest for covering biomaterials used as implants [9, 10] and therefore they will be in contact with cells [11]. Layer-by-layer assembly of polyelectrolytes is a simple and suitable method for coating different substrates such as glass, silicon, ther‐ moplastic and even curved surfaces [12, 13].

It's known that biomaterials must present two main conditions to be admitted for integra‐ tion in the biological system: to be biocompatible with this system and to have definite me‐ chanical and electrical properties depending on their use [14]. The next implants generation has a tendency to be bioactive, besides its biocompatibility, thanks to substrate coating with bioactive substances.

#### *2.3.2. Fabrication method and application fields*

Multilayer polyelectrolyte films are made by alternating oppositely charged polyelectrolytes (polyanions and polycations) on glass slides (Figure 1).

Film's formation is based on charge overcompensation of the newly adsorbed polyions. In‐ deed, a polyanion (negative charge) added to a polycation (positive charge), previously de‐ posited on the substrate, will neutralise the excess of positive charges and therefore create a new negatively charged polyelectrolyte layer. This step can be repeated as many times as the needed number of layers is reached [15].

**Figure 1.** Layer-by-layer polyelectrolyte film's fabrication. This assembly method is based on alternating oppositely charge polyanion (positive charge) and polycations (negative charge) on a solid substrate. One bilayer consists in one polycation associated with one polyanion and the film is a set of n bilayers.

This adsorption mechanism is governed by electrostatic interactions which represent, be‐ sides other secondary interactions (hydrogen bond or dispersive force), a paramount param‐ eter for the final structure of the formed film [16].

Polyelectrolyte multilayer films are used in different fields: orthopedic surgery (hip prosthe‐ sis...), cardiovascular (artificial heart, vascular prosthesis...), odontology (dental restora‐ tion...), ophthalmology (contact lenses...), urology (catheters, artificial kidney...), endocrinology (artificial pancreas, biosensors...), aesthetic surgery and other domains [17].

### **2.4. Polyelectrolyte film surface characterization**

This study is possible by investigating surface wettability and calculating surface free ener‐ gy. Indeed, wettability is the aptitude of a substrate to be coated by a thin liquid film while dipped in a liquid solution. This method is used to follow the substrate behaviour in relation to its environment and can be done thanks to the contact angle measurement. In this paper we are interested in the dynamic contact angle method using Wilhelmy plate method, treat‐ ed later. This method, besides giving information about substrate surface hydrophilicity and hydrophobicity, allows us to evaluate the surface roughness and chemical heterogeneity. Moreover, with the results found, we measured the polyelectrolyte film's surface free ener‐ gy according to Van Oss theory.

#### *2.4.1. Contact angle measurement*

The second interesting physical property is their optical reactions towards light (refraction, reflection angle, polarization, absorption…). Moreover, polymers are characterized by their

From the chemical point of view, Fowkes [8] presumed the existence of different polymers surface structure: polymers with polar surfaces (polyethylene), polymers with acid (polyi‐ mide) or basic (polystyrene) sites dominance and others are regrouping both acid and basic characters (polyamide). These surfaces are governed by specific (dispersive forces attraction)

Polymers properties are strongly influenced by molecular interactions such as Van der Waals interactions (low energy bonds), hydrogen interactions (low energy bonds having an electrostatic origin) and ionic interactions due to electrostatic attractions and repulsions be‐

In recent years, polyelectrolyte multilayer film has been widely developed in different fields and for a variety of purposes. This kind of ultrathin film can be fabricated from oppositely charged polyelectrolytes using a method called self-assembly discovered by Decher and coworkers in 1992 and allows surface modification and therefore controlling their properties at

These films are of a great interest for covering biomaterials used as implants [9, 10] and therefore they will be in contact with cells [11]. Layer-by-layer assembly of polyelectrolytes is a simple and suitable method for coating different substrates such as glass, silicon, ther‐

It's known that biomaterials must present two main conditions to be admitted for integra‐ tion in the biological system: to be biocompatible with this system and to have definite me‐ chanical and electrical properties depending on their use [14]. The next implants generation has a tendency to be bioactive, besides its biocompatibility, thanks to substrate coating with

Multilayer polyelectrolyte films are made by alternating oppositely charged polyelectrolytes

Film's formation is based on charge overcompensation of the newly adsorbed polyions. In‐ deed, a polyanion (negative charge) added to a polycation (positive charge), previously de‐ posited on the substrate, will neutralise the excess of positive charges and therefore create a new negatively charged polyelectrolyte layer. This step can be repeated as many times as

ability as electrical conductors or insulators.

