Phosphates in Biomaterials

**117**

**Chapter 7**

**Abstract**

Calcium Phosphate Cements in

*Manuel Pedro Fernandes Graça and Sílvia Rodrigues Gavinho*

Calcium phosphate cements (CPCs) consist of a combination of calcium phosphates and a liquid phase, allowing it to fit into the body where it was inserted. Several chemical compositions have been synthesized, promoting specific characteristics to the cements for applications such as bone augmentation and reinforcement and metal implant fixation. The hardening reaction mechanism is at low temperatures and makes it capable of incorporating different drugs and other biological molecules. In addition to the abovementioned advantages, CPCs have excellent bioactivity and osteoconductivity and the ability to form a bone bond. Its function as osteoconductor can be improved by insertion of growth factors. In addition, it is possible to functionalize it with silver ions and use it as a coating of implants, conferring antibacterial properties. In this chapter the physical, mechanical, chemical, and biological properties and the possibility of using these cements as

**Keywords:** calcium phosphates, bone cements, tissue regeneration, drug delivery,

Calcium phosphate cements (CPCs) were proposed by Brown and Chow [1] and LeGeros et al. [2] in the 1980s. In 1990, the first CPC was used commercially in the

CPCs consist of a combination of one or more calcium orthophosphate powders in which a liquid phase, usually water or an aqueous solution, is added, allowing it to be set and hardened at the site of the body where it was implanted. This type of cement hardens through a dissolution reaction and a precipitation process, distinguishing itself from other cements that harden through a polymerization reaction. Over time, new compositions have been synthesized promoting specific characteristics to the cements for various applications such as bone augmentation and strengthening [6–14], fixation of metal implants [15, 16], and vertebral fractures [17–19]. The CPCs have the essential advantage of hardening in vivo through a lowtemperature reaction. After mixing, the material becomes moldable, and its

noninvasive injection represents an important advantage over conventional calcium phosphate ceramics. The fact that this type of cement does not present an exothermic reaction also makes it able to incorporate different drugs and other biological molecules, allowing its application in treatments by drug delivery [20]. In addition to the aforementioned advantages, CPCs have excellent bioactivity and osteoconductivity and an excellent ability to form a bone bond. Its rate of resorption is also

Tissue Engineering

drug carriers or biomolecules will be discussed.

treatment of maxillofacial defects and fractures [3–5].

osseointegration, antibacterial properties

**1. Calcium phosphate cements**

#### **Chapter 7**

## Calcium Phosphate Cements in Tissue Engineering

*Manuel Pedro Fernandes Graça and Sílvia Rodrigues Gavinho*

#### **Abstract**

Calcium phosphate cements (CPCs) consist of a combination of calcium phosphates and a liquid phase, allowing it to fit into the body where it was inserted. Several chemical compositions have been synthesized, promoting specific characteristics to the cements for applications such as bone augmentation and reinforcement and metal implant fixation. The hardening reaction mechanism is at low temperatures and makes it capable of incorporating different drugs and other biological molecules. In addition to the abovementioned advantages, CPCs have excellent bioactivity and osteoconductivity and the ability to form a bone bond. Its function as osteoconductor can be improved by insertion of growth factors. In addition, it is possible to functionalize it with silver ions and use it as a coating of implants, conferring antibacterial properties. In this chapter the physical, mechanical, chemical, and biological properties and the possibility of using these cements as drug carriers or biomolecules will be discussed.

**Keywords:** calcium phosphates, bone cements, tissue regeneration, drug delivery, osseointegration, antibacterial properties

#### **1. Calcium phosphate cements**

Calcium phosphate cements (CPCs) were proposed by Brown and Chow [1] and LeGeros et al. [2] in the 1980s. In 1990, the first CPC was used commercially in the treatment of maxillofacial defects and fractures [3–5].

CPCs consist of a combination of one or more calcium orthophosphate powders in which a liquid phase, usually water or an aqueous solution, is added, allowing it to be set and hardened at the site of the body where it was implanted. This type of cement hardens through a dissolution reaction and a precipitation process, distinguishing itself from other cements that harden through a polymerization reaction. Over time, new compositions have been synthesized promoting specific characteristics to the cements for various applications such as bone augmentation and strengthening [6–14], fixation of metal implants [15, 16], and vertebral fractures [17–19].

The CPCs have the essential advantage of hardening in vivo through a lowtemperature reaction. After mixing, the material becomes moldable, and its noninvasive injection represents an important advantage over conventional calcium phosphate ceramics. The fact that this type of cement does not present an exothermic reaction also makes it able to incorporate different drugs and other biological molecules, allowing its application in treatments by drug delivery [20]. In addition to the aforementioned advantages, CPCs have excellent bioactivity and osteoconductivity and an excellent ability to form a bone bond. Its rate of resorption is also

a factor to take into account, since, after modifying its structure, it is possible to modify that rate. Calcium phosphate cements also have disadvantages, namely, their low mechanical performance, which limits their application in bearing situations. Its intrinsic porosity also leads to this material presenting less strength compared to calcium phosphate ceramics [21, 22].

The cements tend to dissolve in order to achieve a stable and less soluble phase, the dissolution being controlled by the pH of the medium. The cements based on hydroxyapatite or brushite are the only final reaction products because they are the most stable at pH > 4.2 and pH < 4.2, respectively, even though there are a large number of formulations. In addition, and because these materials are intended to be used as bone substitutes, it is important to take into account that the values of the compressive strength of the cortical bone vary between 90 and 209 MPa, [23, 24] and the spongy bone varies between 1.5 and 45 MPa. [25]. As reference values, the compressive strength of apatite cements usually ranges from 20 to 50 MPa [26–32]. Brushite CPCs are generally weaker than apatite CPCs being around 25 MPa [33].

In summary, the main characteristics of this type of cement are presented in **Table 1**.

#### **1.1 Physical and mechanical properties**

The mechanical properties are the main properties to take into account when developing a biomaterial to apply surgically. The microstructural characteristics (porosity, quantity, size, morphology, and distribution of the crystals formed) of a biomaterial are the determining factor to define the mechanical properties. These characteristics are controlled by the synthesis process and its intrinsic parameters. In addition, all factors, such as chemical composition of cement, relative proportions of reactants, powder or liquid additives acting as accelerators or retarders, particle size, liquid-powder ratio (L/P ratio) (**Figure 1**), applied pressure during synthesis, and aging conditions, will affect its mechanical properties [36].

Unlike bioceramics, which require sintering at high temperatures, CPCs are formed through a dissolution-precipitation process at room or body temperature. During this process, a crystalline matrix is formed in which with the passage of time and matrix it becomes increasingly dense until reaching the maximum mechanical properties [36].


**119**

**Figure 1.**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

Although porosity is a disadvantage to be applied in load situations, porosity is sought to improve the resorbability of the material and the extent of bioactivity, increasing the surface area available for reaction. The presence of a certain level of porosity makes this material a good carrier for controlled drug systems [34].

*pores are formed due to the increased separation between the aggregates (adapted from [35]).*

*Microstructure and porosity of CPCs according to particle size and L/P ratio. (A) When the particle size is small, there is an increase in the specific surface area, resulting in the degree of supersaturation. This phenomenon favors the nucleation of crystals and leads to the formation of large quantities of small needle-like crystals. When the particles are larger, the formation of larger plate crystals occurs. Small pores are formed in the cement where small particles are used. (B) Porosity and pore size distribution also vary with the L/P ratio. When the L/P ratio is low, the space between the particles in the mixture decreases, leading to a more compact structure of crystal agglomerates. In contrast, when the L/P ratio increases, the total porosity of the cement increases, and the larger* 

It is possible to promote the creation of macropores in CPCs through two techniques. The leaching of porogen after the adjustment creates macropores; however, it is necessary to add large amounts of porogen, which may compromise the injection process [37–47]. The formation of the pores can also be achieved by the formation of gas foam prior to fixation, but the release of the gas after the introduction of the implant may be harmful to the organism [46, 48–52]. In order to overcome the drawbacks of the techniques mentioned above for obtaining macroporous structures of CPCs, Ginebra et al. proposed the use of self-adjusting injectable macropo-

rous foams composed of a protein-based foaming agent and CPC paste [52]. Regardless of the composition of CPC (apatite or brushite), the strength decreases globally with increased porosity, which is a common occurrence in several materials mainly in porous materials used in bone replacement [36]. In addition to the pore fraction effect, pore size also significantly affects the strength of CPCs. Bai et al. [53] found that the compressive strength is inversely proportional to the size of the macropores through a study in materials with equivalent total porosity but with different sizes of macropores. Through Griffith's classical theory [36], which relates strength to the critical size of the fault, macropores can be considered as failures, thus reducing strength. In addition to the quantity and size of the pores, the characteristics of the crystals (quantity, size, morphology, and distribution) also influence the strength of these cements. The growth of the crystals depends on the kinetics of the dissolution-precipitation reaction of the cement, being controlled by many factors. The smaller the particle size of the starting materials, the faster the material will be converted to apatite, and the crystals formed will not have time to grow. The small size of these crystals will lead to a more dense and crystalline

One of the conditions that influence the kinetics of apatite formation is aging. The transformation of the initial reactants into apatite becomes faster at higher

organization, increasing the strength of the cement [54].

Normally, microporosity ranges between 30 and 55% and is dependent on the L/P ratio; the higher this ratio is, the greater the microporosity [37].

#### **Table 1.**

*The nature and properties of calcium phosphate bone cements (adapted from [34]).*

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

#### **Figure 1.**

*Contemporary Topics about Phosphorus in Biology and Materials*

calcium phosphate ceramics [21, 22].

**1.1 Physical and mechanical properties**

around 25 MPa [33].

in **Table 1**.

properties [36].

**Calcium phosphate properties**

Material type Ceramic

Bioactivity Bioactive

a factor to take into account, since, after modifying its structure, it is possible to modify that rate. Calcium phosphate cements also have disadvantages, namely, their low mechanical performance, which limits their application in bearing situations. Its intrinsic porosity also leads to this material presenting less strength compared to

the dissolution being controlled by the pH of the medium. The cements based on hydroxyapatite or brushite are the only final reaction products because they are the most stable at pH > 4.2 and pH < 4.2, respectively, even though there are a large number of formulations. In addition, and because these materials are intended to be used as bone substitutes, it is important to take into account that the values of the compressive strength of the cortical bone vary between 90 and 209 MPa, [23, 24] and the spongy bone varies between 1.5 and 45 MPa. [25]. As reference values, the compressive strength of apatite cements usually ranges from 20 to 50 MPa [26–32]. Brushite CPCs are generally weaker than apatite CPCs being

In summary, the main characteristics of this type of cement are presented

The mechanical properties are the main properties to take into account when developing a biomaterial to apply surgically. The microstructural characteristics (porosity, quantity, size, morphology, and distribution of the crystals formed) of a biomaterial are the determining factor to define the mechanical properties. These characteristics are controlled by the synthesis process and its intrinsic parameters. In addition, all factors, such as chemical composition of cement, relative proportions of reactants, powder or liquid additives acting as accelerators or retarders, particle size, liquid-powder ratio (L/P ratio) (**Figure 1**), applied pressure during synthesis, and aging conditions, will affect its mechanical properties [36].

Unlike bioceramics, which require sintering at high temperatures, CPCs are formed through a dissolution-precipitation process at room or body temperature. During this process, a crystalline matrix is formed in which with the passage of time and matrix it becomes increasingly dense until reaching the maximum mechanical

Normally, microporosity ranges between 30 and 55% and is dependent on the

Reaction products Calcium phosphates, usually hydroxyapatite or brushite (37°C) Stability Resorbable (low or high resorption rate depending on composition and microstructure)

L/P ratio; the higher this ratio is, the greater the microporosity [37].

Setting reaction mechanism Dissolutions and precipitation reaction

*The nature and properties of calcium phosphate bone cements (adapted from [34]).*

Applications Bone regeneration; non-load-bearing applications

Liquid phase Water or aqueous solutions Powder component Calcium phosphate powders

The cements tend to dissolve in order to achieve a stable and less soluble phase,

**118**

**Table 1.**

*Microstructure and porosity of CPCs according to particle size and L/P ratio. (A) When the particle size is small, there is an increase in the specific surface area, resulting in the degree of supersaturation. This phenomenon favors the nucleation of crystals and leads to the formation of large quantities of small needle-like crystals. When the particles are larger, the formation of larger plate crystals occurs. Small pores are formed in the cement where small particles are used. (B) Porosity and pore size distribution also vary with the L/P ratio. When the L/P ratio is low, the space between the particles in the mixture decreases, leading to a more compact structure of crystal agglomerates. In contrast, when the L/P ratio increases, the total porosity of the cement increases, and the larger pores are formed due to the increased separation between the aggregates (adapted from [35]).*

Although porosity is a disadvantage to be applied in load situations, porosity is sought to improve the resorbability of the material and the extent of bioactivity, increasing the surface area available for reaction. The presence of a certain level of porosity makes this material a good carrier for controlled drug systems [34].

It is possible to promote the creation of macropores in CPCs through two techniques. The leaching of porogen after the adjustment creates macropores; however, it is necessary to add large amounts of porogen, which may compromise the injection process [37–47]. The formation of the pores can also be achieved by the formation of gas foam prior to fixation, but the release of the gas after the introduction of the implant may be harmful to the organism [46, 48–52]. In order to overcome the drawbacks of the techniques mentioned above for obtaining macroporous structures of CPCs, Ginebra et al. proposed the use of self-adjusting injectable macroporous foams composed of a protein-based foaming agent and CPC paste [52].

Regardless of the composition of CPC (apatite or brushite), the strength decreases globally with increased porosity, which is a common occurrence in several materials mainly in porous materials used in bone replacement [36]. In addition to the pore fraction effect, pore size also significantly affects the strength of CPCs. Bai et al. [53] found that the compressive strength is inversely proportional to the size of the macropores through a study in materials with equivalent total porosity but with different sizes of macropores. Through Griffith's classical theory [36], which relates strength to the critical size of the fault, macropores can be considered as failures, thus reducing strength. In addition to the quantity and size of the pores, the characteristics of the crystals (quantity, size, morphology, and distribution) also influence the strength of these cements. The growth of the crystals depends on the kinetics of the dissolution-precipitation reaction of the cement, being controlled by many factors. The smaller the particle size of the starting materials, the faster the material will be converted to apatite, and the crystals formed will not have time to grow. The small size of these crystals will lead to a more dense and crystalline organization, increasing the strength of the cement [54].

One of the conditions that influence the kinetics of apatite formation is aging. The transformation of the initial reactants into apatite becomes faster at higher

temperatures, thus making the structure more homogeneous and denser, making it more resistant. In contrast, high temperatures also cause the development of precipitated apatite crystals more rapidly, resulting in larger crystals, which will negatively influence the resistance [55].

The fixation kinetics may be influenced by the presence of accelerating or retarding substances which are added to the mixture and is an important factor in determining the strength of the structure.

Bermudez et al. [56] and Yang et al. [57] found that by adding certain amounts of apatite, the hardening time of the CPCs is lower, and the compressive strength increases considerably. α-Hydroxylic acids (citric acid or glycolic acid) and their salts (sodium citrate) are also used as retarders. The addition of the same allows to mix and to process more easily, reducing the L/P ratio associated with a decrease of the porosity, improving the strength [58–64]. However, it is necessary to take into account an optimum concentration of these additives since the excess may lead to the opposite effect and decrease the force [63]. In summary, the mechanical properties of the CPCs, and in particular the resistance, depend strongly on the microstructure, which is related to the synthesis process, chemical composition, powder or liquid additives acting as accelerators or retarders, particle size, L/P ratio, and aging conditions. In addition, it has been found that crystalline structures have, with smaller crystals, become more compact and homogeneous and appear to give better mechanical properties than those with larger crystals.

Strength has been the main property to be studied when evaluating mechanical performance; however, the CPCs applied in bone defects are also subject to cyclic loading, and the resistance of CPC to fractures cannot be evaluated by strength alone. In order to adequately evaluate the ability to resist fractures, it is also necessary to take into account fracture toughness that describes the strength of a material containing cracks or notches to resist crack propagation [65, 66].

The toughness of a material depends on its nano-/microstructure and on the possibility of promoting the hardening activation mechanism [67, 68]. Without the activation of significant hardening mechanisms, the fracture toughness of CPCs is very low. Due to the low values of toughness, CPCs are very sensitive to defects and failures. The reliability, that is, the likelihood of failure of brittle materials, is also an important factor when one thinks of applying the cement to load-bearing sites. As previously mentioned, it is possible to improve the hardening mechanism by decreasing porosity, as it is the most damaging factor in mechanical performance. To overcome this problem, it is necessary to decrease the volumetric fraction of the pores in order to achieve a more dense matrix by compacting the cement paste prior to hydration.

Studies have shown that compaction pressure would significantly increase tensile strength [69]. However, when the compaction pressure is above 100 MPa, only a slight decrease in porosity is achieved, and the diametral tensile strength is not substantially improved.

In addition, the use of this method to promote the hardening has the same function of decreasing the L/P ratio, which would influence the workability and injectability of cement pastes, which may exclude the application of this cement in minimally invasive surgery.

In order to overcome this disadvantage, researchers added certain amounts of citric acid to the cement liquid to evaluate its effect on the fixability and fixation properties of apatite cements and found that this addition effectively improves the mechanical properties of the cement. According to this prominent effect of citric acid on strength improvement, Barralet et al. [61] and Gbureck et al. [60] added sodium citrate and compacted the resulting cement slurry, obtaining compressive strengths near the resistance of the cortical bone, that this composition can be used

**121**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

**1.2 Chemical and biological properties**

solution of the reactants [74].

function of pH [71–73].

stable and less soluble phase.

drugs or molecules [75–77].

salts have been added [92].

in load-bearing locations. This resistance can also be achieved by other factors, especially in the use of citric acid but without applying external pressure, varying

The main chemical reaction occurring in the setting mechanism is similar in all these systems of cements and can be understood by analyzing the behavior of the solubility of the existing compounds in the composition [71–73]. During the fixation reaction, the two mechanisms present are dissolution and reprecipitation [23]. The dissolution is activated by the release of the calcium and phosphate ions from the starting materials, leading to supersaturation in the solution. After the ionic concentration reaches a critical value, the nucleation of the new phase occurs, usually around the powder particles. This new phase develops in line with the dis-

In the dissolution/reprecipitation mechanism, the formation of the precipitates depends on the relative stability of the various calcium phosphate salts in the system. The existence of a precipitate that grows in the form of the crystal agglomerates determines the force that a cement can acquire [74]. The solubility phase diagram predicts this reaction, describing the evolution of the solubility of a compound through the logarithm of the total concentration of calcium (or phosphate) as a

The less stable phase of calcium phosphate tends to dissolve to form a more

In the industry, the biological responses of materials have been increasingly a property to be taken into account, focusing on improving cell and tissue CPC interactions, as well as their applications in bone tissue engineering [78–82]. The improvement of these interactions is one important factor for the application of

Studies evaluating the in vivo behavior of CPCs show high levels of biocompatibility and osteoconductivity, stimulating tissue regeneration [5, 82–92]. Most apatite cements are reabsorbed by cell-mediated mechanisms. The function of the osteoclastic cells in this process is to degrade the materials layer by layer, starting from the surface which is in contact with the bone to the nucleus. The biodegradability of apatite CPCs is slow but higher than that of the synthesized hydroxyapatite. As noted above, the rate of degradation of apatite cements is controlled by precipitated hydroxyapatite (PHA) crystallinity, specific surface area, and matrix porosity.

The cements based on brushite have a higher reabsorption rate than apatites due to their superior stability in the biological environment [89–91]. However, there is a possibility in vivo of brushite cements to be transformed into PHA and thus reduce their total degradation rate. In order to retard this reaction or to avoid magnesium

At a biological level, the mechanism of dissolution is mediated by the action of

physiological solutions or by cell-mediated processes (phagocytosis) [93].

biomaterials and their commercialization for clinical applications.

As mentioned above, apatite is the most stable calcium phosphate (less soluble at a pH above 4.2 at room temperature), and brushite is most stable at a pH below 4.2 [71]. In these reactions the amount of water consumed is nonexistent, or almost nil, being necessary only for the reagents to become viable and to allow homogeneity in the solution. For this reason, water becomes one of the main contributions to the development of porosity in cement, and, therefore, CPCs are intrinsically porous materials. In addition to this material, in situ, the body temperature allows its molding after mixing, to be injectable and therefore to be used as a carrier for biological

the particle size and distribution of powdered reagents [70].

*Contemporary Topics about Phosphorus in Biology and Materials*

negatively influence the resistance [55].

determining the strength of the structure.

temperatures, thus making the structure more homogeneous and denser, making it more resistant. In contrast, high temperatures also cause the development of precipitated apatite crystals more rapidly, resulting in larger crystals, which will

The fixation kinetics may be influenced by the presence of accelerating or retarding substances which are added to the mixture and is an important factor in

give better mechanical properties than those with larger crystals.

containing cracks or notches to resist crack propagation [65, 66].

Bermudez et al. [56] and Yang et al. [57] found that by adding certain amounts of apatite, the hardening time of the CPCs is lower, and the compressive strength increases considerably. α-Hydroxylic acids (citric acid or glycolic acid) and their salts (sodium citrate) are also used as retarders. The addition of the same allows to mix and to process more easily, reducing the L/P ratio associated with a decrease of the porosity, improving the strength [58–64]. However, it is necessary to take into account an optimum concentration of these additives since the excess may lead to the opposite effect and decrease the force [63]. In summary, the mechanical properties of the CPCs, and in particular the resistance, depend strongly on the microstructure, which is related to the synthesis process, chemical composition, powder or liquid additives acting as accelerators or retarders, particle size, L/P ratio, and aging conditions. In addition, it has been found that crystalline structures have, with smaller crystals, become more compact and homogeneous and appear to

Strength has been the main property to be studied when evaluating mechanical performance; however, the CPCs applied in bone defects are also subject to cyclic loading, and the resistance of CPC to fractures cannot be evaluated by strength alone. In order to adequately evaluate the ability to resist fractures, it is also necessary to take into account fracture toughness that describes the strength of a material

The toughness of a material depends on its nano-/microstructure and on the possibility of promoting the hardening activation mechanism [67, 68]. Without the activation of significant hardening mechanisms, the fracture toughness of CPCs is very low. Due to the low values of toughness, CPCs are very sensitive to defects and failures. The reliability, that is, the likelihood of failure of brittle materials, is also an important factor when one thinks of applying the cement to load-bearing sites. As previously mentioned, it is possible to improve the hardening mechanism by decreasing porosity, as it is the most damaging factor in mechanical performance. To overcome this problem, it is necessary to decrease the volumetric fraction of the pores in order to achieve a more dense matrix by compacting the cement paste prior

Studies have shown that compaction pressure would significantly increase tensile strength [69]. However, when the compaction pressure is above 100 MPa, only a slight decrease in porosity is achieved, and the diametral tensile strength is

In addition, the use of this method to promote the hardening has the same function of decreasing the L/P ratio, which would influence the workability and injectability of cement pastes, which may exclude the application of this cement in

In order to overcome this disadvantage, researchers added certain amounts of citric acid to the cement liquid to evaluate its effect on the fixability and fixation properties of apatite cements and found that this addition effectively improves the mechanical properties of the cement. According to this prominent effect of citric acid on strength improvement, Barralet et al. [61] and Gbureck et al. [60] added sodium citrate and compacted the resulting cement slurry, obtaining compressive strengths near the resistance of the cortical bone, that this composition can be used

**120**

to hydration.

not substantially improved.

minimally invasive surgery.

in load-bearing locations. This resistance can also be achieved by other factors, especially in the use of citric acid but without applying external pressure, varying the particle size and distribution of powdered reagents [70].