210 Advances in Biomaterials Science and Biomedical Applications

tween ions or ionized groups.

*2.3.1. Generality*

bioactive substances.

**2.3. Polyelectrolyte multilayer film**

the molecular (or even the atomic) level.

moplastic and even curved surfaces [12, 13].

*2.3.2. Fabrication method and application fields*

the needed number of layers is reached [15].

(polyanions and polycations) on glass slides (Figure 1).

and non-specific interactions (acid-base interaction).

There are a variety of simple and inexpensive techniques for measuring contact angles, most of which are described in detail in various texts and publications and will be men‐ tioned only briefly here. The most common direct methods (Figure 2) include the sessile drop (a), the captive bubble (b) and the tilting plate (c). Indirect methods include tensi‐ ometry and geometric analysis of the shape of a meniscus. For solids for which the above methods are not applicable, such as powders and porous materials, methods based on capillary pressures, sedimentation rates, wetting times, imbibition rates, and other properties, have been developed [18].

**Figure 2.** The more common systems of contact angle measurement showing the sessile drop (a), the captive bubble (b) and the tilting plate (c). θ is the contact angle to be measured.

#### *2.4.1.1. The sessile drop method*

It's a static contact angle measurement method which consists in putting down a liquid drop on the solid plate we want to characterize its surface by measuring the contact angle made by the drop on this surface. Indeed, when a drop of a liquid is putted down on a solid sur‐ face; three phases system occurs: solid, liquid and gas (Figure 3).

**Figure 3.** Static contact angle measurement with the sessile drop method

The drop's profile is being changed depending on the physico-chemical characters of the solid surface, on the adhesion forces newly created at the interface solid/liquid and on the cohesion forces of the liquid. This change will affect the contact angle value reveal‐ ing the surface state (hydrophobic or hydrophilic, rough or smooth, homogeneous or het‐ erogeneous…) and the different forces occurred are linked together according to Young's equation [19]:

### *γsv* =*γsl* + *γlvcosθ*,

drop (a), the captive bubble (b) and the tilting plate (c). Indirect methods include tensi‐ ometry and geometric analysis of the shape of a meniscus. For solids for which the above methods are not applicable, such as powders and porous materials, methods based on capillary pressures, sedimentation rates, wetting times, imbibition rates, and

**Figure 2.** The more common systems of contact angle measurement showing the sessile drop (a), the captive bubble

It's a static contact angle measurement method which consists in putting down a liquid drop on the solid plate we want to characterize its surface by measuring the contact angle made by the drop on this surface. Indeed, when a drop of a liquid is putted down on a solid sur‐

other properties, have been developed [18].

212 Advances in Biomaterials Science and Biomedical Applications

(b) and the tilting plate (c). θ is the contact angle to be measured.

face; three phases system occurs: solid, liquid and gas (Figure 3).

**Figure 3.** Static contact angle measurement with the sessile drop method

*2.4.1.1. The sessile drop method*

Where γsv, γsl and γlv represent the "surface tensions" of the interface solid/gas, solid/liquid and liquid/gas, respectively, and θ represents the contact angle.

### *2.4.1.2. The captive bubble method*

It's a derivative of the sessile drop method and consists in making an air bubble (or a bubble from a less dense and non miscible liquid such as dodecane, octane and octadecane) on a solid surface immersed in pure water or in other liquid with a well known physico-chemical characters. So, it's possible to measure the contact angle made by this bubble with the im‐ mersed solid surface (see Figure 2).

#### *2.4.1.3. The tilting plate method*

The tilting plate method is to slowly tilt a contact angle sample until the sessile drop on it begins to move in the downhill direction. At that time, the downhill contact angle is the ad‐ vancing angle and the uphill angle the receding contact angle [20].

The principal alternative to the tilting plate method is having the dispense needle remain immersed in the sessile drop and pumping in until the drop expands in base area and pumping out until the drop contracts in base area. Often the tilting plate measurement is carried out on an instrument with a mechanical platform that tilts the stage and the camera together.

It has been shown that these methods are a subject of controversy. However, the dynamic contact angle measurement using the Wilhelmy plate method has been shown to be easier for use and gives more information about the surface characterized.