#### **1.2 Chemical and biological properties**

The main chemical reaction occurring in the setting mechanism is similar in all these systems of cements and can be understood by analyzing the behavior of the solubility of the existing compounds in the composition [71–73]. During the fixation reaction, the two mechanisms present are dissolution and reprecipitation [23]. The dissolution is activated by the release of the calcium and phosphate ions from the starting materials, leading to supersaturation in the solution. After the ionic concentration reaches a critical value, the nucleation of the new phase occurs, usually around the powder particles. This new phase develops in line with the dissolution of the reactants [74].

In the dissolution/reprecipitation mechanism, the formation of the precipitates depends on the relative stability of the various calcium phosphate salts in the system. The existence of a precipitate that grows in the form of the crystal agglomerates determines the force that a cement can acquire [74]. The solubility phase diagram predicts this reaction, describing the evolution of the solubility of a compound through the logarithm of the total concentration of calcium (or phosphate) as a function of pH [71–73].

The less stable phase of calcium phosphate tends to dissolve to form a more stable and less soluble phase.

As mentioned above, apatite is the most stable calcium phosphate (less soluble at a pH above 4.2 at room temperature), and brushite is most stable at a pH below 4.2 [71].

In these reactions the amount of water consumed is nonexistent, or almost nil, being necessary only for the reagents to become viable and to allow homogeneity in the solution. For this reason, water becomes one of the main contributions to the development of porosity in cement, and, therefore, CPCs are intrinsically porous materials.

In addition to this material, in situ, the body temperature allows its molding after mixing, to be injectable and therefore to be used as a carrier for biological drugs or molecules [75–77].

In the industry, the biological responses of materials have been increasingly a property to be taken into account, focusing on improving cell and tissue CPC interactions, as well as their applications in bone tissue engineering [78–82]. The improvement of these interactions is one important factor for the application of biomaterials and their commercialization for clinical applications.

Studies evaluating the in vivo behavior of CPCs show high levels of biocompatibility and osteoconductivity, stimulating tissue regeneration [5, 82–92]. Most apatite cements are reabsorbed by cell-mediated mechanisms. The function of the osteoclastic cells in this process is to degrade the materials layer by layer, starting from the surface which is in contact with the bone to the nucleus. The biodegradability of apatite CPCs is slow but higher than that of the synthesized hydroxyapatite. As noted above, the rate of degradation of apatite cements is controlled by precipitated hydroxyapatite (PHA) crystallinity, specific surface area, and matrix porosity.

The cements based on brushite have a higher reabsorption rate than apatites due to their superior stability in the biological environment [89–91]. However, there is a possibility in vivo of brushite cements to be transformed into PHA and thus reduce their total degradation rate. In order to retard this reaction or to avoid magnesium salts have been added [92].

At a biological level, the mechanism of dissolution is mediated by the action of physiological solutions or by cell-mediated processes (phagocytosis) [93].

**Figure 2.**

*Image of the drill hole with progressive resorption of the calcium phosphate cement matrix, during 8 weeks [81].*

Bone replacement depends on the age, sex, and general metabolic health of the host and the site and volume where it is applied, porosity, crystallinity, chemical composition, particle size, and L/P ratio of the cement. Considering these factors, it can take from 3 to 36 months for the cement to be completely reabsorbed and replaced with bone. However, further studies are needed to confirm the total resorption of the material in order to be applied clinically [94]. Studies have revealed bone development around calcium phosphate cements, demonstrating osteoconductive and osteoinductive characteristics in several cases. It has been shown that within 2 weeks, spicules of living bone with normal bone marrow and gaps in osteocytes can be identified in the cement. After 8 weeks, the cement is almost completely surrounded by new bone. At this stage, no cement reabsorption is typically observed [94]. **Figure 2** shows a progressive resorption of the calcium phosphate cement matrix, with tricalcium phosphate (TCP) granules embedded in a matrix of dicalcium phosphate dihydrate (DCPD) and parallel new bone formation, in a drill hole. After 2 weeks almost the entire surface of the cement was in direct contact with the margins of the bone defect. After 4 weeks, occasional granules of β-TCP and the newly formed bone islets are visible. This area expanded after 6 weeks, involving a progressive reabsorption of the cement matrix and parallel neoformed formation [81].

#### **2. Principal calcium phosphate cements**

#### **2.1 Apatite cements**

The importance given in the use of apatites in bone replacement is due to the fact that this mineral is the base of the main inorganic part of hard tissues. In fact, nonstoichiometric or calcium-deficient hydroxyapatite (CDHA) is the main mineral phase characteristic of human bones [94]. The CPCs consist of a network of calcium phosphate crystals, with chemical composition and crystal size that can be modified to approximate the biological hydroxyapatite that exists in the living bone [95, 96].

In this regard, it is necessary to clarify that even though the stoichiometric hydroxyapatite has a fixed composition, the apatite structure may exist in a variety of compositions. CDHA comprises in its composition the possibility of varying amounts of calcium, where it is possible to present a completely deficient structure based on this base element. The composition may be expressed as Ca10 − *x*(HPO4)(PO4)6 − *x*(OH)2 − *x*, where *x* ranges from 0 to 1, 0 for stoichiometric hydroxyapatite and 1 for hydroxyapatite totally deficient in calcium. Biological apatite is deficient in calcium containing various ionic substitutions such as Na+ , K+ , Mg2+, F<sup>−</sup>, and Cl<sup>−</sup> [34].

Apatite cements may form PHA or CDHA through a precipitation reaction. The synthesis of these cements allows the incorporation of different ions in their composition, depending on the initial compounds. The formation of hydroxyapatite that occurs in the cement is compared to the process of formation of new bone

**123**

**Figure 3.**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

due to the Ca and P ratio being maintained [34, 35].

hydroxyapatite similar to the bone mineral [5].

sintering processes [34].

and is also seen as a biomimetic process, because it occurs at body temperature and physiological environment. This may explain the fact that the hydroxyapatite formed in the reactions of calcium phosphate cements is much more similar to biological apatites than the ceramic hydroxyapatite resulting from high-temperature

The CPCs lead to the formation of PHA or CDHA and can be divided into three systems (single compound, two compounds, more than two compounds), taking into account the number and type of calcium phosphates used in the synthesis [97]. Monocomponent CPCs are those having a single calcium phosphate reagent that hydrolyzes to form PHA or CDHA. Taking into account that at pH 4.2 the hydroxyapatite is less soluble, any other calcium phosphate present will dissolve, and the PHA will tend to precipitate. However, when the formation of PHA occurs, from the hydrolysis of calcium phosphate, the reaction mechanism becomes very slow, due to a decrease in the level of supersaturation, as the reaction proceeds [13]. In this in which only one compound is present, no release of any acid or no base occurs

The second type of cement that may exist is composed of two calcium phosphates, one acid and the other basic, where they adjust after an acid–base reaction. The most commonly used compound is usually tetracalcium phosphate (TTCP), as it is the only calcium phosphate with a Ca/P ratio higher than PHA. Therefore, TTCP can be combined with one or more calcium phosphates with lower Ca/P ratios to obtain PHA or CDHA, avoiding the formation of acids or bases as final products. The combinations that have been more studied seek to produce cements that adjust to the body temperature in a range of pH around the neutral [34]. The third possible system consists of more than two compounds, including calcium phosphates and other salts. For example, a cement proposed by Norian Corporation [5] is used where calcium phosphates with a Ca/P ratio lower than PHA and CaCO3 are added as an additional source of calcium. The initial configuration process involves the formation of DCPD, later forming dahllite, a carbonated

Apatitic CPCs appear as a viscous, easily moldable material; however, their injection is difficult. **Figure 3** shows the microstructure of an apatitic cement after setting. The setting time can also be reduced by means of additives such as with the introduction of PHA particles. These changes in the composition may lead to an

*Microstructure of an apatitic calcium phosphate cement after setting, showing the micro−/nanosize pore* 

*structure formed by the entanglement of the precipitated crystals [71].*

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

*Contemporary Topics about Phosphorus in Biology and Materials*

Bone replacement depends on the age, sex, and general metabolic health of the host and the site and volume where it is applied, porosity, crystallinity, chemical composition, particle size, and L/P ratio of the cement. Considering these factors, it can take from 3 to 36 months for the cement to be completely reabsorbed and replaced with bone. However, further studies are needed to confirm the total resorption of the material in order to be applied clinically [94]. Studies have revealed bone development around calcium phosphate cements, demonstrating osteoconductive and osteoinductive characteristics in several cases. It has been shown that within 2 weeks, spicules of living bone with normal bone marrow and gaps in osteocytes can be identified in the cement. After 8 weeks, the cement is almost completely surrounded by new bone. At this stage, no cement reabsorption is typically observed [94]. **Figure 2** shows a progressive resorption of the calcium phosphate cement matrix, with tricalcium phosphate (TCP) granules embedded in a matrix of dicalcium phosphate dihydrate (DCPD) and parallel new bone formation, in a drill hole. After 2 weeks almost the entire surface of the cement was in direct contact with the margins of the bone defect. After 4 weeks, occasional granules of β-TCP and the newly formed bone islets are visible. This area expanded after 6 weeks, involving a progressive reabsorp-

*Image of the drill hole with progressive resorption of the calcium phosphate cement matrix, during 8 weeks [81].*

tion of the cement matrix and parallel neoformed formation [81].

The importance given in the use of apatites in bone replacement is due to the fact that this mineral is the base of the main inorganic part of hard tissues. In fact, nonstoichiometric or calcium-deficient hydroxyapatite (CDHA) is the main mineral phase characteristic of human bones [94]. The CPCs consist of a network of calcium phosphate crystals, with chemical composition and crystal size that can be modified to approximate the biological hydroxyapatite that exists in the living bone [95, 96]. In this regard, it is necessary to clarify that even though the stoichiometric hydroxyapatite has a fixed composition, the apatite structure may exist in a variety of compositions. CDHA comprises in its composition the possibility of varying amounts of calcium, where it is possible to present a completely deficient structure based on this base element. The composition may be expressed as Ca10 − *x*(HPO4)(PO4)6 − *x*(OH)2 − *x*, where *x* ranges from 0 to 1, 0 for stoichiometric hydroxyapatite and 1 for hydroxyapatite totally deficient in calcium. Biological apatite is deficient in cal-

Apatite cements may form PHA or CDHA through a precipitation reaction. The synthesis of these cements allows the incorporation of different ions in their composition, depending on the initial compounds. The formation of hydroxyapatite that occurs in the cement is compared to the process of formation of new bone

, K+

, Mg2+, F<sup>−</sup>, and Cl<sup>−</sup> [34].

**2. Principal calcium phosphate cements**

cium containing various ionic substitutions such as Na+

**2.1 Apatite cements**

**Figure 2.**

**122**

and is also seen as a biomimetic process, because it occurs at body temperature and physiological environment. This may explain the fact that the hydroxyapatite formed in the reactions of calcium phosphate cements is much more similar to biological apatites than the ceramic hydroxyapatite resulting from high-temperature sintering processes [34].

The CPCs lead to the formation of PHA or CDHA and can be divided into three systems (single compound, two compounds, more than two compounds), taking into account the number and type of calcium phosphates used in the synthesis [97].

Monocomponent CPCs are those having a single calcium phosphate reagent that hydrolyzes to form PHA or CDHA. Taking into account that at pH 4.2 the hydroxyapatite is less soluble, any other calcium phosphate present will dissolve, and the PHA will tend to precipitate. However, when the formation of PHA occurs, from the hydrolysis of calcium phosphate, the reaction mechanism becomes very slow, due to a decrease in the level of supersaturation, as the reaction proceeds [13]. In this in which only one compound is present, no release of any acid or no base occurs due to the Ca and P ratio being maintained [34, 35].

The second type of cement that may exist is composed of two calcium phosphates, one acid and the other basic, where they adjust after an acid–base reaction. The most commonly used compound is usually tetracalcium phosphate (TTCP), as it is the only calcium phosphate with a Ca/P ratio higher than PHA. Therefore, TTCP can be combined with one or more calcium phosphates with lower Ca/P ratios to obtain PHA or CDHA, avoiding the formation of acids or bases as final products. The combinations that have been more studied seek to produce cements that adjust to the body temperature in a range of pH around the neutral [34].

The third possible system consists of more than two compounds, including calcium phosphates and other salts. For example, a cement proposed by Norian Corporation [5] is used where calcium phosphates with a Ca/P ratio lower than PHA and CaCO3 are added as an additional source of calcium. The initial configuration process involves the formation of DCPD, later forming dahllite, a carbonated hydroxyapatite similar to the bone mineral [5].

Apatitic CPCs appear as a viscous, easily moldable material; however, their injection is difficult. **Figure 3** shows the microstructure of an apatitic cement after setting. The setting time can also be reduced by means of additives such as with the introduction of PHA particles. These changes in the composition may lead to an

#### **Figure 3.**

*Microstructure of an apatitic calcium phosphate cement after setting, showing the micro−/nanosize pore structure formed by the entanglement of the precipitated crystals [71].*

adjustment time in the range of about 15 minutes. When hardening of the cement paste occurs too fast, the hardened cement must be milled to render it viscous again. Subsequently, the paste hardens due to the precipitation of PHA.

After implantation, the mechanical properties can be altered. Investigations indicate that the mechanical properties of apatite CPC tend to increase, unlike brushite cement, which initially decrease and increase when the bone develops [98, 99].

#### **2.2 Brushite cements**

Brushite (DCPD) is an acidic calcium phosphate that has been found in some physiological sites, for example, in bones [100]. Unlike hydroxyapatite, brushite is metastable under physiological conditions [101] and for this reason reabsorbed much faster than CPC apatite; however, there are studies that conclude that DCPD in vivo tends to convert to PHA [26]. Some CPCs were designed to provide brushite as the final product.

Several combinations of compounds have been proposed for the formation of brushite cements; most are β-tricalcium phosphate (β-TCP) and an acid component, namely, monocalcium phosphate monohydrate (MCPM) or phosphoric acid [102–104].

The reaction leading to the formation of brushite CPCs is an acid–base reaction. The brushite paste is acidic during sedimentation because brushite can only precipitate at a pH value below ~6 [105]. The pH of the cement paste tends to change slowly toward equilibrium pH [106]. If the slurry contains an excess of basic phase, the pH tends to equilibrate by crossing the solubility of the base phase with that of the DCPD. The time of stabilization of brushite CPC depends greatly on the solubility of the basic phase: the higher the solubility of the basic phase, the faster the defined time. For example, hydroxyapatite (HA) + MCPM blends have an adjustment time of several minutes. The β-TCP + MCPM mixtures have an adjustment time of 30–60 seconds [107, 108]. However, compounds that inhibit the development of DCPD crystals can be added, increasing the settling time of the β-TCP + MCPM mixtures [109]. The brushite CPC can initially be very liquid and still be defined within a short period of time, unlike the apatite CPC. The brushite CPC is slightly weaker (tensile strength of 10 MPa [110] and compressive strength of 60 MPa) than the apatite CPC (tensile strength of 16 MPa [111] and compressive strength of 83 MPa [112]). The mechanical properties of apatite CPC increase [98], whereas those of brushite CPC decrease [99] This latter phenomenon is attributed to the greater solubility of DCPD in relation to that of PHA [113]. After a few weeks of implantation, the mechanical properties of brushite CPC are promoted by bone growth [99]. Although brushite CPC exhibits biocompatible properties, inflammatory reactions have been reported with the excessive addition of brushite CPC [114]. Investigations indicate that these reactions are due to the transformation of DCPD into PHA [115]. This reaction releases large amounts of acid. The transformation of DCPD into PHA can be avoided with the addition of magnesium ions to the cement [116]. Unlike apatite CPC, brushite CPC cannot be reabsorbed exclusively by osteoclastic activity but also by simple dissolution. Therefore, brushite CPCs degrade at a faster rate than the apatite CPC.

Although brushite demonstrates a higher solubility rate than the other calcium phosphate phases, it is a precursor of the most stable HA phase [117–120]. For this reason, DCPD coatings as an initial step to obtaining HA have been widely used. The synthesis of HA through precipitation mechanisms results in compacted crystals but with sizes difficult to control. Using DCPD as a precursor becomes favorable since it is possible to modify the crystal size of the DCPD through homogeneous precipitation and can be converted directly into HA [119]. In environments with a pH > 6–7, brushite becomes unstable and becomes the most favorable HA phase [121, 122].

**125**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

no inflammation occurring [127].

**3.1 Drug delivery**

**Figure 4.**

reagents as well as their final product.

**3. Calcium phosphate cement applications**

material is injectable and biodegradable [77].

bone tumors or osteoporosis [35].

matrix and the inserted drug.

The fact that DCPD is able to be more soluble leads to its use in metal implants as a means of increasing the amount of calcium and phosphate ions available in the surrounding tissue of the implant to promote increased osseointegration [118].

*mechanism, and microstructure evolution during setting (adapted from [35]).*

*Classification of calcium phosphate cements, with examples of the most common formulations. From top to bottom, the cements are classified by the type of end product (apatite or brushite), a number of components in the solid phase (single or multiple), type of setting reaction (hydrolysis or acid–base reaction), setting* 

The biocompatibility of DCPD as a coating has been demonstrated in several cell lines as, for example, in pre-osteoblastic macrophages [123, 124] and fibroblastic cells [125]. The biocompatibility of DCPD has also been demonstrated when used at a cranial defect site in sheep [126], and the formation of new bone was observed in the absence of inflammation [81]. A clinical study in humans in 2010 effectively used a brushite cement for the repair and increase of pterionic craniotomies, with

**Figure 4** summarizes how the CPCs are classified by type and number of initial

The main requirements for a substrate to have potential as a drug carrier are to have the ability to incorporate it, to retain it at a specific site, and to distribute it progressively over time in surrounding tissues. In addition, it is beneficial that the

Calcium phosphate cements, in addition to allowing hardening at room or body temperature, allow the insertion of various components due to their intrinsic porosity. It is possible to incorporate drugs, biologically active molecules, or even cells without their functions being altered by the effect of temperature or even losing their activity during the procedure (**Figure 5**). This change in the CPCs offers new properties in addition to the osteoconductive characteristic, namely, to increase its capacity for bone regeneration or to support in disorders or pathologies, such as

It is necessary to take into account that the performance of the drug delivery depends on the structural characteristics such as the specific surface area, permeability, matrix degradation rate, drug solubility, or the interaction itself between the *Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*


#### **Figure 4.**

*Contemporary Topics about Phosphorus in Biology and Materials*

**2.2 Brushite cements**

as the final product.

[102–104].

Subsequently, the paste hardens due to the precipitation of PHA.

adjustment time in the range of about 15 minutes. When hardening of the cement paste occurs too fast, the hardened cement must be milled to render it viscous again.

After implantation, the mechanical properties can be altered. Investigations indicate that the mechanical properties of apatite CPC tend to increase, unlike brushite cement, which initially decrease and increase when the bone develops [98, 99].

Brushite (DCPD) is an acidic calcium phosphate that has been found in some physiological sites, for example, in bones [100]. Unlike hydroxyapatite, brushite is metastable under physiological conditions [101] and for this reason reabsorbed much faster than CPC apatite; however, there are studies that conclude that DCPD in vivo tends to convert to PHA [26]. Some CPCs were designed to provide brushite

Several combinations of compounds have been proposed for the formation of brushite cements; most are β-tricalcium phosphate (β-TCP) and an acid component, namely, monocalcium phosphate monohydrate (MCPM) or phosphoric acid

The reaction leading to the formation of brushite CPCs is an acid–base reaction. The brushite paste is acidic during sedimentation because brushite can only precipitate at a pH value below ~6 [105]. The pH of the cement paste tends to change slowly toward equilibrium pH [106]. If the slurry contains an excess of basic phase, the pH tends to equilibrate by crossing the solubility of the base phase with that of the DCPD. The time of stabilization of brushite CPC depends greatly on the solubility of the basic phase: the higher the solubility of the basic phase, the faster the defined time. For example, hydroxyapatite (HA) + MCPM blends have an adjustment time of several minutes. The β-TCP + MCPM mixtures have an adjustment time of 30–60 seconds [107, 108]. However, compounds that inhibit the development of DCPD crystals can be added, increasing the settling time of the β-TCP + MCPM mixtures [109]. The brushite CPC can initially be very liquid and still be defined within a short period of time, unlike the apatite CPC. The brushite CPC is slightly weaker (tensile strength of 10 MPa [110] and compressive strength of 60 MPa) than the apatite CPC (tensile strength of 16 MPa [111] and compressive strength of 83 MPa [112]). The mechanical properties of apatite CPC increase [98], whereas those of brushite CPC decrease [99] This latter phenomenon is attributed to the greater solubility of DCPD in relation to that of PHA [113]. After a few weeks of implantation, the mechanical properties of brushite CPC are promoted by bone growth [99]. Although brushite CPC exhibits biocompatible properties, inflammatory reactions have been reported with the excessive addition of brushite CPC [114]. Investigations indicate that these reactions are due to the transformation of DCPD into PHA [115]. This reaction releases large amounts of acid. The transformation of DCPD into PHA can be avoided with the addition of magnesium ions to the cement [116]. Unlike apatite CPC, brushite CPC cannot be reabsorbed exclusively by osteoclastic activity but also by simple dissolution. Therefore, brushite CPCs degrade at a faster rate than the apatite CPC.

Although brushite demonstrates a higher solubility rate than the other calcium phosphate phases, it is a precursor of the most stable HA phase [117–120]. For this reason, DCPD coatings as an initial step to obtaining HA have been widely used. The synthesis of HA through precipitation mechanisms results in compacted crystals but with sizes difficult to control. Using DCPD as a precursor becomes favorable since it is possible to modify the crystal size of the DCPD through homogeneous precipitation and can be converted directly into HA [119]. In environments with a pH > 6–7, brushite becomes unstable and becomes the most favorable HA phase [121, 122].

**124**

*Classification of calcium phosphate cements, with examples of the most common formulations. From top to bottom, the cements are classified by the type of end product (apatite or brushite), a number of components in the solid phase (single or multiple), type of setting reaction (hydrolysis or acid–base reaction), setting mechanism, and microstructure evolution during setting (adapted from [35]).*

The fact that DCPD is able to be more soluble leads to its use in metal implants as a means of increasing the amount of calcium and phosphate ions available in the surrounding tissue of the implant to promote increased osseointegration [118].

The biocompatibility of DCPD as a coating has been demonstrated in several cell lines as, for example, in pre-osteoblastic macrophages [123, 124] and fibroblastic cells [125]. The biocompatibility of DCPD has also been demonstrated when used at a cranial defect site in sheep [126], and the formation of new bone was observed in the absence of inflammation [81]. A clinical study in humans in 2010 effectively used a brushite cement for the repair and increase of pterionic craniotomies, with no inflammation occurring [127].

**Figure 4** summarizes how the CPCs are classified by type and number of initial reagents as well as their final product.

#### **3. Calcium phosphate cement applications**

#### **3.1 Drug delivery**

The main requirements for a substrate to have potential as a drug carrier are to have the ability to incorporate it, to retain it at a specific site, and to distribute it progressively over time in surrounding tissues. In addition, it is beneficial that the material is injectable and biodegradable [77].

Calcium phosphate cements, in addition to allowing hardening at room or body temperature, allow the insertion of various components due to their intrinsic porosity. It is possible to incorporate drugs, biologically active molecules, or even cells without their functions being altered by the effect of temperature or even losing their activity during the procedure (**Figure 5**). This change in the CPCs offers new properties in addition to the osteoconductive characteristic, namely, to increase its capacity for bone regeneration or to support in disorders or pathologies, such as bone tumors or osteoporosis [35].

It is necessary to take into account that the performance of the drug delivery depends on the structural characteristics such as the specific surface area, permeability, matrix degradation rate, drug solubility, or the interaction itself between the matrix and the inserted drug.

#### **Figure 5.**

*Scheme about the possible applications of calcium phosphates in the biomolecules and drug delivery (adapted from [128]).*

The interaction between the drug and the cement matrix is defined by the drug insertion procedure. Usually, the drugs are added as a powder to the solid phase or dissolved in the liquid part. The method that allows better homogenization of the drug in the matrix is in the liquid phase [35].

Another method for inserting the drug into the matrix is by impregnating solid beads or granules from the cement with a drug solution. This method continues to present benefits compared to other conventional methods even with its impaired injection process. The advantages associated with this method are due to the fact that the consolidation of the material (dissolution-precipitation reaction) leads to the production of hydrated matrices with large specific surface areas that allow the loading of drugs and their release mechanism to be favorable [35].

The two drug encapsulation processes differ greatly in the structural properties of the matrix. In one method the drug is incorporated in the initial phase together with the cement reactants while the matrix is still evolving. Thus, the hardening can last for hours or even days until the suspension of particles evolves into a network of interlaced crystals. In the other case, if the drug is added to the preconceived cement, the matrix structure will always be stable throughout the release process, in addition to possible degradation, and therefore the results cannot be extrapolated to the previous situation. This highlights the need for the studies to be carried out, taking into account the actual conditions of application. The alternative method is to incorporate the drug into polymeric microspheres prior to mixing in the cement. This procedure has two advantages over the other methods presented. It is possible to modify the release kinetics of the drug, and the degradation of the microspheres generates an array capable of being more easily reabsorbed and remodeled [35].

The incorporation of drugs can influence the entire mechanism of action of the cement and compromise its purpose, changing the fit kinetics, rheological characteristics, and microstructural development. For example, molecules that interact with calcium or phosphate ions promote a coprecipitation during the fit or form complexes with Ca2+, promoting a delay in precipitation and modifying viscosity, set time, and cement properties [72, 129, 130].

**127**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

synthesized cements.

**3.2 Growth factor addition**

the bone tissue [77].

material [134, 135].

much slower [137].

may influence the functionality of the drug and its release.

In addition to the influence that the drug has on the cement matrix, it is also necessary to take into account the influence of the cement on the stability of the drug or bioactive molecule. Due to the dissolution-precipitation process, there is a change in the surrounding pH as well as the change in ionic concentrations, which

Thus, it is beneficial to study the release of drugs introduced from already

Growth factors are a large group of proteins that interact at the cellular level [35]. The major families of these proteins are the transforming growth factor-beta superfamily responsible for promoting bone regeneration. The BMPs are part of the TGF-β superfamily and have been widely used in bone regeneration. It is known to play a role as an activating agent in the various biological phenomena responsible for bone formation and therefore can accelerate bone growth. These BMPs stand out from the other growth factors of the large TGF-βSF group because they are osteoinductive. That is, the BMPs act at the level of cell differentiation transforming the pluripotential cells into bone-forming cells, aiding in bone formation outside

These proteins have been produced at an industrial level with a high level of purity; however, it is necessary that their administration is controlled and with adequate therapeutic levels as well as adapted to the tissue targets. In fact, it is known that the injection of such substances alone cannot induce the formation and regeneration of tissues since the protein diffuses very rapidly from the site of implantation. Thus, the CPCs present themselves as good substrates and carriers for these bioactive molecules, also improving their function as osteoconduction [77]. Studies have shown that the superfamily of growth factors stimulates osteoblast proliferation and collagen synthesis in vitro [132] and may increase the size of the

This improvement in bone growth, when applied to cement with these molecules, is due to the adsorption of large doses of rhTGF-β1 on the surface of the

Despite this accumulation of the growth factor to the surface, there is a homogenous distribution throughout the cement mass, increasing the time of the release

In contrast to the kinetics of drug release, the release of these factors becomes

It has been determined that in the first days, the release rate of the components is higher because the initial release is only of the material present in the surface layer that is in contact with the medium. This increase in the rhTGF-β1 release was confirmed when the area in contact with the medium occurred by fragmentation. The same phenomenon was observed when BMP-2 human recombinant microspheres were introduced into the CPC. The release of the factor was quite limited due to

cortical bone when applied near the periosteum in vivo [133].

of the growth factors while the degradation of the matrix occurs.

osteoblastic cells using primary mouse bone cells in vitro [136].

Blom et al. showed that the addition of a human recombinant TGF-β<sup>1</sup> (rhTGF-β1) to a CPC in the adjustment phase stimulated the differentiation of pre-

Recently, studies have been developed related to the incorporation of antibiotics, as a preventive method of infections resulting from surgeries or as treatment of bone infections. In addition to antibiotics, studies with anti-inflammatories, antitumor drugs, or hormones have also been disclosed. In another aspect, the incorporation of factors that stimulate bone regeneration, such as bone morphogenetic proteins (BMP) or transforming growth factors-β (TGF-β), has been studied [77, 131].

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

*Contemporary Topics about Phosphorus in Biology and Materials*

drug in the matrix is in the liquid phase [35].

set time, and cement properties [72, 129, 130].

The interaction between the drug and the cement matrix is defined by the drug insertion procedure. Usually, the drugs are added as a powder to the solid phase or dissolved in the liquid part. The method that allows better homogenization of the

*Scheme about the possible applications of calcium phosphates in the biomolecules and drug delivery (adapted* 

Another method for inserting the drug into the matrix is by impregnating solid beads or granules from the cement with a drug solution. This method continues to present benefits compared to other conventional methods even with its impaired injection process. The advantages associated with this method are due to the fact that the consolidation of the material (dissolution-precipitation reaction) leads to the production of hydrated matrices with large specific surface areas that allow the

The two drug encapsulation processes differ greatly in the structural properties of the matrix. In one method the drug is incorporated in the initial phase together with the cement reactants while the matrix is still evolving. Thus, the hardening can last for hours or even days until the suspension of particles evolves into a network of interlaced crystals. In the other case, if the drug is added to the preconceived cement, the matrix structure will always be stable throughout the release process, in addition to possible degradation, and therefore the results cannot be extrapolated to the previous situation. This highlights the need for the studies to be carried out, taking into account the actual conditions of application. The alternative method is to incorporate the drug into polymeric microspheres prior to mixing in the cement. This procedure has two advantages over the other methods presented. It is possible to modify the release kinetics of the drug, and the degradation of the microspheres generates an array capable of being more easily reabsorbed and remodeled [35]. The incorporation of drugs can influence the entire mechanism of action of the cement and compromise its purpose, changing the fit kinetics, rheological characteristics, and microstructural development. For example, molecules that interact with calcium or phosphate ions promote a coprecipitation during the fit or form complexes with Ca2+, promoting a delay in precipitation and modifying viscosity,

loading of drugs and their release mechanism to be favorable [35].

**126**

**Figure 5.**

*from [128]).*

In addition to the influence that the drug has on the cement matrix, it is also necessary to take into account the influence of the cement on the stability of the drug or bioactive molecule. Due to the dissolution-precipitation process, there is a change in the surrounding pH as well as the change in ionic concentrations, which may influence the functionality of the drug and its release.

Thus, it is beneficial to study the release of drugs introduced from already synthesized cements.

Recently, studies have been developed related to the incorporation of antibiotics, as a preventive method of infections resulting from surgeries or as treatment of bone infections. In addition to antibiotics, studies with anti-inflammatories, antitumor drugs, or hormones have also been disclosed. In another aspect, the incorporation of factors that stimulate bone regeneration, such as bone morphogenetic proteins (BMP) or transforming growth factors-β (TGF-β), has been studied [77, 131].

#### **3.2 Growth factor addition**

Growth factors are a large group of proteins that interact at the cellular level [35]. The major families of these proteins are the transforming growth factor-beta superfamily responsible for promoting bone regeneration. The BMPs are part of the TGF-β superfamily and have been widely used in bone regeneration. It is known to play a role as an activating agent in the various biological phenomena responsible for bone formation and therefore can accelerate bone growth. These BMPs stand out from the other growth factors of the large TGF-βSF group because they are osteoinductive. That is, the BMPs act at the level of cell differentiation transforming the pluripotential cells into bone-forming cells, aiding in bone formation outside the bone tissue [77].

These proteins have been produced at an industrial level with a high level of purity; however, it is necessary that their administration is controlled and with adequate therapeutic levels as well as adapted to the tissue targets. In fact, it is known that the injection of such substances alone cannot induce the formation and regeneration of tissues since the protein diffuses very rapidly from the site of implantation. Thus, the CPCs present themselves as good substrates and carriers for these bioactive molecules, also improving their function as osteoconduction [77].

Studies have shown that the superfamily of growth factors stimulates osteoblast proliferation and collagen synthesis in vitro [132] and may increase the size of the cortical bone when applied near the periosteum in vivo [133].

This improvement in bone growth, when applied to cement with these molecules, is due to the adsorption of large doses of rhTGF-β1 on the surface of the material [134, 135].

Despite this accumulation of the growth factor to the surface, there is a homogenous distribution throughout the cement mass, increasing the time of the release of the growth factors while the degradation of the matrix occurs.

Blom et al. showed that the addition of a human recombinant TGF-β<sup>1</sup> (rhTGF-β1) to a CPC in the adjustment phase stimulated the differentiation of preosteoblastic cells using primary mouse bone cells in vitro [136].

In contrast to the kinetics of drug release, the release of these factors becomes much slower [137].

It has been determined that in the first days, the release rate of the components is higher because the initial release is only of the material present in the surface layer that is in contact with the medium. This increase in the rhTGF-β1 release was confirmed when the area in contact with the medium occurred by fragmentation. The same phenomenon was observed when BMP-2 human recombinant microspheres were introduced into the CPC. The release of the factor was quite limited due to

the possible physical entrapment of the microparticles inside the porous cement. According to the authors, the nanoporosity of CPC not only did not facilitate the release of the protein but could also limit it because of the high binding affinity of the protein by CPC [138].

Haddad et al. [139] investigated the action of implantation of cement loaded with BMP-2 in the bone repair of a critical-sized calvarial vault defect in rabbits. Compared with control, an increase in bone formation was observed at 45% after 12 weeks of implantation.

Other investigations by Seeherman et al. [140, 141] also demonstrated the efficiency of these combined systems (BMP-2/cement). For example, these composites accelerated the filling of a bone defect by 40% after approximately 4 months of implantation, compared to cement without the protein. This study was done in a primate fibula osteotomy [140].

The composite used in the abovementioned study was also used in rabbit bone defects. After 4 weeks of implantation, an acceleration of reabsorption was observed as well as filling of the defect compared to the base cement. This acceleration led to the complete filling of the defect with new bone 8 weeks after the implant.

#### **3.3 Ion addition**

To avoid infections resulting from orthopedic surgery, which usually lead to bone loss or subsequent removal of the implant, alternatives such as antibiotic delivery have been used on the site [142–144]. This transport is usually done using poly(methyl methacrylate) (PMMA) or by encapsulating the drug in the CPC matrix. PMMA beads have the drawback that they are not resorbable and require further surgery to remove them and place new antibiotic-loaded spheres if the goal is to prolong the treatment [145–148]. Faced with this drawback, CPCs have been widely studied as degradable materials capable of carrying antibiotics [77, 145, 148–155]. However, there is a risk of creating bacterial resistance due to low doses of release [156–158]. Thus, the use of surface functionalization of biomaterials as well as the coating of implant surfaces with silver ions has been recurring, conferring antibacterial properties [159–163].

The antimicrobial properties of Ag<sup>+</sup> ions have been investigated and studied in the field of biomedical engineering [164]. It has been found that bacteria hardly gain resistance to silver-based products and low concentrations are required to have a bactericidal effect [165].

Therefore, both metallic silver and ionic silver were incorporated in several biomaterials, HA [166–169], and bioactive glasses [164, 170–172]. In both structures, silver has relatively low toxicity to human cells [173–176].

Ewald et al. evaluated the antimicrobial properties of silver-loaded bone cements as well as their osteoconductive and resorbable properties. The reagents used were both α- and β-tricalcium phosphates combined with slightly acidic compounds to form HA or brushite cement. This study revealed that it is possible to synthesize cements with antimicrobial activity with effects comparable to antibiotic treatments. **Figure 6** shows the inhibition of both *S. aureus* and *S. epidermidis* cultured on the surfaces of the silver-doped cements. Ag-brushite exhibits more antibacterial properties than Ag–HA. In addition, in the case of brushite cements, silver ions allow the cements to increase in compressive strength by approximately 30% [177].

Several studies have been carried out using silver-doped calcium phosphate cements, and the results have been satisfactory, demonstrating inhibitory effect against certain bacteria [178].

In addition to silver, other ions have been incorporated into materials composed of calcium phosphate. Doping with Co2+ showed proangiogenic effects [179, 180].

**129**

and differentiation.

**Figure 6.**

*LB medium [177].*

evaluating material remodeling in vivo [183].

material for application in bone regeneration.

**3.4 3D printing of CPCs**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

Modification with Cu2+ increased the rate of vascularization [181]. The influence of the introduction of Co2+, Cu2+, and Cr3+ on calcium phosphate cement was investigated, evaluating the material properties, proliferation, and osteogenic differentiation of human mesenchymal stem cells in vitro [182]. The Cr3+ and Cu2+ ions, in this case with less evidence, had positive effects on osteogenic proliferation

*Bacterial activity of* S. aureus *and* S. epidermidis *on cement surfaces (n = 4) determined by the WST-1-test in* 

Despite all the positive results of this incorporation, the examinations that evaluate the osteogenic capacity in vitro are not enough to estimate the clinical performance of these materials in the bone graft. Since the balance between bone neoformation and material resorption is crucial for successful remodeling, in vitro analysis of osteoclast-mediated degradation of materials is a logical next step in

However, it is necessary to take into account the dose that is used in the doping process, as investigators have concluded that doses in certain amounts of Cu2+ and Co2+ can cause cytotoxic reactions to osteoclasts and progenitors of osteoclasts during the initial release of Cu2+ and Co2+, respectively. Cements at lower doses are beneficial to bone regeneration since Cu2+ at 18 μM completely inhibits reabsorption (but not the formation of osteoclasts), which may be beneficial for patients with osteoporosis and imbalance between bone formation and resorption. The addition of Cr3+ to brushite cements increases osteoclastic reabsorption and increases the viability of osteoprogenitor cells compared to cement without the addition of ions [182]. Therefore, Cr3+-doped brushite cements are suggested as a promising new

The 3D printing technique, or additive manufacture (AM), is based on the addition of layers of powder material producing solid materials with adjustable porosity, through a digital geometric model. This method has been the subject of much research in the medical field such as in the synthesis of the customized scaffold [184]. In addition to the economical and fast production that 3D printing allows, it also allows

#### **Figure 6.**

*Contemporary Topics about Phosphorus in Biology and Materials*

the protein by CPC [138].

12 weeks of implantation.

**3.3 Ion addition**

primate fibula osteotomy [140].

antibacterial properties [159–163].

a bactericidal effect [165].

against certain bacteria [178].

The antimicrobial properties of Ag<sup>+</sup>

silver has relatively low toxicity to human cells [173–176].

the possible physical entrapment of the microparticles inside the porous cement. According to the authors, the nanoporosity of CPC not only did not facilitate the release of the protein but could also limit it because of the high binding affinity of

Haddad et al. [139] investigated the action of implantation of cement loaded with BMP-2 in the bone repair of a critical-sized calvarial vault defect in rabbits. Compared with control, an increase in bone formation was observed at 45% after

Other investigations by Seeherman et al. [140, 141] also demonstrated the efficiency of these combined systems (BMP-2/cement). For example, these composites accelerated the filling of a bone defect by 40% after approximately 4 months of implantation, compared to cement without the protein. This study was done in a

The composite used in the abovementioned study was also used in rabbit bone defects. After 4 weeks of implantation, an acceleration of reabsorption was observed as well as filling of the defect compared to the base cement. This acceleration led to

To avoid infections resulting from orthopedic surgery, which usually lead to bone loss or subsequent removal of the implant, alternatives such as antibiotic delivery have been used on the site [142–144]. This transport is usually done using poly(methyl methacrylate) (PMMA) or by encapsulating the drug in the CPC matrix. PMMA beads have the drawback that they are not resorbable and require further surgery to remove them and place new antibiotic-loaded spheres if the goal is to prolong the treatment [145–148]. Faced with this drawback, CPCs have been widely studied as degradable materials capable of carrying antibiotics [77, 145, 148–155]. However, there is a risk of creating bacterial resistance due to low doses of release [156–158]. Thus, the use of surface functionalization of biomaterials as well as the coating of implant surfaces with silver ions has been recurring, conferring

the field of biomedical engineering [164]. It has been found that bacteria hardly gain resistance to silver-based products and low concentrations are required to have

Therefore, both metallic silver and ionic silver were incorporated in several biomaterials, HA [166–169], and bioactive glasses [164, 170–172]. In both structures,

Ewald et al. evaluated the antimicrobial properties of silver-loaded bone cements as well as their osteoconductive and resorbable properties. The reagents used were both α- and β-tricalcium phosphates combined with slightly acidic compounds to form HA or brushite cement. This study revealed that it is possible to synthesize cements with antimicrobial activity with effects comparable to antibiotic treatments. **Figure 6** shows the inhibition of both *S. aureus* and *S. epidermidis* cultured on the surfaces of the silver-doped cements. Ag-brushite exhibits more antibacterial properties than Ag–HA. In addition, in the case of brushite cements, silver ions allow the cements to increase in compressive strength by approximately 30% [177]. Several studies have been carried out using silver-doped calcium phosphate cements, and the results have been satisfactory, demonstrating inhibitory effect

In addition to silver, other ions have been incorporated into materials composed of calcium phosphate. Doping with Co2+ showed proangiogenic effects [179, 180].

ions have been investigated and studied in

the complete filling of the defect with new bone 8 weeks after the implant.

**128**

*Bacterial activity of* S. aureus *and* S. epidermidis *on cement surfaces (n = 4) determined by the WST-1-test in LB medium [177].*

Modification with Cu2+ increased the rate of vascularization [181]. The influence of the introduction of Co2+, Cu2+, and Cr3+ on calcium phosphate cement was investigated, evaluating the material properties, proliferation, and osteogenic differentiation of human mesenchymal stem cells in vitro [182]. The Cr3+ and Cu2+ ions, in this case with less evidence, had positive effects on osteogenic proliferation and differentiation.

Despite all the positive results of this incorporation, the examinations that evaluate the osteogenic capacity in vitro are not enough to estimate the clinical performance of these materials in the bone graft. Since the balance between bone neoformation and material resorption is crucial for successful remodeling, in vitro analysis of osteoclast-mediated degradation of materials is a logical next step in evaluating material remodeling in vivo [183].

However, it is necessary to take into account the dose that is used in the doping process, as investigators have concluded that doses in certain amounts of Cu2+ and Co2+ can cause cytotoxic reactions to osteoclasts and progenitors of osteoclasts during the initial release of Cu2+ and Co2+, respectively. Cements at lower doses are beneficial to bone regeneration since Cu2+ at 18 μM completely inhibits reabsorption (but not the formation of osteoclasts), which may be beneficial for patients with osteoporosis and imbalance between bone formation and resorption. The addition of Cr3+ to brushite cements increases osteoclastic reabsorption and increases the viability of osteoprogenitor cells compared to cement without the addition of ions [182]. Therefore, Cr3+-doped brushite cements are suggested as a promising new material for application in bone regeneration.