#### *2.4.1.4. The dynamic contact angle method: The tensiometer*

In our study, we used the Wilhelmy plate method (Tensiometer 3S, GBX, France) which al‐ lows a dynamic measurement of the contact angle hysteresis. Indeed, the tensiometer used for the measurement will measure the force applied to the substrate while immerged in a liquid thanks to a balance where the substrate was hanged (Figure 4)

In each case, the polyelectrolyte film coated glass slide was immersed into and then drawn out of the measurement liquid. Therefore, the tensiometer will evaluate the ad‐ vancing angle (θa) when the liquid moves forward the substrate surface and thereafter the receding angle (θr) when the liquid resorbs from the substrate. The difference be‐ tween θa and θr is called contact angle hysteresis H (H = θa - θr) and is useful for under‐ standing the wettability of the film. It gives us information about the surface film mobility, its reorganization and roughness [21].

**Figure 4.** The Wilhelmy plate method for dynamic contact angle measurement. The surface plate is partially im‐ merged in the up down moving liquid container. Curves (Loops) are automatically drawn by a software associated to the Tensiometer according to F = f (Immersion depth)

When a substrate is immersed in a liquid, three forces occur (see Figure 4): the gravity force, the upthrust buoyancy and the capillary forces. Therefore, by measuring the applied force according to the immersion depth and as we previously know the dimension of the sub‐ strate; one can calculate the wetting forces according to the equation [22]:

$$\mathbf{F} = \mathbf{m} \mathbf{g} + \mathbf{p} \; \text{\*} \; \gamma\_{\text{LV}} \; \text{\*} \; \cos \theta \; \text{--} \; \text{F}\_{\text{b}} \tag{1}$$

Where F represents the force measured (mN/m), m is the substratum mass, g is the accelera‐ tion constant induced by the gravity, p is the substratum perimeter (cm2 ), γLV is the surface free energy (mN/m) of the liquid used for measurement (constant), θ: the contact angle be‐ tween the liquid and the substratum (°) and Fb is the force related to the upthrust buoyancy.

Usually, we make several immersion/emersion cycles for the substratum we are investi‐ gating and the different loops (one loop corresponds to one immersion/emersion cycle) are drawn by a software associated to the Tensiometer according to Force = f ( immer‐ sion depth). Moreover, the substratum weight is assumed to be nil by a direct correction fixing the pre-immersion force to the value of zero. Therefore, the previous equation ([Eq. 1]) becomes:

F(zero immersion)=p\*γLV\*cosθ

As the surface energy of the liquid of measurement is previously known, therefore the con‐ tact angle could be deduced.

It has been shown that the contact angle changes depending on the nature of the film and on its charges and thickness. The nature of liquid of measurement, the speed and temperature of measurement are also involved in this change [23]. Indeed, the thickness of the film can affect its elasticity which will induce a difference in the liquid diffusion into this film and therefore the film's swelling level changes affecting the contact angle. A previous study made by Elbert et al.[24] has shown a clear effect of the film layers' number on the wettabili‐ ty of the film.

The liquid used for measurement can affect the surface wettability by the mean of its pH which varies from a liquid to another and controls the acid or base character as well as the liquid polarity. These parameters are responsible for the rearrangement of the biomaterial's groups at its contact. This reorganization is also depending on the liquid diffusion into the polymer and on the effect of solubilization induced by the liquid to this polymer. This phe‐ nomenon represents an interesting mechanism for explaining contact angle hysteresis espe‐ cially when the liquid concerned is water. Indeed, water has small molecules which allow it to diffuse easily. Therefore, after diffusion into a polymer, water will confer its hydrophilic character to this polymer which is being to have some kind of elasticity responsible for the reorganization of its polar groups as a reaction to the high surface energy level of water which is responsible for the high energy level at the interface [25]. Concerning the dynamic contact angle measurement speed, it affects the contact period between the biomaterial and the liquid and therefore it will change the period of time needed for the rearrangement of the surface polar groups during contact with the liquid. As each film has its own defined reorganization time, therefore different contact angles were found for the same surface at different measurement speeds. Moreover, every polymer has a defined glass transition tem‐ perature (Tg) able to induce a change on the surface wettability depending on the tempera‐ ture of measurement [26].

#### *2.4.2. Surface free energy calculation*

tween θa and θr is called contact angle hysteresis H (H = θa - θr) and is useful for under‐ standing the wettability of the film. It gives us information about the surface film

**Figure 4.** The Wilhelmy plate method for dynamic contact angle measurement. The surface plate is partially im‐ merged in the up down moving liquid container. Curves (Loops) are automatically drawn by a software associated to

When a substrate is immersed in a liquid, three forces occur (see Figure 4): the gravity force, the upthrust buoyancy and the capillary forces. Therefore, by measuring the applied force according to the immersion depth and as we previously know the dimension of the sub‐