#### **3.4 3D printing of CPCs**

The 3D printing technique, or additive manufacture (AM), is based on the addition of layers of powder material producing solid materials with adjustable porosity, through a digital geometric model. This method has been the subject of much research in the medical field such as in the synthesis of the customized scaffold [184]. In addition to the economical and fast production that 3D printing allows, it also allows

#### **Figure 7.**

*Scheme of the 3D printing system of the calcium phosphate bone cement scaffolds and its advantages due to the possibility of improving the settings.*

the manufacture of pieces with high geometric complexity. All these advantages have made 3D printing very useful in the field of biomedical and tissue engineering due to the ability to replicate the complicated architecture as well as the cellular heterogeneity present in tissues and organs. The bone presents itself as an exquisite structure due to the existence of a complex compound of minerals and an organic matrix. Taking into account this organization, there are several size scales, and it is possible to easily reproduce this structure through 3D printing. In addition, 3D printing is suitable for producing structures derived from medical images, such as CT scans [185, 186].

Recently, the manufacture of the customized scaffolds of the calcium phosphate cements by 3D printing has been described. During printing, highly viscous or pasty materials are dispensed through an adjustable dosing nozzle, resulting in the deposition of CPC layers on a platform with a liquid or air as a plotting medium. After the stabilization of the scaffolds, the water adjustment reaction begins. These stabilization and hardening conditions allow the integration of biological molecules at specific sites [187].

In this way, it is possible to create scaffolds, based on calcium phosphates, more complex with specificities that can promote higher bioactivity or introduce drugs or other therapeutic components, taking into account the patients' needs (**Figure 7**).

In addition, this technique allows the combination of components (i.e., CPC and alginate) to produce structures more resistant to compression and with improved toughness compared to pure CPC supports [188].

#### **4. Conclusions and future perspective**

Due to its bioactivity, biocompatibility, osteoconductivity, and osteoinductivity, calcium phosphate cements present an advantageous option in the field of bone tissue engineering, taking into account all the needs that this application demands. In addition, it may be used as scaffolds and transport medium for various biological molecules such as stem cells, drugs, or growth factors. It is also worth mentioning the possibility of producing structures of these cements through 3D printing technology, where it is possible to manufacture intrinsically complex biomimetic structures due to the degree of precision of this technique.

The possibility of building calcium phosphate cements, involving the incorporation of several types of cells, growth factors, molecules, or bioactive glasses, allows favorable results in the vascularization of bone tissues and, consequently, in bone regeneration. This feature, particularly appreciated in large bone regenerations, will allow a considerable increase in the use of these structures in clinical applications. However, more research is needed to consolidate and understand all the information associated with the fundamental mechanisms that promote the development of tissue engineering and regenerative medicine.

**131**

**Author details**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

AM manufacture additive

HA hydroxyapatite L/P ratio liquid-powder ratio

BMP bone morphogenetic proteins CDHA calcium-deficient hydroxyapatite CPCs calcium phosphate cements DCPA dicalcium phosphate anhydrous DCPD dicalcium phosphate dihydrate

PHA precipitated hydroxyapatite PMMA poly(methyl methacrylate) TCP tricalcium phosphate

TTCP tetracalcium phosphate rhTGF-β<sup>1</sup> human recombinant TGF-β<sup>1</sup>

TGF-β transforming growth factors-beta

MCPM monocalcium phosphate monohydrate

Manuel Pedro Fernandes Graça\* and Sílvia Rodrigues Gavinho I3N and Physics Department, University of Aveiro, Aveiro, Portugal

© 2020 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium,

\*Address all correspondence to: mpfg@ua.pt

provided the original work is properly cited.

investigated.

**Acronyms**

The success associated with biomaterials has been underestimated maybe because they have been used clinically for more than 40 years. However, there is still a huge diversity of calcium phosphate-based materials that have not been fully *Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

The success associated with biomaterials has been underestimated maybe because they have been used clinically for more than 40 years. However, there is still a huge diversity of calcium phosphate-based materials that have not been fully investigated.

#### **Acronyms**

*Contemporary Topics about Phosphorus in Biology and Materials*

the manufacture of pieces with high geometric complexity. All these advantages have made 3D printing very useful in the field of biomedical and tissue engineering due to the ability to replicate the complicated architecture as well as the cellular heterogeneity present in tissues and organs. The bone presents itself as an exquisite structure due to the existence of a complex compound of minerals and an organic matrix. Taking into account this organization, there are several size scales, and it is possible to easily reproduce this structure through 3D printing. In addition, 3D printing is suitable for producing structures derived from medical images, such as CT scans [185, 186].

*Scheme of the 3D printing system of the calcium phosphate bone cement scaffolds and its advantages due to the* 

Recently, the manufacture of the customized scaffolds of the calcium phosphate

In this way, it is possible to create scaffolds, based on calcium phosphates, more complex with specificities that can promote higher bioactivity or introduce drugs or other therapeutic components, taking into account the patients' needs (**Figure 7**). In addition, this technique allows the combination of components (i.e., CPC and alginate) to produce structures more resistant to compression and with improved

Due to its bioactivity, biocompatibility, osteoconductivity, and osteoinductivity, calcium phosphate cements present an advantageous option in the field of bone tissue engineering, taking into account all the needs that this application demands. In addition, it may be used as scaffolds and transport medium for various biological molecules such as stem cells, drugs, or growth factors. It is also worth mentioning the possibility of producing structures of these cements through 3D printing technology, where it is possible to manufacture intrinsically complex biomimetic

The possibility of building calcium phosphate cements, involving the incorporation of several types of cells, growth factors, molecules, or bioactive glasses, allows favorable results in the vascularization of bone tissues and, consequently, in bone regeneration. This feature, particularly appreciated in large bone regenerations, will allow a considerable increase in the use of these structures in clinical applications. However, more research is needed to consolidate and understand all the information associated with the fundamental mechanisms that promote the development of

cements by 3D printing has been described. During printing, highly viscous or pasty materials are dispensed through an adjustable dosing nozzle, resulting in the deposition of CPC layers on a platform with a liquid or air as a plotting medium. After the stabilization of the scaffolds, the water adjustment reaction begins. These stabilization and hardening conditions allow the integration of biological molecules

**130**

at specific sites [187].

**Figure 7.**

*possibility of improving the settings.*

toughness compared to pure CPC supports [188].

structures due to the degree of precision of this technique.

**4. Conclusions and future perspective**

tissue engineering and regenerative medicine.


#### **Author details**

Manuel Pedro Fernandes Graça\* and Sílvia Rodrigues Gavinho I3N and Physics Department, University of Aveiro, Aveiro, Portugal

\*Address all correspondence to: mpfg@ua.pt

© 2020 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

#### **References**

[1] Brown WE, Chow LC. A new calcium phosphate setting cement. Journal of Dental Research. 1983;**62**:672

[2] LeGeros RZ, Chohayeb A, Shulman A. Apatitic calcium phosphates: Possible dental restorative materials. Journal of Dental Research. 1982;**61**:343

[3] Friedman CD, Costantino PD, Takagi S, Chow LC. BoneSource hydroxyapatite cement: A novel biomaterial for craniofacial skeletal tissue engineering and reconstruction. Journal of Biomedical Materials Research. 1998;**43**:428-432

[4] Kamerer DB, Hirsch BE, Snyderman CH, Costantino P, Friedman CD. Hydroxyapatite cement: A new method for achieving watertight closure in transtemporal surgery. The American Journal of Otology. 1994;**15**:47-49

[5] Constantz BR, Ison IC, Fulmer MT, Poser RD, Smith ST, VanWagoner M, et al. Skeletal repair by in in situ formation of the mineral phase of bone. Science. 1995;**267**:1796-1799

[6] Horstmann WG, Verheyen CCPM, Leemans R. An injectable calcium phosphate cement as a bone-graft substitute in the treatment of displaced lateral tibial plateau fractures. Injury. 2003;**34**:141-144

[7] Strauss EJ, Egol KA. The management of ankle fractures in the elderly. Injury. 2007;**38**:S2-S9

[8] Liverneaux PA. Osteoporotic distal radius curettage—Filling with an injectable calcium phosphate cement. A cadaveric study. European Journal of Orthopaedic Surgery and Traumatology. 2004;**15**:1-6

[9] Welch RD, Zhang H, Bronson DG. Experimental tibial plateau fractures

augmented with calcium phosphate cement or autologous bone graft. The Journal of Bone and Joint Surgery. 2003;**85**:222

[10] Aral A, Yalçin S, Karabuda ZC, Anil A, Jansen JA, Mutlu Z. Injectable calcium phosphate cement as a graft material for maxillary sinus augmentation: An experimental pilot study. Clinical Oral Implants Research. 2008;**19**:612-617

[11] Bai B, Jazrawi LM, Kummer FJ, Spivak JM. The use of an injectable, biodegradable calcium phosphate bone substitute for the prophylactic augmentation of osteoporotic vertebrae and the management of vertebral compression fractures. Spine. 1999;**24**:1521-1526

[12] Schildhauer T, Bennett A, Wright T, Lane J, O'Leary P. Intravertebral body reconstruction with an injectable in situ-setting carbonated apatite: Biomechanical evaluation of a minimally invasive technique. Journal of Orthopaedic Research. 1999;**17**:67-72

[13] Maestretti G, Cremer C, Otten P, Jakob RP. Prospective study of standalone balloon kyphoplasty with calcium phosphate cement augmentation in traumatic fractures. European Spine Journal. 2007;**16**:601-610

[14] Libicher M, Hillmeier J, Liegibel U, Sommer U, Pyerin W, Vetter M, et al. Osseous integration of calcium phosphate in osteoporotic vertebral fractures after kyphoplasty: Initial results from a clinical and experimental pilot study. Osteoporosis International. 2006;**17**:1208-1215

[15] Mermelstein LE, Chow LC, Friedman CD, Crisco JJ. The reinforcement of cancellous bone screws with calcium phosphate cement. Journal of Orthopaedic Trauma. 1996;**10**:15-20

**133**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

Oxford, UK: Elsevier Science Ltd; 2000.

[24] Burstein AH, Reilly DT, Martens M. Aging of bone tissue: Mechanical properties. The Journal of Bone and Joint Surgery. 1976;**58A**:82-86

[25] Carter DR, Hayes WC. The compressive behavior of bone as a two-phase porous structure. Clinical Orthopaedics and Related Research.

[26] Constantz BR, Barr BM, Ison IC, Fulmer MT, Baker J, Mckinney LA, et al. Histological, chemical, and

[27] Ginebra MP, Fernández E, De Maeyer EAP, Verbeeck RMH, Boltong MG, Ginebra J, et al. Setting reaction and hardening of an apatitic calcium phosphate cement. Journal of Dental Research. 1997;**76**(4):905-912

[28] Ginebra MP, Driessens FCM, Planell JA. Effect of the particle size on the micro and nanostructural

A kinetic analysis. Biomaterials.

[29] Driessens FCM, Planell JA, Boltong MG, Khairoun I, Ginebra MP. Osteotransductive bone cements. Proceedings of the Institution of Mechanical Engineers, Part H: Journal of Engineering in Medicine.

[30] Ishikawa K, Miyamoto Y, Kon M, Nagayama M, Asaoka K. Non-decay type fast-setting calcium phosphate cement: composite with sodium alginate. Biomaterials. 1995;**16**:527-532

[31] Khairoun I, Boltong MG, Driessens FC, Planell JA. Effect of

2004;**25**:3453-3462

1998;**212**:427-435

features of a calcium phosphate cement:

crystallographic analysis of four calcium phosphate cements in different rabbit osseous sites. Journal of Biomedical Materials Research Part B: Applied Biomaterials. 1998;**43**:451-461

1977;**59A**:954-962

pp. 31-71

[16] Ooms E, Wolke J, Van der Waerden J, Jansen J. Use of injectable calcium-phosphate cement for the fixation of titanium implants: An experimental study in goats. Journal of Biomedical Materials Research. Part B, Applied Biomaterials. 2003;**66**:447-456

[17] Takemasa R, Kiyasu K, Tani T, Inoue S. Validity of calcium phosphate cement vertebroplasty for vertebral non-union after osteoporotic fracture with middle column involvement. The

Spine Journal. 2007;**7**:148S

Science. 2003;**8**:192-197

2006;**76**:456-468

2006;**27**:2171-2177

pp. 271-308

2011;**4**:1658-1671

for use in vertebroplasty and

Calcium phosphate cements: Competitive drug carriers for the musculoskeletal system? Biomaterials.

[19] Lewis G. Injectable bone cements

kyphoplasty: State-of-the-art review. Journal of Biomedical Materials

Research. Part B, Applied Biomaterials.

[20] Ginebra MP, Traykova T, Planell JA.

[21] Ginebra MP. Cements as bone repair materials. In: Planell JA, editor. Bone Repair Biomaterials. Cambridge, UK: Woodhead Publishing Limited; 2009.

[22] Canal C, Ginebra MP. Fibrereinforced calcium phosphate cements: A review. The Journal of the Mechanical Behavior of Biomedical Materials.

[23] Ontañón M, Aparicio C,

Ginebra MP, Planell JA. Structure and mechanical properties of bone. In: Elices M, editor. Structural Biological Materials, Pergamon Material Series.

[18] Tomita S, Kin A, Yazu M, Abe M. Biomechanical evaluation of kyphoplasty and vertebroplasty with calcium phosphate cement in a simulated osteoporotic compression fracture. Journal of Orthopaedic

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

[16] Ooms E, Wolke J, Van der Waerden J, Jansen J. Use of injectable calcium-phosphate cement for the fixation of titanium implants: An experimental study in goats. Journal of Biomedical Materials Research. Part B, Applied Biomaterials. 2003;**66**:447-456

[17] Takemasa R, Kiyasu K, Tani T, Inoue S. Validity of calcium phosphate cement vertebroplasty for vertebral non-union after osteoporotic fracture with middle column involvement. The Spine Journal. 2007;**7**:148S

[18] Tomita S, Kin A, Yazu M, Abe M. Biomechanical evaluation of kyphoplasty and vertebroplasty with calcium phosphate cement in a simulated osteoporotic compression fracture. Journal of Orthopaedic Science. 2003;**8**:192-197

[19] Lewis G. Injectable bone cements for use in vertebroplasty and kyphoplasty: State-of-the-art review. Journal of Biomedical Materials Research. Part B, Applied Biomaterials. 2006;**76**:456-468

[20] Ginebra MP, Traykova T, Planell JA. Calcium phosphate cements: Competitive drug carriers for the musculoskeletal system? Biomaterials. 2006;**27**:2171-2177

[21] Ginebra MP. Cements as bone repair materials. In: Planell JA, editor. Bone Repair Biomaterials. Cambridge, UK: Woodhead Publishing Limited; 2009. pp. 271-308

[22] Canal C, Ginebra MP. Fibrereinforced calcium phosphate cements: A review. The Journal of the Mechanical Behavior of Biomedical Materials. 2011;**4**:1658-1671

[23] Ontañón M, Aparicio C, Ginebra MP, Planell JA. Structure and mechanical properties of bone. In: Elices M, editor. Structural Biological Materials, Pergamon Material Series.

Oxford, UK: Elsevier Science Ltd; 2000. pp. 31-71

[24] Burstein AH, Reilly DT, Martens M. Aging of bone tissue: Mechanical properties. The Journal of Bone and Joint Surgery. 1976;**58A**:82-86

[25] Carter DR, Hayes WC. The compressive behavior of bone as a two-phase porous structure. Clinical Orthopaedics and Related Research. 1977;**59A**:954-962

[26] Constantz BR, Barr BM, Ison IC, Fulmer MT, Baker J, Mckinney LA, et al. Histological, chemical, and crystallographic analysis of four calcium phosphate cements in different rabbit osseous sites. Journal of Biomedical Materials Research Part B: Applied Biomaterials. 1998;**43**:451-461

[27] Ginebra MP, Fernández E, De Maeyer EAP, Verbeeck RMH, Boltong MG, Ginebra J, et al. Setting reaction and hardening of an apatitic calcium phosphate cement. Journal of Dental Research. 1997;**76**(4):905-912

[28] Ginebra MP, Driessens FCM, Planell JA. Effect of the particle size on the micro and nanostructural features of a calcium phosphate cement: A kinetic analysis. Biomaterials. 2004;**25**:3453-3462

[29] Driessens FCM, Planell JA, Boltong MG, Khairoun I, Ginebra MP. Osteotransductive bone cements. Proceedings of the Institution of Mechanical Engineers, Part H: Journal of Engineering in Medicine. 1998;**212**:427-435

[30] Ishikawa K, Miyamoto Y, Kon M, Nagayama M, Asaoka K. Non-decay type fast-setting calcium phosphate cement: composite with sodium alginate. Biomaterials. 1995;**16**:527-532

[31] Khairoun I, Boltong MG, Driessens FC, Planell JA. Effect of

**132**

2004;**15**:1-6

*Contemporary Topics about Phosphorus in Biology and Materials*

augmented with calcium phosphate cement or autologous bone graft. The Journal of Bone and Joint Surgery.

[10] Aral A, Yalçin S, Karabuda ZC, Anil A, Jansen JA, Mutlu Z. Injectable calcium phosphate cement as a graft material for maxillary sinus augmentation: An experimental pilot study. Clinical Oral Implants Research.

[11] Bai B, Jazrawi LM, Kummer FJ, Spivak JM. The use of an injectable, biodegradable calcium phosphate bone substitute for the prophylactic augmentation of osteoporotic vertebrae and the management of vertebral compression fractures. Spine.

[12] Schildhauer T, Bennett A, Wright T, Lane J, O'Leary P. Intravertebral body reconstruction with an injectable in situ-setting carbonated apatite: Biomechanical evaluation of a

minimally invasive technique. Journal of Orthopaedic Research. 1999;**17**:67-72

Otten P, Jakob RP. Prospective study of standalone balloon kyphoplasty with calcium phosphate cement augmentation in traumatic fractures. European Spine

[14] Libicher M, Hillmeier J, Liegibel U, Sommer U, Pyerin W, Vetter M, et al. Osseous integration of calcium phosphate in osteoporotic vertebral fractures after kyphoplasty: Initial results from a clinical and experimental pilot study. Osteoporosis International.

[13] Maestretti G, Cremer C,

Journal. 2007;**16**:601-610

2006;**17**:1208-1215

[15] Mermelstein LE, Chow LC, Friedman CD, Crisco JJ. The reinforcement of cancellous bone screws with calcium phosphate cement. Journal of Orthopaedic Trauma. 1996;**10**:15-20

2003;**85**:222

2008;**19**:612-617

1999;**24**:1521-1526

[1] Brown WE, Chow LC. A new calcium phosphate setting cement. Journal of

Shulman A. Apatitic calcium phosphates: Possible dental restorative materials. Journal of Dental Research. 1982;**61**:343

Dental Research. 1983;**62**:672

[2] LeGeros RZ, Chohayeb A,

[3] Friedman CD, Costantino PD, Takagi S, Chow LC. BoneSource hydroxyapatite cement: A novel biomaterial for craniofacial skeletal tissue engineering and reconstruction. Journal of Biomedical Materials Research. 1998;**43**:428-432

[4] Kamerer DB, Hirsch BE, Snyderman CH, Costantino P,

1994;**15**:47-49

2003;**34**:141-144

Friedman CD. Hydroxyapatite cement: A new method for achieving watertight closure in transtemporal surgery. The American Journal of Otology.

[5] Constantz BR, Ison IC, Fulmer MT, Poser RD, Smith ST, VanWagoner M, et al. Skeletal repair by in in situ

formation of the mineral phase of bone.

[6] Horstmann WG, Verheyen CCPM, Leemans R. An injectable calcium phosphate cement as a bone-graft substitute in the treatment of displaced lateral tibial plateau fractures. Injury.

management of ankle fractures in the

[8] Liverneaux PA. Osteoporotic distal radius curettage—Filling with an injectable calcium phosphate cement. A cadaveric study. European Journal of Orthopaedic Surgery and Traumatology.

[9] Welch RD, Zhang H, Bronson DG. Experimental tibial plateau fractures

Science. 1995;**267**:1796-1799

[7] Strauss EJ, Egol KA. The

elderly. Injury. 2007;**38**:S2-S9

**References**

calcium carbonate on clinical compliance of apatitic calcium phosphate bone cement. Journal of Biomedical Materials Research. 1997;**38**(4):356-360

[32] Khairoun I, Driessens FCM, Boltong MG, Planell JA, Wenz R. Addition of cohesion promotors to calcium phosphate cements. Biomaterials. 1999;**20**:393-398

[33] Grover LM, Gbureck U, Wright AJ, Tremayne M, Barralet JE. Biologically mediated resorption of brushite cement in vitro. Biomaterials. 2006;**27**:2178-2185

[34] Ginebra MP. Calcium phosphate bone cements. In: Deb S, editor. Ortopedic Bone Cements. Boca Raton, FL: CRC Press; 2008. pp. 206-230

[35] Ginebra MP, Canal C, Espanol M, Pastorino D, Edgar B. Calcium phosphate cements as drug delivery materials. Advanced Drug Delivery Reviews. 2012;**64**:1090-1110

[36] Zhang J, Liu W, Schnitzler V, Tancret F, Bouler JM. Calcium phosphate cements for bone substitution: Chemistry, handling and mechanical properties. Acta Biomaterialia. 2014;**10**: 1035-1049

[37] Zhang JT, Tancret F, Bouler JM. Fabrication and mechanical properties of calcium phosphate cements (CPC) for bone substitution. Materials Science & Engineering. C, Materials for Biological Applications. 2011;**31**:740-747

[38] Guo H, Su JC, Wei J, Kong H, Liu CS. Biocompatibility and osteogenicity of degradable Ca-deficient hydroxyapatite scaffolds from calcium phosphate cement for bone tissue engineering. Acta Biomaterialia. 2009;**5**:268-278

[39] Barralet JE, Grover L, Gaunt T, Wright AJ, Gibson IR. Preparation of macroporous calcium phosphate cement tissue engineering scaffold. Biomaterials. 2002;**23**:3063-3072

[40] Takagi S, Chow LC. Formation of macropores in calcium phosphate cement implants. The Journal of Materials Science: Materials in Medicine. 2001;**12**:135-139

[41] Li M, Liu XY, Liu XD, Ge BF, Chen KM. Creation of macroporous calcium phosphate cements as bone substitutes by using genipin-crosslinked gelatin microspheres. The Journal of Materials Science: Materials in Medicine. 2009;**20**:925-934

[42] Xu H, Quinn JB, Takagi S, Chow LC, Eichmiller FC. Strong and macroporous calcium phosphate cement: Effects of porosity and fiber reinforcement on mechanical properties. Journal of Biomedical Materials Research. 2001;**57**:457-466

[43] Cama G, Barberis F, Botter R, Cirillo P, Capurro M, Quarto R, et al. Preparation and properties of macroporous brushite bone cements. Acta Biomaterialia. 2009;**5**:2161-2168

[44] Félix Lanao RP, Leeuwenburgh SCG, Wolke JGC, Jansen JA. In vitro degradation rate of apatitic calcium phosphate cement with incorporated PLGA microspheres. Acta Biomaterialia. 2011;**7**:3459-3468

[45] Habraken WJEM, Liao HB, Zhang Z, Wolke JGC, Grijpma DW, Mikos AG, et al. In vivo degradation of calcium phosphate cement incorporated into biodegradable microspheres. Acta Biomaterialia. 2010;**6**:2200-2211

[46] Klijn RJ, van den Beucken J, Lanao R, Veldhuis G, Leeuwenburgh SC, Wolke J, et al. Three different strategies to obtain porous calcium phosphate cements: Comparison of performance in a rat skull bone augmentation model. Tissue Engineering. Part A. 2012;**18**: 1171-1182

[47] Qi XP, Ye JD. Mechanical and rheological properties and injectability

**135**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

phosphate cement. Biomaterials.

[55] TenHuisen KS, Brown PW. Formation of calcium-deficient hydroxyapatite from α-tricalcium phosphate. Biomaterials. 1998;**19**:

[56] Bermudez O, Boltong MG,

Driessens FCM, Planell JA. Development of some calcium-phosphate cements from combinations of α-TCP, MCPM and CaO. The Journal of Materials Science: Materials in Medicine. 1994;**5**:160-163

[57] Yang QZ, Troczynski T, Liu DM. Influence of apatite seeds on the

Biomaterials. 2002;**23**:2751-2760

2012;**8**:474-487

[58] Tamimi F, Sheikh Z, Barralet J. Dicalcium phosphate cements: Brushite and monetite. Acta Biomaterialia.

[59] Sarda S, Fernandez E, Nilsson M, Balcells M, Planell JA. Kinetic study of citric acid influence on calcium phosphate bone cements as waterreducing agent. Journal of Biomedical Materials Research. 2002;**61**:653-659

[60] Gbureck U, Barralet JE, Spatz K, Grover LM, Thull R. Ionic modification of calcium phosphate cement viscosity. Part I: Hypodermic injection and strength improvement of apatite cement. Biomaterials.

[61] Barralet JE, Hofmann M,

Grover LM, Gbureck U. High-strength apatitic cement by modification with α-hydroxy acid salts. Advanced Materials. 2003;**15**:2091-2094

[62] Barralet JE, Grover LM, Gbureck U. Ionic modification of calcium phosphate cement viscosity. Part II: Hypodermic injection and strength improvement of brushite cement. Biomaterials.

2004;**25**:2187-2195

2004;**25**:2197-2203

synthesis of calcium phosphate cement.

2003;**24**:4103-4113

2209-2217

containing poly (lactic-co-glycolic acid) microspheres. Materials Science & Engineering. C, Materials for Biological Applications. 2009;**29**:1901-1906

[48] Almirall A, Larrecq G, Delgado JA, Martinez S, Planell JA, Ginebra MP. Fabrication of low temperature macroporous hydroxyapatite scaffolds by foaming and hydrolysis of an a-TCP paste. Biomaterials. 2004;**25**:3671-3680

[49] Del Real RP, Wolke J, Vallet-Regi M, Jansen JA. A new method to produce macropores in calcium phosphate cements. Biomaterials. 2002;**23**:

[50] Chen WC, Zhou HZ, Tang MH, Weir MD, Bao CY, Xu H. Gas-foaming calcium phosphate cement scaffold encapsulating human umbilical cord stem cells. Tissue Engineering. Part A.

[51] Hesaraki S, Moztarzadeh F,

macropores in apatitic calcium phosphate bone cement with the use of an effervescent additive. Journal of Biomedical Materials Research. Part A.

[52] Ginebra M, Delgado J, Harr I, Almirall A, Del Valle S, Planell JA. Factors affecting the structure and properties of an injectable self-setting calcium phosphate foam. Journal of Biomedical Materials Research. Part A.

[53] Bai F, Meng GL, Yuan YA, Liu CS, Wang Z, Liu JA. Role of macropore size in the mechanical properties and in vitro degradation of porous calcium phosphate cements. Materials Letters.

Sharifi D. Formation of interconnected

3673-3680

2012;**18**:816-827

2007;**83A**:80-87

2007;**80A**:351-361

2010;**64**:2028-2031

[54] Liu CS, Shao HF, Chen FY, Zheng HY. Effects of the granularity of raw materials on the hydration and hardening process of calcium

of calcium phosphate cement

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

*Contemporary Topics about Phosphorus in Biology and Materials*

[40] Takagi S, Chow LC. Formation of macropores in calcium phosphate cement implants. The Journal of Materials Science: Materials in Medicine. 2001;**12**:135-139

[41] Li M, Liu XY, Liu XD, Ge BF, Chen KM. Creation of macroporous calcium phosphate cements as bone substitutes by using genipin-crosslinked gelatin microspheres. The Journal of Materials Science: Materials in Medicine. 2009;**20**:925-934

[42] Xu H, Quinn JB, Takagi S, Chow LC, Eichmiller FC. Strong and macroporous calcium phosphate cement: Effects of porosity and fiber reinforcement on mechanical properties. Journal of Biomedical Materials Research.

[43] Cama G, Barberis F, Botter R, Cirillo P, Capurro M, Quarto R, et al. Preparation and properties of macroporous brushite bone cements. Acta Biomaterialia. 2009;**5**:2161-2168

Wolke JGC, Jansen JA. In vitro degradation rate of apatitic calcium phosphate cement with incorporated PLGA microspheres. Acta Biomaterialia.

[45] Habraken WJEM, Liao HB, Zhang Z, Wolke JGC, Grijpma DW, Mikos AG, et al. In vivo degradation of calcium phosphate cement incorporated into biodegradable microspheres. Acta Biomaterialia. 2010;**6**:2200-2211

[46] Klijn RJ, van den Beucken J,

[47] Qi XP, Ye JD. Mechanical and rheological properties and injectability

Lanao R, Veldhuis G, Leeuwenburgh SC, Wolke J, et al. Three different strategies to obtain porous calcium phosphate cements: Comparison of performance in a rat skull bone augmentation model. Tissue Engineering. Part A. 2012;**18**:

[44] Félix Lanao RP, Leeuwenburgh SCG,

2001;**57**:457-466

2011;**7**:3459-3468

1171-1182

calcium carbonate on clinical compliance of apatitic calcium phosphate bone cement. Journal of Biomedical Materials

Research. 1997;**38**(4):356-360

[32] Khairoun I, Driessens FCM, Boltong MG, Planell JA, Wenz R. Addition of cohesion promotors to calcium phosphate cements. Biomaterials. 1999;**20**:393-398

[33] Grover LM, Gbureck U, Wright AJ, Tremayne M, Barralet JE. Biologically mediated resorption of brushite cement in vitro. Biomaterials. 2006;**27**:2178-2185

[34] Ginebra MP. Calcium phosphate bone cements. In: Deb S, editor. Ortopedic Bone Cements. Boca Raton, FL: CRC Press; 2008. pp. 206-230

[35] Ginebra MP, Canal C, Espanol M, Pastorino D, Edgar B. Calcium phosphate cements as drug delivery materials. Advanced Drug Delivery Reviews.

[36] Zhang J, Liu W, Schnitzler V,

[37] Zhang JT, Tancret F, Bouler JM. Fabrication and mechanical properties of calcium phosphate cements (CPC) for bone substitution. Materials Science

& Engineering. C, Materials for

Biomaterialia. 2009;**5**:268-278

[39] Barralet JE, Grover L, Gaunt T, Wright AJ, Gibson IR. Preparation of macroporous calcium phosphate cement tissue engineering scaffold. Biomaterials. 2002;**23**:3063-3072

Biological Applications. 2011;**31**:740-747

[38] Guo H, Su JC, Wei J, Kong H, Liu CS. Biocompatibility and osteogenicity of degradable Ca-deficient hydroxyapatite scaffolds from calcium phosphate cement for bone tissue engineering. Acta

cements for bone substitution: Chemistry, handling and mechanical properties. Acta Biomaterialia. 2014;**10**:

Tancret F, Bouler JM. Calcium phosphate

2012;**64**:1090-1110

1035-1049

**134**

of calcium phosphate cement containing poly (lactic-co-glycolic acid) microspheres. Materials Science & Engineering. C, Materials for Biological Applications. 2009;**29**:1901-1906

[48] Almirall A, Larrecq G, Delgado JA, Martinez S, Planell JA, Ginebra MP. Fabrication of low temperature macroporous hydroxyapatite scaffolds by foaming and hydrolysis of an a-TCP paste. Biomaterials. 2004;**25**:3671-3680

[49] Del Real RP, Wolke J, Vallet-Regi M, Jansen JA. A new method to produce macropores in calcium phosphate cements. Biomaterials. 2002;**23**: 3673-3680

[50] Chen WC, Zhou HZ, Tang MH, Weir MD, Bao CY, Xu H. Gas-foaming calcium phosphate cement scaffold encapsulating human umbilical cord stem cells. Tissue Engineering. Part A. 2012;**18**:816-827

[51] Hesaraki S, Moztarzadeh F, Sharifi D. Formation of interconnected macropores in apatitic calcium phosphate bone cement with the use of an effervescent additive. Journal of Biomedical Materials Research. Part A. 2007;**83A**:80-87

[52] Ginebra M, Delgado J, Harr I, Almirall A, Del Valle S, Planell JA. Factors affecting the structure and properties of an injectable self-setting calcium phosphate foam. Journal of Biomedical Materials Research. Part A. 2007;**80A**:351-361

[53] Bai F, Meng GL, Yuan YA, Liu CS, Wang Z, Liu JA. Role of macropore size in the mechanical properties and in vitro degradation of porous calcium phosphate cements. Materials Letters. 2010;**64**:2028-2031

[54] Liu CS, Shao HF, Chen FY, Zheng HY. Effects of the granularity of raw materials on the hydration and hardening process of calcium

phosphate cement. Biomaterials. 2003;**24**:4103-4113

[55] TenHuisen KS, Brown PW. Formation of calcium-deficient hydroxyapatite from α-tricalcium phosphate. Biomaterials. 1998;**19**: 2209-2217

[56] Bermudez O, Boltong MG, Driessens FCM, Planell JA. Development of some calcium-phosphate cements from combinations of α-TCP, MCPM and CaO. The Journal of Materials Science: Materials in Medicine. 1994;**5**:160-163

[57] Yang QZ, Troczynski T, Liu DM. Influence of apatite seeds on the synthesis of calcium phosphate cement. Biomaterials. 2002;**23**:2751-2760

[58] Tamimi F, Sheikh Z, Barralet J. Dicalcium phosphate cements: Brushite and monetite. Acta Biomaterialia. 2012;**8**:474-487

[59] Sarda S, Fernandez E, Nilsson M, Balcells M, Planell JA. Kinetic study of citric acid influence on calcium phosphate bone cements as waterreducing agent. Journal of Biomedical Materials Research. 2002;**61**:653-659

[60] Gbureck U, Barralet JE, Spatz K, Grover LM, Thull R. Ionic modification of calcium phosphate cement viscosity. Part I: Hypodermic injection and strength improvement of apatite cement. Biomaterials. 2004;**25**:2187-2195

[61] Barralet JE, Hofmann M, Grover LM, Gbureck U. High-strength apatitic cement by modification with α-hydroxy acid salts. Advanced Materials. 2003;**15**:2091-2094

[62] Barralet JE, Grover LM, Gbureck U. Ionic modification of calcium phosphate cement viscosity. Part II: Hypodermic injection and strength improvement of brushite cement. Biomaterials. 2004;**25**:2197-2203

[63] Qi XP, Ye JD, Wang YJ. Improved injectability and in vitro degradation of a calcium phosphate cement containing poly(lactide-co-glycolide) microspheres. Acta Biomaterialia. 2008;**4**:1837-1845

[64] Marino FT, Torres J, Hamdan M, Rodriguez CR, Cabarcos EL. Advantages of using glycolic acid as a retardant in a brushite forming cement. Journal of Biomedical Materials Research. 2007;**83B**:571-579

[65] Munz D, Fett D. Ceramics: Mechanical Properties, Failure Behavior, Materials Selection. Germany: Springer; 2001

[66] Munz D. What can we learn from r-curve measurements? Journal of the American Ceramic Society. 2007;**90**:1-15

[67] Ritchie RO. The conflicts between strength and toughness. Nature Materials. 2011;**10**:817-822

[68] Launey ME, Ritchie RO. On the fracture toughness of advanced materials. Advanced Materials. 2009;**21**:2103-2110

[69] Chow LC, Hirayama S, Takagi S, Parry E. Diametral tensile strength and compressive strength of a calcium phosphate cement: Effect of applied pressure. Journal of Biomedical Materials Research. 2000;**53**:511-517

[70] Hofmann MP, Mohammed AR, Perrie Y, Gbureck U, Barralet JE. Highstrength resorbable brushite bone cement with controlled drug-releasing capabilities. Acta Biomaterialia. 2009;**5**:43-49

[71] Chow LC. Development of selfsetting calcium phosphate cements. Journal of the Ceramic Society of Japan (International Edition). 1991;**99**:927-936

[72] Ginebra MP, Rilliard A, Fernández E, Elvira C, San Román J, Planell JA. Mechanical and rheological improvement of a calcium phosphate cement by the addition of a polymeric drug. Journal of Biomedical Materials Research. 2001;**57**:113-118

[73] Driessens FCM, Boltong MG, Bermúdez O, Planell JA, Ginebra MP, Fernández E. Effective formulations for the preparation of calcium phosphate bone cements. Journal of Materials Science. Materials in Medicine. 1994;**5**:164-170

[74] Ginebra MP. Desarrollo y caracterización de un cemento óseo basado en fosfato tricálcico-α para aplicaciones quirúrgicas [PhD thesis]. Barcelona, Spain: Universitat Politècnica de Catalunya; 1996

[75] Chow LC. Calcium phosphate cements: Chemistry, properties, and applications. MRS Proceedings. 1999;**599**:27

[76] Xu HH, Weir MD, Burguera EF, et al. Injectable and macroporous calcium phosphate cement scaffold. Biomaterials. 2006;**27**:4279-4287

[77] Ginebra MP, Traykova T, Planell JA. Calcium phosphate cements as bone drug delivery systems: A review. Journal of Controlled Release. 2006;**113**:102-110

[78] Weir MD, Xu HH. Osteoblastic induction on calcium phosphate cement-chitosan constructs for bone tissue engineering. Journal of Biomedical Materials Research. Part A. 2010;**94**:223-233

[79] Mestres G, Le Van C, Ginebra MP. Silicon-stabilized α-tricalcium phosphate and its use in a calcium phosphate cement: Characterization and cell response. Acta Biomaterialia. 2012;**8**:1169-1179

[80] Lanao RPF, Leeuwenburgh SC, Wolke JG, et al. Bone response to fast-degrading, injectable calcium

**137**

2002;**395**:23-32

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

> implanted in cortical bone. Biomaterials. 2003;**24**:989-1000

1998;**80**:1112

337-344

2003;**24**:3463-3474

[88] Frankenburg EP, Goldstein SA, Bauer TW, Harris SA, Poser RD. Biomechanical and histological evaluation of a calcium phosphate cement. The Journal of Bone and Joint Surgery. American Volume.

[89] Apelt D, Theiss F, El-warrak AO, Zlinszky K, Bettschart-Wolfisberger R, Bohner M, et al. In vivo behavior of three different injectable hydraulic calcium phosphate cements. Biomaterials. 2004;**25**:1439-1451

[90] Ohura K, Bohner M, Hardouin P, Lemaitre J, Pasquier G, Flautre B. Resorption of, and bone formation from, new betatricalcium phosphatemonocalcium phosphate cements: An in vivo study. Journal of Biomedical Materials Research. 1996;**30**:193-200

[91] Munting E, Mirtchi AA, Lemaitre J. Bone repair of defects filled with phosphocalcic hydraulic cement: an in vivo study. Journal of Materials Science. Materials in Medicine. 1993;**4**:

[92] Bohner M, Theiss F, Apelt D, Hirsiger W, Houriet R, Rizzoli G, et al. Compositional changes of a dicalcium phosphate dihydrate cement after implantation in sheep. Biomaterials.

[93] Lu J, Descamps M, Dejou J, et al. The biodegradation mechanism of calcium phosphate biomaterials in bone. Journal of Biomedical Materials Research. Part A. 2002;**63**:408-412

orthophosphate cements and concretes.

[95] Morgan E, Yetkinler D, Constantz B, Dauskardt R. Mechanical properties of carbonated apatite bone mineral

[94] Dorozhkin SV. Calcium

Materials. 2009;**2**:221-291

phosphate cements containing PLGA microparticles. Biomaterials.

[81] Theiss F, Apelt D, Brand B, et al. Biocompatibility and resorption of a brushite calcium phosphate cement. Biomaterials. 2005;**26**:4383-4394

[82] Noetzel J, Özer K, Reisshauer B-H, et al. Tissue responses to an experimental calcium phosphate cement and mineral trioxide aggregate as materials for furcation perforation repair: A histological study in dogs. Clinical Oral Investigations. 2006;**10**:77

[83] Kurashina K, Kurita H, Kotani A, Klein CPAT, Groot K. In vivo study of calcium phosphate cements: Implantation of an α-tricalcium phosphate/dicalcium phosphate dibasic/tetracalcium phosphate monoxide cement paste. Biomaterials.

[84] Jansen JA, Ruijter JE, Schaeken HG, van der Waerden JPC, Planell JA, Driessens FCM. Evaluation of tricalciumphosphate/hydroxyapatite cement for tooth replacement: An experimental animal study. Journal of Materials Science: Materials in

[85] Friedman CD, Costantino PD, Takagi S, Chow LC. Bonesource hydroxyapatite cement: A novel biomaterial for cranofacial skeletal tissue engineering

Biomedical Materials Research Part B: Applied Biomaterials. 1998;**43**:428-432

2011;**32**:8839-8847

1997;**18**:539-543

Medicine. 1995;**6**:653-657

and reconstruction. Journal of

[86] Larsson S, Bauer TW. Use of injectable calcium phosphate cement for fracture fixation: A review. Clinical Orthopaedics and Related Research.

[87] Ooms EM, Wolke JGC, van de Heuvel MT, Jeschke B, Jansen JA. Histological evaluation of the bone response to calcium phosphate cement *Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

phosphate cements containing PLGA microparticles. Biomaterials. 2011;**32**:8839-8847

*Contemporary Topics about Phosphorus in Biology and Materials*

Planell JA. Mechanical and rheological improvement of a calcium phosphate cement by the addition of a polymeric drug. Journal of Biomedical Materials

Research. 2001;**57**:113-118

1994;**5**:164-170

de Catalunya; 1996

1999;**599**:27

2010;**94**:223-233

2012;**8**:1169-1179

[73] Driessens FCM, Boltong MG, Bermúdez O, Planell JA, Ginebra MP, Fernández E. Effective formulations for the preparation of calcium phosphate bone cements. Journal of Materials Science. Materials in Medicine.

[74] Ginebra MP. Desarrollo y caracterización de un cemento óseo basado en fosfato tricálcico-α para aplicaciones quirúrgicas [PhD thesis]. Barcelona, Spain: Universitat Politècnica

[75] Chow LC. Calcium phosphate cements: Chemistry, properties, and applications. MRS Proceedings.

[76] Xu HH, Weir MD, Burguera EF, et al. Injectable and macroporous calcium phosphate cement scaffold. Biomaterials. 2006;**27**:4279-4287

[77] Ginebra MP, Traykova T, Planell JA. Calcium phosphate cements as bone drug delivery systems: A review. Journal of Controlled Release. 2006;**113**:102-110

[78] Weir MD, Xu HH. Osteoblastic induction on calcium phosphate cement-chitosan constructs for bone tissue engineering. Journal of Biomedical Materials Research. Part A.

[79] Mestres G, Le Van C, Ginebra MP.

[80] Lanao RPF, Leeuwenburgh SC, Wolke JG, et al. Bone response to fast-degrading, injectable calcium

Silicon-stabilized α-tricalcium phosphate and its use in a calcium phosphate cement: Characterization and cell response. Acta Biomaterialia.

[63] Qi XP, Ye JD, Wang YJ. Improved injectability and in vitro degradation of a calcium phosphate cement containing poly(lactide-co-glycolide) microspheres. Acta Biomaterialia.

[64] Marino FT, Torres J, Hamdan M, Rodriguez CR, Cabarcos EL. Advantages of using glycolic acid as a retardant in a brushite forming cement. Journal of Biomedical Materials Research.

[65] Munz D, Fett D. Ceramics: Mechanical Properties, Failure Behavior, Materials Selection. Germany: Springer; 2001

[66] Munz D. What can we learn from r-curve measurements? Journal of the American Ceramic Society.

[67] Ritchie RO. The conflicts between strength and toughness. Nature Materials. 2011;**10**:817-822

[68] Launey ME, Ritchie RO. On the fracture toughness of advanced materials. Advanced Materials. 2009;**21**:2103-2110

[69] Chow LC, Hirayama S, Takagi S, Parry E. Diametral tensile strength and compressive strength of a calcium phosphate cement: Effect of applied pressure. Journal of Biomedical Materials Research. 2000;**53**:511-517

[70] Hofmann MP, Mohammed AR, Perrie Y, Gbureck U, Barralet JE. Highstrength resorbable brushite bone cement with controlled drug-releasing capabilities. Acta Biomaterialia.

[71] Chow LC. Development of selfsetting calcium phosphate cements. Journal of the Ceramic Society of Japan (International Edition).