Where F represents the force measured (mN/m), m is the substratum mass, g is the accelera‐

free energy (mN/m) of the liquid used for measurement (constant), θ: the contact angle be‐ tween the liquid and the substratum (°) and Fb is the force related to the upthrust buoyancy. Usually, we make several immersion/emersion cycles for the substratum we are investi‐ gating and the different loops (one loop corresponds to one immersion/emersion cycle) are drawn by a software associated to the Tensiometer according to Force = f ( immer‐ sion depth). Moreover, the substratum weight is assumed to be nil by a direct correction fixing the pre-immersion force to the value of zero. Therefore, the previous equation

F= mg + p \* γLV \* cos θ− Fb (1)

), γLV is the surface

strate; one can calculate the wetting forces according to the equation [22]:

tion constant induced by the gravity, p is the substratum perimeter (cm2

mobility, its reorganization and roughness [21].

214 Advances in Biomaterials Science and Biomedical Applications

the Tensiometer according to F = f (Immersion depth)

([Eq. 1]) becomes:

F(zero immersion)=p\*γLV\*cosθ

It's interesting to know the value of surface free energy of a biomaterial because it has an effect on wettability as shown by Van Oss [27]. While the contact between the biomaterial and the liquid generates an interface solid/liquid which will consume, during its formation, a defined energy called the interface energy. The reversible adhesion force represents, there‐ fore, the difference in the energy level between the initial state characterized by two surfaces [28, 29]: solid surface with the energy (γs) and liquid surface with the energy (γ<sup>l</sup> ); and the final state (γsl).

The surface free energy is a kind of an attraction force of the surface which cannot be meas‐ ure directly but calculated after contact angle measurement in different measurement liq‐ uids (with different surface free energies) according to Owends et Wendt or Van Oss' approaches. Their theories are complementary but Van Oss' approach has been shown to give more information. It consists in the following equation [27]:

*γ<sup>S</sup>* =*γ<sup>S</sup> LW* + 2 (*γ<sup>S</sup>* +.*γ<sup>S</sup>* <sup>−</sup>)<sup>½</sup>

where γS represents the surface free energy of the biomaterial surface, γ<sup>S</sup> WL : the dispersive component and γ<sup>S</sup> + , γ<sup>S</sup> represent the polar components (acid- base).

The different components of the solid and the liquid surface free energies as well as the con‐ tact angle are related by this equation:

$$\gamma\_L \text{ (1 + \cos\theta) = 2 (\{\gamma\_S^{-LW} \cdot \gamma\_L^{-LW}\}^\natural + \{\gamma\_S^\* \cdot \gamma\_L^\* \cdot \gamma\_L^{-}\}^\natural + \{\gamma\_L^\* \cdot \gamma\_S^\* \cdot \}^\natural)$$

This equation contains three unknown parameters: γ LW, γ <sup>+</sup> and γ - ; the contact angle meas‐ urement must be done with three different measurement liquids in order to solve this equa‐ tion and calculate the surface free energy of our polyelectrolyte film. For this purpose, we used three different liquids: water, diiodomethane and formamide.

#### *2.4.3. Evaluation of the surface roughness and heterogeneity*

Theses parameters are deduced from the shapes of the curves drawn (loops). Indeed, the more the surface is rough; the more the curve is deformed (non linear curve). However, the more the surface is smooth; the more the curve presents a linear shape (no deformations ob‐ served). Otherwise, a roughness of about 100 nm has been shown to induce contact angle hysteresis. As for surface heterogeneity, it can be concluded from the different contact angle hysteresis values measured in the case of a negligible roughness.

Concerning the different polyelectrolyte films used in this study, a previous investigation was made by Picart and coworkers [30]. They measured the roughness by the AFM tech‐ nique, refractive index and thickness are estimated by optical waveguide light mode spectroscopy, and zeta potential is measured by streaming potential measurements. In‐ deed, these parameters give us information about the chemical heterogeneity of the poly‐ electrolyte used.

Many studies had observed an important dependence of the contact angle hysteresis on the surface composition and topography (roughness) [31, 32]. Therefore, the more the sur‐ face is rough; the more it's hydrophilic and vise versa and the more this surface is com‐ posed of small molecules, the less the liquid diffusion in the biomaterial surface is disturbed leading to a low contact angle value. According to Morra et al.[33], this is may be due to existence of two different effects while studying the wettability of rough and homogeneous biomaterials: the barrier effect, where hysteresis increases with increasing the surface roughness due to an important rigidity of the substrate, and the capillary at‐ traction at the surface which can affect Young's concept. Indeed, the capillary effect indu‐ ces an increase of both the advancing and receding contact angles in the case of a surface presenting a contact angle superior to 90° at the equilibrium state. In the opposite case (contact angle inferior to 90° at the equilibrium state), the inverse situation will happen and the contact angle variations will be less important than those corresponding to the barrier effect. Only in the case of a contact angle equals to 90°, the capillary effect is neg‐ ligible.