Fernández E, Elvira C, San Román J,

2008;**4**:1837-1845

2007;**83B**:571-579

2007;**90**:1-15

**136**

2009;**5**:43-49

1991;**99**:927-936

[72] Ginebra MP, Rilliard A,

[81] Theiss F, Apelt D, Brand B, et al. Biocompatibility and resorption of a brushite calcium phosphate cement. Biomaterials. 2005;**26**:4383-4394

[82] Noetzel J, Özer K, Reisshauer B-H, et al. Tissue responses to an experimental calcium phosphate cement and mineral trioxide aggregate as materials for furcation perforation repair: A histological study in dogs. Clinical Oral Investigations. 2006;**10**:77

[83] Kurashina K, Kurita H, Kotani A, Klein CPAT, Groot K. In vivo study of calcium phosphate cements: Implantation of an α-tricalcium phosphate/dicalcium phosphate dibasic/tetracalcium phosphate monoxide cement paste. Biomaterials. 1997;**18**:539-543

[84] Jansen JA, Ruijter JE, Schaeken HG, van der Waerden JPC, Planell JA, Driessens FCM. Evaluation of tricalciumphosphate/hydroxyapatite cement for tooth replacement: An experimental animal study. Journal of Materials Science: Materials in Medicine. 1995;**6**:653-657

[85] Friedman CD, Costantino PD, Takagi S, Chow LC. Bonesource hydroxyapatite cement: A novel biomaterial for cranofacial skeletal tissue engineering and reconstruction. Journal of Biomedical Materials Research Part B: Applied Biomaterials. 1998;**43**:428-432

[86] Larsson S, Bauer TW. Use of injectable calcium phosphate cement for fracture fixation: A review. Clinical Orthopaedics and Related Research. 2002;**395**:23-32

[87] Ooms EM, Wolke JGC, van de Heuvel MT, Jeschke B, Jansen JA. Histological evaluation of the bone response to calcium phosphate cement implanted in cortical bone. Biomaterials. 2003;**24**:989-1000

[88] Frankenburg EP, Goldstein SA, Bauer TW, Harris SA, Poser RD. Biomechanical and histological evaluation of a calcium phosphate cement. The Journal of Bone and Joint Surgery. American Volume. 1998;**80**:1112

[89] Apelt D, Theiss F, El-warrak AO, Zlinszky K, Bettschart-Wolfisberger R, Bohner M, et al. In vivo behavior of three different injectable hydraulic calcium phosphate cements. Biomaterials. 2004;**25**:1439-1451

[90] Ohura K, Bohner M, Hardouin P, Lemaitre J, Pasquier G, Flautre B. Resorption of, and bone formation from, new betatricalcium phosphatemonocalcium phosphate cements: An in vivo study. Journal of Biomedical Materials Research. 1996;**30**:193-200

[91] Munting E, Mirtchi AA, Lemaitre J. Bone repair of defects filled with phosphocalcic hydraulic cement: an in vivo study. Journal of Materials Science. Materials in Medicine. 1993;**4**: 337-344

[92] Bohner M, Theiss F, Apelt D, Hirsiger W, Houriet R, Rizzoli G, et al. Compositional changes of a dicalcium phosphate dihydrate cement after implantation in sheep. Biomaterials. 2003;**24**:3463-3474

[93] Lu J, Descamps M, Dejou J, et al. The biodegradation mechanism of calcium phosphate biomaterials in bone. Journal of Biomedical Materials Research. Part A. 2002;**63**:408-412

[94] Dorozhkin SV. Calcium orthophosphate cements and concretes. Materials. 2009;**2**:221-291

[95] Morgan E, Yetkinler D, Constantz B, Dauskardt R. Mechanical properties of carbonated apatite bone mineral

substitute: Strength, fracture and fatigue behavior. Journal of Materials Science. Materials in Medicine. 1997;**8**:559-570

[96] Montufar EB, Traykova T, Schacht E, Ambrosio L, Santin M, Planell JA, et al. Self-hardening calcium deficient hydroxyapatite/gelatine foams for bone regeneration. Journal of Materials Science. Materials in Medicine. 2010;**21**:863-869

[97] Fernández E, Gil FJ, Ginebra MP, Driessens FCM, Planell JA, Best S. Calcium phosphate bone cements for clinical applications. Part II: Precipitate formation during setting reactions. Journal of Materials Science: Materials in Medicine. 1999;**10**:177-183

[98] Miyamoto Y, Ishikawa K, Fukao H, Sawada M, Nagagama M, Kon M, et al. In vivo setting behaviour of fastsetting calcium phosphate cement. Biomaterials. 1995;**16**:855-860

[99] Ikenaga M, Hardouin I, Lemaitre J, Andrianjatovo H, Flautre B. Biomechanical characterization of a biodegradable calcium phosphate hydraulic cement: A comparison with porous biphasic calcium phosphate ceramics. Journal of Biomedical Materials Research. 1998;**40**:139-144

[100] Muenzenberg K, Gebhardt M. Brushite, octocalcium phosphate, and carbonate containing apatite in bone. Clinical Orthopaedics and Related Research. 1973;**90**:271-273

[101] Terjesen T. Bone healing after metal plate fixation and external fixation of the osteotomized rabbit tibia. Acta Orthopaedica Scandinavica. 1984;**55**:69

[102] Lemaitre J, Mirtchi A, Mortier A. Calcium phosphate cements for medical use: State of the art and perspectives of development. Silicates Industriels. 1987;**52**:141-146

[103] Mirtchi AA, Lemaitre J, Terao N. Calcium phosphate cements: Study of the [beta]-tricalcium phosphate– monocalcium phosphate system. Biomaterials. 1989;**10**:475-480

[104] Bajpai P, Fuchs C, McCullum DE. Development of tricalcium phosphate ceramic cements. In: Lemons J, editor. Quantitative Characterization and Performance of Porous Implants for Hard Tissue Applications. Philadelphia, USA: American Society for Testing Materials. 1987. pp. 377-388

[105] Elliott J. Structure and Chemistry of the Apatites and Other Calcium Orthophosphates. Amsterdam: Elsevier; 1994

[106] Bohner M, Van Landuyt P, Merkle H, Lemaitre J. Composition effects on the pH of a hydraulic calcium phosphate cement. The Journal of Materials Science: Materials in Medicine. 1997;**8**:675-681

[107] Lemaitre J, Mirtchi A, Mortier A. Calcium phosphate cements for medical use: state of the art and perspectives of development. Silicates Industriels. 1987;**9-10**:141-6

[108] Bohner M. Proprietes physicochimiques et osteogeniques d'un biociment hydraulique a base de phosphates de calcium [PhD thesis No. 1171]. Lausanne: Swiss Federal Institute of Technology of Lausanne (EPFL); 1993

[109] Bohner M, Van Landuyt P, Trophardy G, Merkle H, Lemaitre J. Effect of several additives and their admixtures on the physico-chemical properties of a calcium phosphate cement. Journal of Materials Science. Materials in Medicine. 2000;**11**:111-116

[110] Andrianjatovo H, Jose F, Lemaitre J. Effect of b-TCP granulometry on setting time and strength of calcium phosphate hydraulic cements. Journal of Materials Science. Materials in Medicine. 1996;**7**:34-39

**139**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

> coatings on orthopedic alloys. Journal of Biomedial Materials Research Part A.

[120] Xie J, Riley C, Chittur K. Effect of albumin on brushite transformation to hydroxyapatite. Journal of Biomedical Materials Research. 2001;**57**:357-365

[121] Levinskas G, Neuman W. The solubility of bone mineral. I. Solubility studies of synthetic hydroxylapatite. The Journal of Physical Chemistry.

[122] Strates B, Neuman W, Levinskas G. The solubility of bone mineral. II. Precipitation of near-neutral solutions of calcium and phosphate. The Journal of Physical Chemistry. 1957;**61**:279-282

[123] Xia Z, Grover L, Huang Y, Adamopoulos I, Gbureck U, Triffitt J, et al. In vitro biodegradation of three brushite calcium phosphate cements by a macrophage cell-line. Biomaterials.

[124] Tamimi F, Kumarasami B,

Biomaterialia. 2008;**4**:1315

Surgery. 2005;**33**:37-44

Doillon C, Gbureck U, Le Nihouannen D, Cabarcos E, et al. Brushite–collagen composites for bone regeneration. Acta

[125] Klammert U, Reuther T, Jahn C, Kraski B, Kbler A, Gbureck U. Cytocompatibility of brushite and monetite cell culture scaffolds made by three-dimensional powder printing. Acta Biomaterialia. 2009;**5**:727

[126] Kuemmerle J, Oberle A, Oechslin C, Bohner M, Frei C, Boecken I, et al. Assessment of the suitability of a new brushite calcium phosphate cement for cranioplasty—An experimental study in sheep. Journal of Cranio-Maxillo-Facial

[127] Ji C, Ahn J. Clinical experience of the brushite calcium phosphate cement for the repair and augmentation of surgically induced cranial defects

1996;**30**:287-294

1955;**59**:164-168

2006;**27**:4557-4565

[111] Ishikawa K, Takagi S, Chow L, Ishikawa Y, Eanes E, Asaoka K. Behavior of a calcium phosphate cement in simulated blood plasma in vitro. Dental

[112] Driessens F. Chemistry and applied aspects of calcium phosphate bone cements. In: Concepts and

Presented at the 15th European Conference on Biomaterials. ESB 99,

Arcachon, France; 1999

Bone. 1999;**25**:35-39

[115] Bohner M. Calcium

Clinical Applications of Ionic Cements.

[113] Shadanbaz S, Dias GJ. Calcium phosphate coatings on magnesium alloys for biomedical applications: A review. Acta Biomaterialia. 2012;**8**:20-30

[114] Hardouin P, Delecourt C, Blary M, Van Landuyt I, Lemaitre J, Hardouin L. Volume effect on biological properties of a calcium phosphate hydraulic cement: Experimental study in sheep.

orthophosphates in medicine: From ceramics to calcium phosphate cements. Injury-International Journal of the Care of the Injured. 2000;**31**:S-D37-47

[116] Bohner M, Matter S. Brushite hydraulic cement stabilized with a magnesium salt. PCT application PCT/ CH99/00595, Switzerland. 1999

[117] Kumar M, Xie J, Chittur K, Riley C. Transformation of modified brushite to hydroxyapatite in aqueous solution: Effects of potassium substitution. Biomaterials. 1999;**20**:1389-1399

[118] Kumar M, Dasarathy H, Riley C.

[119] Redepenning J, Schlessinger T, Burnham S, Lippiello L, Miyano J. Characterization of electrolytically prepared brushite and hydroxyapatite

Electrodeposition of brushite coatings and their transformation to hydroxyapatite in aqueous solutions. Journal of Biomedial Materials Research

Part A. 1999;**45**:302-310

Materials. 1994;**10**:26-32

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

*Contemporary Topics about Phosphorus in Biology and Materials*

[103] Mirtchi AA, Lemaitre J, Terao N. Calcium phosphate cements: Study of the [beta]-tricalcium phosphate– monocalcium phosphate system. Biomaterials. 1989;**10**:475-480

[104] Bajpai P, Fuchs C, McCullum DE. Development of tricalcium phosphate ceramic cements. In: Lemons J, editor. Quantitative Characterization and Performance of Porous Implants for Hard Tissue Applications. Philadelphia, USA: American Society for Testing Materials. 1987. pp. 377-388

[105] Elliott J. Structure and Chemistry of the Apatites and Other Calcium Orthophosphates. Amsterdam: Elsevier;

calcium phosphate cement. The Journal of Materials Science: Materials in

[107] Lemaitre J, Mirtchi A, Mortier A. Calcium phosphate cements for medical use: state of the art and perspectives of development. Silicates Industriels.

[108] Bohner M. Proprietes physicochimiques et osteogeniques d'un biociment hydraulique a base de phosphates de calcium [PhD thesis No. 1171]. Lausanne: Swiss Federal Institute of Technology of Lausanne (EPFL); 1993

[109] Bohner M, Van Landuyt P, Trophardy G, Merkle H, Lemaitre J. Effect of several additives and their admixtures on the physico-chemical properties of a calcium phosphate cement. Journal of Materials Science. Materials in Medicine. 2000;**11**:111-116

Science. Materials in Medicine.

1996;**7**:34-39

[110] Andrianjatovo H, Jose F, Lemaitre J. Effect of b-TCP granulometry on setting time and strength of calcium phosphate hydraulic cements. Journal of Materials

[106] Bohner M, Van Landuyt P, Merkle H, Lemaitre J. Composition effects on the pH of a hydraulic

Medicine. 1997;**8**:675-681

1987;**9-10**:141-6

1994

substitute: Strength, fracture and fatigue behavior. Journal of Materials Science. Materials in Medicine.

Science. Materials in Medicine.

in Medicine. 1999;**10**:177-183

[99] Ikenaga M, Hardouin I,

[97] Fernández E, Gil FJ, Ginebra MP, Driessens FCM, Planell JA, Best S. Calcium phosphate bone cements for clinical applications. Part II: Precipitate formation during setting reactions. Journal of Materials Science: Materials

[98] Miyamoto Y, Ishikawa K, Fukao H, Sawada M, Nagagama M, Kon M, et al. In vivo setting behaviour of fastsetting calcium phosphate cement. Biomaterials. 1995;**16**:855-860

Lemaitre J, Andrianjatovo H, Flautre B. Biomechanical characterization of a biodegradable calcium phosphate hydraulic cement: A comparison with porous biphasic calcium phosphate ceramics. Journal of Biomedical Materials Research. 1998;**40**:139-144

[100] Muenzenberg K, Gebhardt M. Brushite, octocalcium phosphate, and carbonate containing apatite in bone. Clinical Orthopaedics and Related

[101] Terjesen T. Bone healing after metal plate fixation and external fixation of the osteotomized rabbit tibia. Acta Orthopaedica Scandinavica. 1984;**55**:69

[102] Lemaitre J, Mirtchi A, Mortier A. Calcium phosphate cements for medical use: State of the art and perspectives of development. Silicates Industriels.

Research. 1973;**90**:271-273

[96] Montufar EB, Traykova T, Schacht E, Ambrosio L, Santin M, Planell JA, et al. Self-hardening calcium deficient hydroxyapatite/gelatine foams for bone regeneration. Journal of Materials

1997;**8**:559-570

2010;**21**:863-869

**138**

1987;**52**:141-146

[111] Ishikawa K, Takagi S, Chow L, Ishikawa Y, Eanes E, Asaoka K. Behavior of a calcium phosphate cement in simulated blood plasma in vitro. Dental Materials. 1994;**10**:26-32

[112] Driessens F. Chemistry and applied aspects of calcium phosphate bone cements. In: Concepts and Clinical Applications of Ionic Cements. Presented at the 15th European Conference on Biomaterials. ESB 99, Arcachon, France; 1999

[113] Shadanbaz S, Dias GJ. Calcium phosphate coatings on magnesium alloys for biomedical applications: A review. Acta Biomaterialia. 2012;**8**:20-30

[114] Hardouin P, Delecourt C, Blary M, Van Landuyt I, Lemaitre J, Hardouin L. Volume effect on biological properties of a calcium phosphate hydraulic cement: Experimental study in sheep. Bone. 1999;**25**:35-39

[115] Bohner M. Calcium orthophosphates in medicine: From ceramics to calcium phosphate cements. Injury-International Journal of the Care of the Injured. 2000;**31**:S-D37-47

[116] Bohner M, Matter S. Brushite hydraulic cement stabilized with a magnesium salt. PCT application PCT/ CH99/00595, Switzerland. 1999

[117] Kumar M, Xie J, Chittur K, Riley C. Transformation of modified brushite to hydroxyapatite in aqueous solution: Effects of potassium substitution. Biomaterials. 1999;**20**:1389-1399

[118] Kumar M, Dasarathy H, Riley C. Electrodeposition of brushite coatings and their transformation to hydroxyapatite in aqueous solutions. Journal of Biomedial Materials Research Part A. 1999;**45**:302-310

[119] Redepenning J, Schlessinger T, Burnham S, Lippiello L, Miyano J. Characterization of electrolytically prepared brushite and hydroxyapatite coatings on orthopedic alloys. Journal of Biomedial Materials Research Part A. 1996;**30**:287-294

[120] Xie J, Riley C, Chittur K. Effect of albumin on brushite transformation to hydroxyapatite. Journal of Biomedical Materials Research. 2001;**57**:357-365

[121] Levinskas G, Neuman W. The solubility of bone mineral. I. Solubility studies of synthetic hydroxylapatite. The Journal of Physical Chemistry. 1955;**59**:164-168

[122] Strates B, Neuman W, Levinskas G. The solubility of bone mineral. II. Precipitation of near-neutral solutions of calcium and phosphate. The Journal of Physical Chemistry. 1957;**61**:279-282

[123] Xia Z, Grover L, Huang Y, Adamopoulos I, Gbureck U, Triffitt J, et al. In vitro biodegradation of three brushite calcium phosphate cements by a macrophage cell-line. Biomaterials. 2006;**27**:4557-4565

[124] Tamimi F, Kumarasami B, Doillon C, Gbureck U, Le Nihouannen D, Cabarcos E, et al. Brushite–collagen composites for bone regeneration. Acta Biomaterialia. 2008;**4**:1315

[125] Klammert U, Reuther T, Jahn C, Kraski B, Kbler A, Gbureck U. Cytocompatibility of brushite and monetite cell culture scaffolds made by three-dimensional powder printing. Acta Biomaterialia. 2009;**5**:727

[126] Kuemmerle J, Oberle A, Oechslin C, Bohner M, Frei C, Boecken I, et al. Assessment of the suitability of a new brushite calcium phosphate cement for cranioplasty—An experimental study in sheep. Journal of Cranio-Maxillo-Facial Surgery. 2005;**33**:37-44

[127] Ji C, Ahn J. Clinical experience of the brushite calcium phosphate cement for the repair and augmentation of surgically induced cranial defects

following the pterional craniotomy. Journal of Korean Neurosurgical Association. 2010;**47**:180

[128] Bose S, Tarafder S. Calcium phosphate ceramic systems in growth factor and drug delivery for bone tissue engineering: A review. Acta Biomaterialia. 2012;**8**:1401-1421

[129] Bigi A, Bracci B, Panzavolta S. Effect of added gelatin on the properties of calcium phosphate cement. Biomaterials. 2004;**25**:2893-2899

[130] Ratier A, Best S, Freche M, Lacout J, Rodriguez F. Behaviour of a calcium phosphate bone cement containing tetracycline hydrochloride or tetracycline complexed with calcium ions. Biomaterials. 2001;**22**:897-901

[131] Nimni ME. Polypeptide growth factors: Targeted delivery systems. Biomaterials. 1997;**18**:1201-1225

[132] Centrella M, Massague J, Canalis E. Human platelet-derived transforming growth factor β stimulates parameters of bone growth in fetal rat calvaria. Endocrinology. 1986;**119**:2306-2312

[133] Noda M, Camilliere JJ. In vivo stimulation of bone formation by transforming growth factor β. Endocrynology. 1989;**124**:2991-2994

[134] Bosch C, Melsen B, Gibbons R, Vargervik K. Human recombinant transforming growth factor beta 1 in healing of calvarial bone defects. Journal of Craniofacial Surgery. 1996;**7**:300-310

[135] Lind M, Overgaard S, Soballe K, Nguyen T, Ongpipattanakul B, Bunger C. Transforming growth factor-beta 1 enhances bone healing to unloaded tricalcium phosphate coated implants: An experimental study in dogs. Journal of Orthopaedic Research. 1996;**14**:343-350

[136] Blom EJ, Klein-Nulend J, Klein CPAT, Kurashina K, van Waas MAJ, Burger EH. Transforming growth factor-β1 incorporated during setting in calcium phosphate cement stimulates bone cell differentiation in vitro. Journal of Biomedical Materials Research. 2000;**50**:67-74

[137] Blom EJ, Klein-Nulend J, Yin L, van Waas MAJ, Burger EH. Transforming growth factor-β1 in calcium phosphate cement stimulates bone regeneration. Journal of Dental Research. 2000;**79**:255

[138] Ruhe PQ, Hedberg EL, Padron NT, Spauwen PH, Jansen JA, Mikos AG. rhBMP-2 release from injectable poly(DL-lactic-co-glycolic acid)/ calcium-phosphate cement composites. The Journal of Bone and Joint Surgery. 2003;**85**:75-82

[139] Haddad AJ et al. Closure of rabbit calvarial critical-sized defects using protective composite allogeneic and alloplastic bone substitutes. The Journal of Craniofacial Surgery. 2006;**17**:926-934

[140] Seeherman HJ et al. Recombinant human bone morphogenetic protein-2 delivered in an injectable calcium phosphate paste accelerates osteotomysite healing in a nonhuman primate model. The Journal of Bone and Joint Surgery. American Volume. 2004;**86-A**: 1961-1972

[141] Seeherman HJ et al. rhBMP-2 delivered in a calcium phosphate cement accelerates bridging of critical-sized defects in rabbit radii. The Journal of Bone and Joint Surgery. American Volume. 2006;**88**:1553-1565

[142] Ruchholtz S, Tager G, Nast-Kolb D. The periprosthetic total hip infection. Unfallchirurg. 2004;**107**:307-317

[143] Harris W, Sledge CB. Total hip and total knee replacement (Part II). The New England Journal of Medicine. 1990;**323**:801-807

**141**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

[152] Schnieders J, Gbureck U,

[153] Anttipoika I, Josefsson G, Konttinen Y, Lidgren L, Santavirta S, Sanzen L. Hip-arthroplasty infection— Current concepts. Acta Orthopaedica Scandinavica. 1990;**61**:163-169

[154] Barralet JE, Aldred S,

hydroxyapatite gel. Journal of Biomedical Materials Research.

[155] Gitelis S, Brebach GT. The

implant. Journal of Orthopaedic

Surgery. 2002;**10**:53-60

treatment of chronic osteomyelitis with a biodegradable antibiotic-impregnated

[156] Hendriks JGE, van Horn JR, van der Mei HC, Busscher HJ. Backgrounds of antibiotic-loaded bone cement and prosthesis-related infection. Biomaterials. 2004;**25**:545-556

[157] Pitto RP, Spika IA. Antibioticloaded bone cement spacers in two-stage management of infected total knee arthroplasty, International Orthopaedics. 2004;**28**:129-133

[158] Durbhakula SM, Czajka J, Fuchs MD, Uhl RL. Spacer

endoprosthesis for the treatment of infected total hip arthroplasty. The Journal of Arthroplasty. 2004;**19**:760-767

[159] Ewald A, Gluckermann SK, Thull R, Gbureck U. Antimicrobial titanium/silver PVD coatings on titanium, BioMedical Engineering

[160] Yorganci K, Krepel C, Weigelt JA, Edmiston CE. Activity of antibacterial impregnated central venous catheters

Online. 2006;**5**:22

2002;**60**:360-367

Wright AJ, Coombes AGA. In vitro behavior of albumin-loaded carbonate

2006;**27**:4239-4249

Thull R, Kissel T. Controlled release of gentamicin from calcium phosphate– poly(lactic acid-co-glycolic acid) composite bone cement. Biomaterials.