#### **2.5. Conclusion**

*γ<sup>S</sup>* =*γ<sup>S</sup> LW* + 2 (*γ<sup>S</sup>* +.*γ<sup>S</sup>* <sup>−</sup>)<sup>½</sup>

+ , γ<sup>S</sup> -

216 Advances in Biomaterials Science and Biomedical Applications

tact angle are related by this equation:

component and γ<sup>S</sup>

electrolyte used.

ligible.

where γS represents the surface free energy of the biomaterial surface, γ<sup>S</sup>

*γ<sup>L</sup>* (1 + *cosθ*) = 2 ((*γ<sup>S</sup> LW* .*γ<sup>L</sup> LW* )<sup>½</sup> + (*γ<sup>S</sup>* +.*γ<sup>L</sup>* <sup>−</sup>)<sup>½</sup> + (*γ<sup>L</sup>* +.*γ<sup>S</sup>* <sup>−</sup>)½),

This equation contains three unknown parameters: γ LW, γ <sup>+</sup>

*2.4.3. Evaluation of the surface roughness and heterogeneity*

used three different liquids: water, diiodomethane and formamide.

hysteresis values measured in the case of a negligible roughness.

represent the polar components (acid- base).

The different components of the solid and the liquid surface free energies as well as the con‐

urement must be done with three different measurement liquids in order to solve this equa‐ tion and calculate the surface free energy of our polyelectrolyte film. For this purpose, we

Theses parameters are deduced from the shapes of the curves drawn (loops). Indeed, the more the surface is rough; the more the curve is deformed (non linear curve). However, the more the surface is smooth; the more the curve presents a linear shape (no deformations ob‐ served). Otherwise, a roughness of about 100 nm has been shown to induce contact angle hysteresis. As for surface heterogeneity, it can be concluded from the different contact angle

Concerning the different polyelectrolyte films used in this study, a previous investigation was made by Picart and coworkers [30]. They measured the roughness by the AFM tech‐ nique, refractive index and thickness are estimated by optical waveguide light mode spectroscopy, and zeta potential is measured by streaming potential measurements. In‐ deed, these parameters give us information about the chemical heterogeneity of the poly‐

Many studies had observed an important dependence of the contact angle hysteresis on the surface composition and topography (roughness) [31, 32]. Therefore, the more the sur‐ face is rough; the more it's hydrophilic and vise versa and the more this surface is com‐ posed of small molecules, the less the liquid diffusion in the biomaterial surface is disturbed leading to a low contact angle value. According to Morra et al.[33], this is may be due to existence of two different effects while studying the wettability of rough and homogeneous biomaterials: the barrier effect, where hysteresis increases with increasing the surface roughness due to an important rigidity of the substrate, and the capillary at‐ traction at the surface which can affect Young's concept. Indeed, the capillary effect indu‐ ces an increase of both the advancing and receding contact angles in the case of a surface presenting a contact angle superior to 90° at the equilibrium state. In the opposite case (contact angle inferior to 90° at the equilibrium state), the inverse situation will happen and the contact angle variations will be less important than those corresponding to the barrier effect. Only in the case of a contact angle equals to 90°, the capillary effect is neg‐

and γ -

WL : the dispersive

; the contact angle meas‐

When a drop of liquid is placed on a solid surface, the liquid will either spread across the surface to form a thin, approximately uniform film or spread to a limited extent but remain as a discrete drop on the surface. The final condition of the applied liquid to the surface is taken as an indication of the wettability of the surface by the liquid or the wet‐ ting ability of the liquid on the surface. The quantitative measure of the wetting process is taken to be the contact angle, which the drop makes with the solid as measured through the liquid in question.

The wetting of a surface by a liquid and the ultimate extent of spreading of that liquid are very important aspects of practical surface chemistry. Many of the phenomenological aspects of the wetting processes have been recognized and quantified since early in the history of observation of such processes. However, the microscopic details of what is oc‐ curring at the various interfaces and lines of contact among phases has been more a sub‐ ject of conjecture and theory than of known facts until the latter part of this century when quantum electrodynamics and elegant analytical procedures began to provide a great deal of new insight into events at the molecular level. Even with all the new infor‐ mation of the last 20 years, however, there still remains a great deal to learn about the mechanisms of movement of a liquid across a surface.