[144] Lew DP, Waldvogel FA. Osteomyelitis. The Lancet. 2004;**364**:369-379

[145] Bohner M, Lemaitre J, Van Landuyt P, Zambelli PY, Merkle HP, Gander B. Gentamicin-loaded hydraulic

calcium phosphate bone cement as antibiotic delivery system. Journal of Pharmaceutical Sciences.

[146] Alt V, Bechert T, Steinrcke P, Wagener M, Seidel P, Dingeldein E, et al. An in vitro assessment of the antibacterial properties and cytotoxicity of nanoparticulate silver bone cement. Biomaterials. 2004;**25**:4383-4391

[147] Itokazu M, Wenyi Y, Aoki T, Ohara A, Kato N. Synthesis of antibiotic-loaded interporous hydroxyapatite blocks by vacuum method and in vitro drug release testing. Biomaterials. 1998;**19**:817-819

[148] Penner MJ, Masri BA, Duncan CP. Elution characteristics of vancomycin

and tobramycin combined in acrylic bone-cement. The Journal of Arthroplasty. 1996;**11**:939-944

2005;**26**:7276-7285

2002;**23**:3787-3797

[149] Dion A, Langman M, Hall G, Filiaggi M. Vancomycin release behaviour from amorphous calcium polyphosphate matrices intended for osteomyelitis treatment. Biomaterials.

[150] Jiang PJ, Patel S, Gbureck U,

[151] Frutos P, Pena E, Frutos G, Barrales-Rienda JM. Release of gentamicin sulphate from modified commercial bone cement. Effect of (2-hydroxyethyl methacrylate) comonomer and poly(N-vinyl-2 pyrrolidone) additive on release mechanism and kinetics. Biomaterials.

Grover LM. A comparison of the efficacy of hydroxyapatite based cements and gels as drug delivery matrices. Key Engineering Materials. 2008;**93**:327-330

1997;**86**:565-572

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

*Contemporary Topics about Phosphorus in Biology and Materials*

[136] Blom EJ, Klein-Nulend J,

2000;**50**:67-74

2003;**85**:75-82

2006;**17**:926-934

1961-1972

Klein CPAT, Kurashina K, van Waas MAJ, Burger EH. Transforming growth factor-β1 incorporated during setting in calcium phosphate cement stimulates bone cell differentiation in vitro. Journal of Biomedical Materials Research.

[137] Blom EJ, Klein-Nulend J, Yin L, van Waas MAJ, Burger EH. Transforming growth factor-β1 in calcium phosphate cement stimulates bone regeneration. Journal of Dental Research. 2000;**79**:255

[138] Ruhe PQ, Hedberg EL, Padron NT, Spauwen PH, Jansen JA, Mikos AG. rhBMP-2 release from injectable poly(DL-lactic-co-glycolic acid)/ calcium-phosphate cement composites. The Journal of Bone and Joint Surgery.

[139] Haddad AJ et al. Closure of rabbit calvarial critical-sized defects using protective composite allogeneic and alloplastic bone substitutes. The Journal of Craniofacial Surgery.

[140] Seeherman HJ et al. Recombinant human bone morphogenetic protein-2 delivered in an injectable calcium phosphate paste accelerates osteotomysite healing in a nonhuman primate model. The Journal of Bone and Joint Surgery. American Volume. 2004;**86-A**:

[141] Seeherman HJ et al. rhBMP-2 delivered in a calcium phosphate cement accelerates bridging of critical-sized defects in rabbit radii. The Journal of Bone and Joint Surgery. American

[142] Ruchholtz S, Tager G, Nast-Kolb D. The periprosthetic total hip infection. Unfallchirurg. 2004;**107**:307-317

[143] Harris W, Sledge CB. Total hip and total knee replacement (Part II). The New England Journal of Medicine.

Volume. 2006;**88**:1553-1565

1990;**323**:801-807

following the pterional craniotomy. Journal of Korean Neurosurgical

[128] Bose S, Tarafder S. Calcium phosphate ceramic systems in growth factor and drug delivery for bone tissue engineering: A review. Acta Biomaterialia. 2012;**8**:1401-1421

[129] Bigi A, Bracci B, Panzavolta S. Effect of added gelatin on the properties

of calcium phosphate cement. Biomaterials. 2004;**25**:2893-2899

[130] Ratier A, Best S, Freche M, Lacout J, Rodriguez F. Behaviour of a calcium phosphate bone cement containing tetracycline hydrochloride or tetracycline complexed with calcium ions. Biomaterials. 2001;**22**:897-901

[131] Nimni ME. Polypeptide growth factors: Targeted delivery systems. Biomaterials. 1997;**18**:1201-1225

[132] Centrella M, Massague J, Canalis E. Human platelet-derived transforming growth factor β stimulates parameters of bone growth in fetal rat calvaria. Endocrinology. 1986;**119**:2306-2312

[133] Noda M, Camilliere JJ. In vivo stimulation of bone formation by transforming growth factor β. Endocrynology. 1989;**124**:2991-2994

[134] Bosch C, Melsen B, Gibbons R, Vargervik K. Human recombinant transforming growth factor beta 1 in healing of calvarial bone defects. Journal of Craniofacial Surgery.

[135] Lind M, Overgaard S, Soballe K, Nguyen T, Ongpipattanakul B, Bunger C. Transforming growth factor-beta 1 enhances bone healing to unloaded tricalcium phosphate coated implants: An experimental study in dogs. Journal of Orthopaedic Research.

Association. 2010;**47**:180

**140**

1996;**7**:300-310

1996;**14**:343-350

[144] Lew DP, Waldvogel FA. Osteomyelitis. The Lancet. 2004;**364**:369-379

[145] Bohner M, Lemaitre J, Van Landuyt P, Zambelli PY, Merkle HP, Gander B. Gentamicin-loaded hydraulic calcium phosphate bone cement as antibiotic delivery system. Journal of Pharmaceutical Sciences. 1997;**86**:565-572

[146] Alt V, Bechert T, Steinrcke P, Wagener M, Seidel P, Dingeldein E, et al. An in vitro assessment of the antibacterial properties and cytotoxicity of nanoparticulate silver bone cement. Biomaterials. 2004;**25**:4383-4391

[147] Itokazu M, Wenyi Y, Aoki T, Ohara A, Kato N. Synthesis of antibiotic-loaded interporous hydroxyapatite blocks by vacuum method and in vitro drug release testing. Biomaterials. 1998;**19**:817-819

[148] Penner MJ, Masri BA, Duncan CP. Elution characteristics of vancomycin and tobramycin combined in acrylic bone-cement. The Journal of Arthroplasty. 1996;**11**:939-944

[149] Dion A, Langman M, Hall G, Filiaggi M. Vancomycin release behaviour from amorphous calcium polyphosphate matrices intended for osteomyelitis treatment. Biomaterials. 2005;**26**:7276-7285

[150] Jiang PJ, Patel S, Gbureck U, Grover LM. A comparison of the efficacy of hydroxyapatite based cements and gels as drug delivery matrices. Key Engineering Materials. 2008;**93**:327-330

[151] Frutos P, Pena E, Frutos G, Barrales-Rienda JM. Release of gentamicin sulphate from modified commercial bone cement. Effect of (2-hydroxyethyl methacrylate) comonomer and poly(N-vinyl-2 pyrrolidone) additive on release mechanism and kinetics. Biomaterials. 2002;**23**:3787-3797

[152] Schnieders J, Gbureck U, Thull R, Kissel T. Controlled release of gentamicin from calcium phosphate– poly(lactic acid-co-glycolic acid) composite bone cement. Biomaterials. 2006;**27**:4239-4249

[153] Anttipoika I, Josefsson G, Konttinen Y, Lidgren L, Santavirta S, Sanzen L. Hip-arthroplasty infection— Current concepts. Acta Orthopaedica Scandinavica. 1990;**61**:163-169

[154] Barralet JE, Aldred S, Wright AJ, Coombes AGA. In vitro behavior of albumin-loaded carbonate hydroxyapatite gel. Journal of Biomedical Materials Research. 2002;**60**:360-367

[155] Gitelis S, Brebach GT. The treatment of chronic osteomyelitis with a biodegradable antibiotic-impregnated implant. Journal of Orthopaedic Surgery. 2002;**10**:53-60

[156] Hendriks JGE, van Horn JR, van der Mei HC, Busscher HJ. Backgrounds of antibiotic-loaded bone cement and prosthesis-related infection. Biomaterials. 2004;**25**:545-556

[157] Pitto RP, Spika IA. Antibioticloaded bone cement spacers in two-stage management of infected total knee arthroplasty, International Orthopaedics. 2004;**28**:129-133

[158] Durbhakula SM, Czajka J, Fuchs MD, Uhl RL. Spacer endoprosthesis for the treatment of infected total hip arthroplasty. The Journal of Arthroplasty. 2004;**19**:760-767

[159] Ewald A, Gluckermann SK, Thull R, Gbureck U. Antimicrobial titanium/silver PVD coatings on titanium, BioMedical Engineering Online. 2006;**5**:22

[160] Yorganci K, Krepel C, Weigelt JA, Edmiston CE. Activity of antibacterial impregnated central venous catheters

against Klebsiella pneumonia, Intensive Care Medicine. 2002;**28**:438-442

[161] Davenport K, Keeley FX. Evidence for the use of silver-alloy-coated urethral catheters. The Journal of Hospital Infection. 2005;**60**:298-303

[162] Cook G, Costerton JW, Darouiche RO. Direct confocal microscopy studies of the bacterial colonization in vitro of a silvercoated heart valve sewing cuff. The International Journal of Antimicrobial Agents. 2000;**13**:169-173

[163] Boswald M, Lugauer S, Regenfus A, Braun GG, Martus P, Geis C, et al. Reduced rates of catheter-associated infection by use of a new silverimpregnated central venous catheter. Infection. 1999;**27**:56-60

[164] Blaker JJ, Nazhat SN, Boccaccini AR. Development and characterisation of silver-doped bioactive glass-coated sutures for tissue engineering and wound healing applications. Biomaterials. 2003;**25**:1319-1329

[165] Clement JL, Jarrett PS. Antibacterial silver. Metal-Based Drugs. 1994;**1**:467-482

[166] Mahabole M, Aiyer R, Ramakrishna C, Sreedhar B, Khairnar R. Synthesis, characterization and gas sensing property of hydroxyapatite ceramic. The Bulletin of Materials Science. 2005;**28**:535-545

[167] Kim TN, Feng QL, Kim JO, Wu J, Wang H, Chen GC, et al. Antimicrobial effects of metal ions (Ag<sup>+</sup> , Cu2+, Zn2+) in hydroxyapatite. Journal of Materials Science. Materials in Medicine. 1998;**9**: 129-134

[168] Chen W, Oh S, Ong AP, Oh N, Liu Y, Courtney HS, et al. Antibacterial and osteogenic properties of silvercontaining hydroxyapatite coatings produced using a sol gel process. Journal of Biomedical Materials Research. Part A. 2007;**82**:899-906

[169] Chen W, Liu Y, Courtney HS, Bettenga M, Agrawal CM, Bumgardner JD, et al. In vitro antibacterial and biological properties of magnetron co-sputtered silvercontaining hydroxyapatite coating. Biomaterials. 2006;**27**:5512-5517

[170] Clupper DC, Hench LL. Bioactive response of Ag-doped tape cast bioglass-45S5 following heat treatment. Journal of Materials Science. Materials in Medicine. 2001;**12**:917-921

[171] Sharma VK, Yngard RA, Lin Y. Silver nanoparticles: green synthesis and their antimicrobial activities. Advances in Colloid and Interface Science. 2009;**145**:83-96

[172] Ciceo Lucacel R, Hulpus AO, Simon V, Ardelean I. Structural characterization of phosphate glasses doped with silver. Journal of Non-Crystalline Solids. 2009;**355**:425-429

[173] Hollinger MA. Toxicological aspects of topical silver pharmaceuticals. Critical Reviews in Toxicology. 1996;**26**:255-260

[174] Faust RA. Toxicity Summary for Silver, Prepared for the Oak Ridge Reservation Environmental Restoration Programme. Oak Ridge National Laboratory, US Department of Energy; 1992. Available from: http://rais.ornl. gov/tox/profiles/silver\_c\_V1.html [Accessed: 19 August 2019]

[175] Berger TJ, Spadaro JA, Chapin SE, Becher RO. Electrically generated silver ions: quantitative effects on bacterial and mammalian cells. Antimicrobial Agents and Chemotherapy. 1976;**9**:357-358

[176] Williams RL, Doherty PJ, Vince DG, Grashoff GJ, Williams DF. The biocompatibility of silver. Critical

**143**

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

> Cu2+ and Cr3+ incorporated into calcium phosphate bone cements. PLoS One.

[185] Trombetta R, Inzana J, Schwarz E, Kates S, Awad H. 3D printing of calcium phosphate ceramics for bone tissue engineering and drug delivery. Annals of Biomedical Engineering. 2017;**45**:23-44

[186] Ngo T, Kashani A, Imbalzano G,

manufacturing (3D printing): A review of materials, methods, applications and challenges. Composites Part B: Engineering. 2018;**143**:172-196

[188] Luo Y, Lode A, Sonntag F, Nies B, Gelinsky M. Well-ordered biphasic calcium phosphate–alginate scaffolds fabricated by multi-channel 3D plotting under mild conditions. Journal of Materials Chemistry B. 2013;**1**:4088

Nguyena K, Hui D. Additive

[187] Akkinenia A, Luoa Y, Schumacher M, Nies B, Lode A, Gelinsky M. 3D plotting of growth factor loaded calcium phosphate cement scaffolds. Acta Biomaterialia.

2015;**27**:264-274

[184] Bergmanna C, Lindnera M, Zhangb W, Koczura K, Kirstena A, Telleb R, et al. 3D printing of bone substitute implants using calcium phosphate and bioactive glasses. Journal of the European Ceramic Society.

2017;**12**:e0182109

2010;**30**:2563-2567

Reviews in Biocompatibility.

[177] Ewald A, Hösel D, Patel S, Grover LM, Jake E, Barralet JE, et al. Silver-doped calcium phosphate cements with antimicrobial activity. Acta Biomaterialia. 2011;**7**:4064-4070

[178] Rau JV, Fosca M, Graziani V, Egorov AA, Zobkov V, Fedotov A, et al. Silver-doped calcium phosphate bone cements with antibacterial properties. Journal of Functional Biomaterials.

[179] Kulanthaivel S, Roy B, Agarwal T, Giri S, Pramanik K, Pal K, et al. Cobalt doped proangiogenic hydroxyapatite for bone tissue engineering application. Materials Science & Engineering. C, Materials for Biological Applications.

1989;**5**:221-223

2016;**7**:10

2016;**58**:648-658

[180] Birgani ZT, Gharraee N, Malhotra A, van Blitterswijk CA, Habibovic P. Combinatorial incorporation of fluoride and cobalt ions into calcium phosphates to stimulate osteogenesis and angiogenesis. Biomedical Materials (Bristol, England). 2016;**11**:015020

[181] Barralet J, Gbureck U,

2009;**15**:1601-1609

2017;**73**:99-110

[183] Bernhardt A, Schamel M, Gbureck U, Gelinsky M. Osteoclastic differentiation and resorption is modulated by bioactive metal ions Co2+,

Habibovic P, Vorndran E, Gerard C, Doillon CJ. Angiogenesis in calcium phosphate scaffolds by inorganic copper ion release. Tissue Engineering. Part A.

[182] Schamel M, Bernhardt A, Quade M, Wuerkner C, Gbureck U, Moseke C, et al. Cu2+, Co2+ and Cr3+ doping of a calcium phosphate cement influences materials properties and response of human mesenchymal stromal cells. Materials Science and Engineering: C.

*Calcium Phosphate Cements in Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.89131*

Reviews in Biocompatibility. 1989;**5**:221-223

*Contemporary Topics about Phosphorus in Biology and Materials*

of Biomedical Materials Research.

[169] Chen W, Liu Y, Courtney HS,

Bumgardner JD, et al. In vitro antibacterial and biological properties of magnetron co-sputtered silvercontaining hydroxyapatite coating. Biomaterials. 2006;**27**:5512-5517

[170] Clupper DC, Hench LL. Bioactive response of Ag-doped tape cast

bioglass-45S5 following heat treatment. Journal of Materials Science. Materials

[171] Sharma VK, Yngard RA, Lin Y. Silver nanoparticles: green synthesis and their antimicrobial activities. Advances in Colloid and Interface Science.

[172] Ciceo Lucacel R, Hulpus AO, Simon V, Ardelean I. Structural characterization of phosphate glasses doped with silver. Journal of Non-Crystalline Solids. 2009;**355**:425-429

[173] Hollinger MA. Toxicological

pharmaceuticals. Critical Reviews in

[174] Faust RA. Toxicity Summary for Silver, Prepared for the Oak Ridge Reservation Environmental Restoration Programme. Oak Ridge National Laboratory, US Department of Energy; 1992. Available from: http://rais.ornl. gov/tox/profiles/silver\_c\_V1.html [Accessed: 19 August 2019]

[175] Berger TJ, Spadaro JA, Chapin SE, Becher RO. Electrically generated silver ions: quantitative effects on bacterial and mammalian cells. Antimicrobial

Agents and Chemotherapy.

[176] Williams RL, Doherty PJ, Vince DG, Grashoff GJ, Williams DF. The biocompatibility of silver. Critical

1976;**9**:357-358

aspects of topical silver

Toxicology. 1996;**26**:255-260

in Medicine. 2001;**12**:917-921

2009;**145**:83-96

Part A. 2007;**82**:899-906

Bettenga M, Agrawal CM,

against Klebsiella pneumonia, Intensive

[161] Davenport K, Keeley FX. Evidence for the use of silver-alloy-coated urethral catheters. The Journal of Hospital Infection. 2005;**60**:298-303

[163] Boswald M, Lugauer S, Regenfus A, Braun GG, Martus P, Geis C, et al. Reduced rates of catheter-associated infection by use of a new silverimpregnated central venous catheter.

[164] Blaker JJ, Nazhat SN, Boccaccini AR. Development and characterisation of silver-doped bioactive glass-coated sutures for tissue engineering and wound healing applications. Biomaterials.

[165] Clement JL, Jarrett PS. Antibacterial silver. Metal-Based Drugs.

Ramakrishna C, Sreedhar B, Khairnar R. Synthesis, characterization and gas sensing property of hydroxyapatite ceramic. The Bulletin of Materials

[167] Kim TN, Feng QL, Kim JO, Wu J, Wang H, Chen GC, et al. Antimicrobial

in hydroxyapatite. Journal of Materials Science. Materials in Medicine. 1998;**9**:

[168] Chen W, Oh S, Ong AP, Oh N, Liu Y, Courtney HS, et al. Antibacterial and osteogenic properties of silvercontaining hydroxyapatite coatings produced using a sol gel process. Journal

, Cu2+, Zn2+)

[166] Mahabole M, Aiyer R,

Science. 2005;**28**:535-545

effects of metal ions (Ag<sup>+</sup>

Care Medicine. 2002;**28**:438-442

[162] Cook G, Costerton JW, Darouiche RO. Direct confocal microscopy studies of the bacterial colonization in vitro of a silvercoated heart valve sewing cuff. The International Journal of Antimicrobial

Agents. 2000;**13**:169-173

Infection. 1999;**27**:56-60

2003;**25**:1319-1329

1994;**1**:467-482

**142**

129-134

[177] Ewald A, Hösel D, Patel S, Grover LM, Jake E, Barralet JE, et al. Silver-doped calcium phosphate cements with antimicrobial activity. Acta Biomaterialia. 2011;**7**:4064-4070

[178] Rau JV, Fosca M, Graziani V, Egorov AA, Zobkov V, Fedotov A, et al. Silver-doped calcium phosphate bone cements with antibacterial properties. Journal of Functional Biomaterials. 2016;**7**:10

[179] Kulanthaivel S, Roy B, Agarwal T, Giri S, Pramanik K, Pal K, et al. Cobalt doped proangiogenic hydroxyapatite for bone tissue engineering application. Materials Science & Engineering. C, Materials for Biological Applications. 2016;**58**:648-658

[180] Birgani ZT, Gharraee N, Malhotra A, van Blitterswijk CA, Habibovic P. Combinatorial incorporation of fluoride and cobalt ions into calcium phosphates to stimulate osteogenesis and angiogenesis. Biomedical Materials (Bristol, England). 2016;**11**:015020

[181] Barralet J, Gbureck U, Habibovic P, Vorndran E, Gerard C, Doillon CJ. Angiogenesis in calcium phosphate scaffolds by inorganic copper ion release. Tissue Engineering. Part A. 2009;**15**:1601-1609

[182] Schamel M, Bernhardt A, Quade M, Wuerkner C, Gbureck U, Moseke C, et al. Cu2+, Co2+ and Cr3+ doping of a calcium phosphate cement influences materials properties and response of human mesenchymal stromal cells. Materials Science and Engineering: C. 2017;**73**:99-110

[183] Bernhardt A, Schamel M, Gbureck U, Gelinsky M. Osteoclastic differentiation and resorption is modulated by bioactive metal ions Co2+, Cu2+ and Cr3+ incorporated into calcium phosphate bone cements. PLoS One. 2017;**12**:e0182109

[184] Bergmanna C, Lindnera M, Zhangb W, Koczura K, Kirstena A, Telleb R, et al. 3D printing of bone substitute implants using calcium phosphate and bioactive glasses. Journal of the European Ceramic Society. 2010;**30**:2563-2567

[185] Trombetta R, Inzana J, Schwarz E, Kates S, Awad H. 3D printing of calcium phosphate ceramics for bone tissue engineering and drug delivery. Annals of Biomedical Engineering. 2017;**45**:23-44

[186] Ngo T, Kashani A, Imbalzano G, Nguyena K, Hui D. Additive manufacturing (3D printing): A review of materials, methods, applications and challenges. Composites Part B: Engineering. 2018;**143**:172-196

[187] Akkinenia A, Luoa Y, Schumacher M, Nies B, Lode A, Gelinsky M. 3D plotting of growth factor loaded calcium phosphate cement scaffolds. Acta Biomaterialia. 2015;**27**:264-274

[188] Luo Y, Lode A, Sonntag F, Nies B, Gelinsky M. Well-ordered biphasic calcium phosphate–alginate scaffolds fabricated by multi-channel 3D plotting under mild conditions. Journal of Materials Chemistry B. 2013;**1**:4088

**145**

**Chapter 8**

**Abstract**

wider biomedical applications.

dental resin, remineralization

are provided in Refs. [1–4].

**1. Introduction**

Amorphous Calcium Phosphate

*Diane R. Bienek, Anthony A. Giuseppetti and Drago Skrtic*

As biocompatible and osteo-inductive precursor to biological apatite formation, amorphous calcium phosphate (ACP) resorbs at the rate that closely coincides with the rate of new bone formation and is more osteo-conductive than its crystalline counterpart. In addition, in the oral environment, ACP intrinsically provides a protracted supply of the remineralizing calcium and phosphate ions needed for regeneration of mineral lost to tooth decay. These features make ACP composites a strong remineralizing tool at the site of caries attack. Our group has been on the forefront of the research on bioactive, remineralizing, polymeric ACP-based dental materials for over two decades. This entry describes methods for filler, polymer, and composite fabrication and a battery of physicochemical and biological tests involved in evaluation of ACP-based restoratives. Also presented is our most recent design of ACP remineralizing composites with added antimicrobial capability that shows promise for extended dental and, potentially,

**Keywords:** amorphous calcium phosphate, bioactivity, dental composite,

Due to their abundance in nature as phosphate minerals and their existence in living organisms, calcium phosphates (CaPs) are of special significance to humans. CaPs are involved in normal (bones, teeth, antlers) as well as pathological (calcification/mineral deposition in soft tissues) mineralization. Both normal and pathological calcifications represent in vivo crystallization. In contrast, dental caries and osteoporosis are manifestations of in vivo dissolution where less soluble CaPs are being replaced by the more soluble ones. The extensive overviews of the current knowledge on CaP structures, properties, and biomedical and dental importance

Amorphous calcium phosphates (ACPs) are unique members of CaP family with glass-like physical properties and variable chemistry [5]. Evidence of ACPs being an integral mineral component of bones and teeth is, however, ambiguous [6, 7]. Consequently, ACPs are considered the transient precursors in biomineralization [8–11] with the majority of information on their possible roles in biomineralization

originating from the synthetic and/or in vitro studies [12].

as Bioactive Filler in Polymeric

Dental Composites

#### **Chapter 8**

## Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites

*Diane R. Bienek, Anthony A. Giuseppetti and Drago Skrtic*

#### **Abstract**

As biocompatible and osteo-inductive precursor to biological apatite formation, amorphous calcium phosphate (ACP) resorbs at the rate that closely coincides with the rate of new bone formation and is more osteo-conductive than its crystalline counterpart. In addition, in the oral environment, ACP intrinsically provides a protracted supply of the remineralizing calcium and phosphate ions needed for regeneration of mineral lost to tooth decay. These features make ACP composites a strong remineralizing tool at the site of caries attack. Our group has been on the forefront of the research on bioactive, remineralizing, polymeric ACP-based dental materials for over two decades. This entry describes methods for filler, polymer, and composite fabrication and a battery of physicochemical and biological tests involved in evaluation of ACP-based restoratives. Also presented is our most recent design of ACP remineralizing composites with added antimicrobial capability that shows promise for extended dental and, potentially, wider biomedical applications.

**Keywords:** amorphous calcium phosphate, bioactivity, dental composite, dental resin, remineralization

#### **1. Introduction**

Due to their abundance in nature as phosphate minerals and their existence in living organisms, calcium phosphates (CaPs) are of special significance to humans. CaPs are involved in normal (bones, teeth, antlers) as well as pathological (calcification/mineral deposition in soft tissues) mineralization. Both normal and pathological calcifications represent in vivo crystallization. In contrast, dental caries and osteoporosis are manifestations of in vivo dissolution where less soluble CaPs are being replaced by the more soluble ones. The extensive overviews of the current knowledge on CaP structures, properties, and biomedical and dental importance are provided in Refs. [1–4].

Amorphous calcium phosphates (ACPs) are unique members of CaP family with glass-like physical properties and variable chemistry [5]. Evidence of ACPs being an integral mineral component of bones and teeth is, however, ambiguous [6, 7]. Consequently, ACPs are considered the transient precursors in biomineralization [8–11] with the majority of information on their possible roles in biomineralization originating from the synthetic and/or in vitro studies [12].

The use of CaPs in dentistry is roughly a century old with the first scientific article published in 1925 [3]. Majority of the literature on dental use of CaPs focuses on crystalline CaPs. Today, two ACP-based remineralization systems have been commercialized as a toothpaste: a casein phosphopeptide-stabilized ACP and an unstabilized ACP. Other ACP applications include various biocompatible formulations where ACP acts as anticariogenic, remineralizing agent (polymeric composites, chewing gums, sugar confections, bleaching gels, and/or mouth rinses).

In this chapter, we present an overview of our group's up-to-date work on ACPbased dental composites with the emphasis on fabrication and characterization of ACP filler, fine-tuning of polymer phase, and physicochemical, mechanical, and biological evaluation of the ensuing polymeric ACP composites. Attention is also given to our most recent efforts to develop bioactive composites with both remineralizing and antimicrobial (AM) capabilities.

#### **2. ACP-based polymeric dental composites**

The approaches to synthesize ACP include precipitation from supersaturated calcium and phosphate solutions (wet synthesis), spray drying of acidified aqueous solutions of soluble CaPs, precipitation from nonaqueous solutions and solvents (sol-gel technique), and dry chemical techniques (mechanochemical methods, high-energy processing at elevated temperatures). Morphology, chemical composition, atomic structure, thermal properties, mechanical properties, and kinetics of ACP's transformation into crystalline CaPs depend on the preparation method [5]. ACPs fabricated by wet chemical method typically have a relatively constant chemical composition, suggesting the existence of well-defined structural units/ clusters [12, 13]. ACPs are thermodynamically unstable in solutions and convert spontaneously into crystalline CaPs, predominantly apatite. As result of ACP to apatite transformation, both crystallinity and Ca/P ratios of the solid increase with time. The release of remineralizing ions that accompany ACP's transformation in solutions, coupled with their excellent biocompatibility and bioresorbability, makes ACP-based dental materials a powerful remineralizing tool. However, the unique benefit of ACP-based materials, i.e., their sustained remineralization capability, is compromised by the inherently low strength and toughness of ACP composites due to the uncontrolled agglomeration of ACP particles and enhanced water sorption (WS) on exposure to aqueous environment [14]. To overcome these deficiencies, we explored multiple approaches to reduce the heterogeneity of filler particle size distribution (PSD) and better control the stability of the ACP/resin interface. These attempts are described in Sections 2.1–2.3.

#### **2.1 ACP filler: synthesis, characterization, and surface modification**

ACPs utilized in our studies were synthesized according to the protocol originally proposed [15]. It was modified to include ab initio addition of cations or nonionic surfactants to Ca reactant and anionic surfactant or polymer to PO4 reactant. In brief, ACP precipitated instantly upon mixing equal volumes of the reactant solutions {Ca reactant, 80 mmol/L Ca(NO3)2; PO4 reactant, 54 mmol/L Na2HPO4 + 2 mol% Na4P2O7 (to inhibit the precipitation of apatite simultaneously with ACP)} at 23°C and pH ≥ 8.5. After filtering, solid was subsequently washed with ice-cold ammoniated water, then with acetone, freeze-dried, and lyophilized. Dry, as-synthesized ACP (as-ACP) was kept in dessicator under vacuum before being subjected to additional treatments: silanization [16], grinding [17], or mechanical milling [14]. Before being utilized for composite fabrication, ACPs were

**147**

**Figure 1.**

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites*

characterized (validated) by multiple screenings. The amorphousness of the solids was confirmed by X-ray diffraction (XRD) analysis and Fourier-transform infrared (FTIR) spectroscopy. PSD of ACP dispersed in isopropanol was determined by laser light scattering, and their morphology/topology of gold-sputtered specimens was assessed by scanning electron microscopy (SEM). ACP's water content was measured by thermogravimetric analysis (TGA), and Ca/PO4 ratio was calculated from atomic emission spectroscopy data obtained upon dissolving ACP powder in HCl. Typical XRD pattern, FTIR spectrum, and SEM image of ACP are presented in **Figure 1**. More detailed description of ACP preparatory protocols and filler's characterization/validation is provided [18]. The same reference is also a good source of information on methods and techniques utilized in preparation and evaluation of copolymers and composites and their physicochemical, mechanical, and biological

Introducing additives during ACP synthesis and/or applying secondary treatments (**Table 1**) was expected to yield less clustered dry powders (more homogeneous and narrower PSD) and lead to better dispersion of ACP in the resin. Experimental details on surface modification protocols are provided in [16, 19]. More intimate ACP/resin contact is expected to enhance hydrolytic stability of ACP through lowering the WS, thus improving the mechanical performance of

The coprecipitation of Ag and Fe phosphates and premature ACP to apatite conversion in the presence of Fe2+ and Fe3+ disqualified these cations from further evaluation. Results of PSD analysis and TGA results obtained with surface-modified ACPs and strength testing of model 2,2-bis[p-(2′-hydroxy-3′ methacryloxypropoxy) phenyl] propane (Bis-GMA)-based composites employing

Regardless of the type of treatment, practically all ACPs preserved their heterogeneous PSDs. However, values of the median diameter (dm), i.e., the midpoint of the volume size distribution, of Zn-ACP and Al-ACP were significantly lower than the dms of all other ACPs subjected to surface modification and marginally lower than the dms of ground- (g-ACP) or milled- (m-ACP) solids. By comparing the PSD histograms and the corresponding SEM images (not shown here), it is concluded that the degree of particle agglomeration in Zn-ACP and Al-ACP has been modestly reduced—the quantification of the effect was not attempted. A similar conclusion was made after examining the PSDs and SEMs of g- and m-ACP vs. the control.

*Representative XRD pattern (lower left corner) and FTIR spectrum (upper right corner) of ACP powder* 

*DOI: http://dx.doi.org/10.5772/intechopen.86640*

assessments discussed in Sections 2.2 and 2.3.

these fillers are summarized in **Figure 2**.

*(center) with the corresponding SEM image (background).*

composites.

#### *Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites DOI: http://dx.doi.org/10.5772/intechopen.86640*

characterized (validated) by multiple screenings. The amorphousness of the solids was confirmed by X-ray diffraction (XRD) analysis and Fourier-transform infrared (FTIR) spectroscopy. PSD of ACP dispersed in isopropanol was determined by laser light scattering, and their morphology/topology of gold-sputtered specimens was assessed by scanning electron microscopy (SEM). ACP's water content was measured by thermogravimetric analysis (TGA), and Ca/PO4 ratio was calculated from atomic emission spectroscopy data obtained upon dissolving ACP powder in HCl. Typical XRD pattern, FTIR spectrum, and SEM image of ACP are presented in **Figure 1**. More detailed description of ACP preparatory protocols and filler's characterization/validation is provided [18]. The same reference is also a good source of information on methods and techniques utilized in preparation and evaluation of copolymers and composites and their physicochemical, mechanical, and biological assessments discussed in Sections 2.2 and 2.3.

Introducing additives during ACP synthesis and/or applying secondary treatments (**Table 1**) was expected to yield less clustered dry powders (more homogeneous and narrower PSD) and lead to better dispersion of ACP in the resin. Experimental details on surface modification protocols are provided in [16, 19]. More intimate ACP/resin contact is expected to enhance hydrolytic stability of ACP through lowering the WS, thus improving the mechanical performance of composites.

The coprecipitation of Ag and Fe phosphates and premature ACP to apatite conversion in the presence of Fe2+ and Fe3+ disqualified these cations from further evaluation. Results of PSD analysis and TGA results obtained with surface-modified ACPs and strength testing of model 2,2-bis[p-(2′-hydroxy-3′ methacryloxypropoxy) phenyl] propane (Bis-GMA)-based composites employing these fillers are summarized in **Figure 2**.

Regardless of the type of treatment, practically all ACPs preserved their heterogeneous PSDs. However, values of the median diameter (dm), i.e., the midpoint of the volume size distribution, of Zn-ACP and Al-ACP were significantly lower than the dms of all other ACPs subjected to surface modification and marginally lower than the dms of ground- (g-ACP) or milled- (m-ACP) solids. By comparing the PSD histograms and the corresponding SEM images (not shown here), it is concluded that the degree of particle agglomeration in Zn-ACP and Al-ACP has been modestly reduced—the quantification of the effect was not attempted. A similar conclusion was made after examining the PSDs and SEMs of g- and m-ACP vs. the control.

#### **Figure 1.**

*Representative XRD pattern (lower left corner) and FTIR spectrum (upper right corner) of ACP powder (center) with the corresponding SEM image (background).*

*Contemporary Topics about Phosphorus in Biology and Materials*

alizing and antimicrobial (AM) capabilities.

attempts are described in Sections 2.1–2.3.

**2.1 ACP filler: synthesis, characterization, and surface modification**

ACPs utilized in our studies were synthesized according to the protocol originally proposed [15]. It was modified to include ab initio addition of cations or nonionic surfactants to Ca reactant and anionic surfactant or polymer to PO4 reactant. In brief, ACP precipitated instantly upon mixing equal volumes of the reactant solutions {Ca reactant, 80 mmol/L Ca(NO3)2; PO4 reactant, 54 mmol/L Na2HPO4 + 2 mol% Na4P2O7 (to inhibit the precipitation of apatite simultaneously with ACP)} at 23°C and pH ≥ 8.5. After filtering, solid was subsequently washed with ice-cold ammoniated water, then with acetone, freeze-dried, and lyophilized. Dry, as-synthesized ACP (as-ACP) was kept in dessicator under vacuum before being subjected to additional treatments: silanization [16], grinding [17], or

mechanical milling [14]. Before being utilized for composite fabrication, ACPs were

**2. ACP-based polymeric dental composites**

The use of CaPs in dentistry is roughly a century old with the first scientific article published in 1925 [3]. Majority of the literature on dental use of CaPs focuses on crystalline CaPs. Today, two ACP-based remineralization systems have been commercialized as a toothpaste: a casein phosphopeptide-stabilized ACP and an unstabilized ACP. Other ACP applications include various biocompatible formulations where ACP acts as anticariogenic, remineralizing agent (polymeric composites, chewing gums, sugar confections, bleaching gels, and/or mouth rinses).

In this chapter, we present an overview of our group's up-to-date work on ACPbased dental composites with the emphasis on fabrication and characterization of ACP filler, fine-tuning of polymer phase, and physicochemical, mechanical, and biological evaluation of the ensuing polymeric ACP composites. Attention is also given to our most recent efforts to develop bioactive composites with both reminer-

The approaches to synthesize ACP include precipitation from supersaturated calcium and phosphate solutions (wet synthesis), spray drying of acidified aqueous solutions of soluble CaPs, precipitation from nonaqueous solutions and solvents (sol-gel technique), and dry chemical techniques (mechanochemical methods, high-energy processing at elevated temperatures). Morphology, chemical composition, atomic structure, thermal properties, mechanical properties, and kinetics of ACP's transformation into crystalline CaPs depend on the preparation method [5]. ACPs fabricated by wet chemical method typically have a relatively constant chemical composition, suggesting the existence of well-defined structural units/ clusters [12, 13]. ACPs are thermodynamically unstable in solutions and convert spontaneously into crystalline CaPs, predominantly apatite. As result of ACP to apatite transformation, both crystallinity and Ca/P ratios of the solid increase with time. The release of remineralizing ions that accompany ACP's transformation in solutions, coupled with their excellent biocompatibility and bioresorbability, makes ACP-based dental materials a powerful remineralizing tool. However, the unique benefit of ACP-based materials, i.e., their sustained remineralization capability, is compromised by the inherently low strength and toughness of ACP composites due to the uncontrolled agglomeration of ACP particles and enhanced water sorption (WS) on exposure to aqueous environment [14]. To overcome these deficiencies, we explored multiple approaches to reduce the heterogeneity of filler particle size distribution (PSD) and better control the stability of the ACP/resin interface. These

**146**


#### **Table 1.**

*Surface modification of ACP: ab initio additives and secondary treatments.*

The significant increase in dm of polymer-ACPs, on the other hand, could be attributed to the effect similar to "polymer bridging" seen in apatite/high-molecularweight (Mw) polyacrylate system [18]. In theory, poly(ethylene oxide) (PEO)/ACP interactions are controlled by the conformational changes in the adsorbed polymer, collision rate between the particles, and/or aggregate breakup due to fluid shear [20]. However, the exact controlling mechanism is yet to be determined.

The total content of surface-bound and/or structural water in all ACPs was unaffected by the treatment (on average (16.0 ± 1.2) mass%) and compared well with the control ((15.8 ± 3.9) mass%). Ideally, the lower intrinsic water content associated with less agglomerated ACP filler is desired to ensure favorable WS and ion release. WS of dental composites is generally controlled by structure/ composition of the resin matrix. In case of ACP composites, the hydrophilic filler increases the amount of water absorbed. In ACP/Bis-GMA composites [21], filler/ resin interface is inundated by the numerous voids that may enhance water diffusion and hydration of the filler. These processes are better controlled in m-ACP composites [14].

A simple BFS [22] screening of Bis-GMA-based ACP composites (**Figure 2**) revealed the following order of mechanical stability upon extended (up to 3 months) exposure to aqueous environment: (m-ACP ≥ g-ACP ≥ Zn-ACP = Zr-ACP = silanized ACP) ≥ control (unmodified) ACP ≥ (Si-ACP ≥ surfactant-ACPs) ≥ (Al-ACP = PEO-ACP). The observed decrease in BFS of ACP composites can, generally, be attributed to either reduction in ACP's intactness caused by spatial changes that occur in parallel to ion release, the internal ACP to apatite conversion, or increased WS. The experimental finding that, after aqueous immersion, the BFS of Al-ACP and PEO-ACP composites deteriorated independently of their PSDs (dm of PEO-ACP was more than six times larger than that of Al-ACP, while their BFS was practically identical) suggests that fillers PSD has only a minor, if any, role in composite's ability to resist plasticization/degradation upon water exposure.

**149**

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites*

Our findings contradict an earlier report [23] that PEO strengthened the interface with ACP interface. Better understanding of the observed behavior may require further in-depth mechanical testing and water sorption/desorption evaluations. It appears prudent to consider approaches apart from the undertaken filler treatments, such as choice of monomers utilized in resin formulations, to design

*Effect of additives/treatments on PSD and water content of ACP solids and the biaxial flexure strength (BFS) following the aqueous immersion of composite specimens formulated with these fillers and bis-GMA resin. Indicated are mean values of minimum three independent runs and standard deviations (SD, represented by* 

**2.2 Resin matrix: fine-tuning a means of improving copolymer properties**

The experimental resins were formulated from the commercially available base, diluent, and adhesive monomers activated for light-, chemical-, and/or dual (light and chemical)-cure (LC, CC, and DC, respectively). Monomers utilized in resin formulations and the ranges of their concentrations are listed in **Table 2** (indicated acronyms will be used throughout this chapter). All commercial monomers were

In LC formulations, all components were blended together and stirred at room temperature until reaching a uniform consistency. In CC formulations, monomers were first combined, their mixture homogenized and then divided into two equal parts by mass. The components of the CC initiating system are added separately to each monomer mixture and stirred magnetically until blended fully. In DC systems, LC and CC step were combined. LC and DC preparations required the use of yellow

Copolymer specimens were prepared by packing resins into Teflon molds, covering each side of the mold with Mylar film and a glass slide, clamping assembly together, and, for LC and DC specimens, curing each side of the assembly with visible light for 2 minutes. Physicochemical tests routinely performed with copolymers involved assessment of degree of vinyl conversion (DVC; near-IR spectroscopy),

Methacrylate-based dental resins typically entail high-viscosity base monomer

(due to its larger molecular volume decreases polymerization shrinkage (PS) and enhances modulus of cured copolymer) and low-viscosity diluent monomer (due to its smaller molecular volume and greater flexibility enhances DVC and handling properties). Majority of contemporary dental resins is Bis-GMA-/ TEGDMA-based. Known drawbacks of Bis-GMA-based systems are relatively low DVCs, high PS, and susceptibility to hydrolytic and enzymatic degradation upon

*DOI: http://dx.doi.org/10.5772/intechopen.86640*

composites with optimized performance.

**Figure 2.**

utilized as received, without any additional purification.

*bars). Control ACP: as-ACP made without additives or a secondary treatment.*

light to prevent the premature photopolymerization.

WS (gravimetric measurements), and BFS.

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites DOI: http://dx.doi.org/10.5772/intechopen.86640*

**Figure 2.**

*Contemporary Topics about Phosphorus in Biology and Materials*

**Additives Expected interactions/effect(s)**

and their coalescence with water Nonionic-Triton100, Tween 80,

Grinding Particle size reduced by friction

*Surface modification of ACP: ab initio additives and secondary treatments.*

*Acronyms are defined in the appended list of abbreviations.*

The significant increase in dm of polymer-ACPs, on the other hand, could be attributed to the effect similar to "polymer bridging" seen in apatite/high-molecularweight (Mw) polyacrylate system [18]. In theory, poly(ethylene oxide) (PEO)/ACP interactions are controlled by the conformational changes in the adsorbed polymer, collision rate between the particles, and/or aggregate breakup due to fluid shear

Anionic-Zonyl FSP As dispersants, surfactants prevent agglomeration of ACP particles

(APTMS, MPTMS) Silane agents stimulate formation of durable bonds at ACP/resin

Ball milling Particle size reduced and PSD homogenized via high-energy impact

PEO (Mw 8 K, 100 K, 1000 K) Polymer stabilizes ACP particles via multiple chelation

interface

and collision

ACP/cation interactions controlled by the cation's ionic potential

The total content of surface-bound and/or structural water in all ACPs was unaffected by the treatment (on average (16.0 ± 1.2) mass%) and compared well with the control ((15.8 ± 3.9) mass%). Ideally, the lower intrinsic water content associated with less agglomerated ACP filler is desired to ensure favorable WS and ion release. WS of dental composites is generally controlled by structure/ composition of the resin matrix. In case of ACP composites, the hydrophilic filler increases the amount of water absorbed. In ACP/Bis-GMA composites [21], filler/ resin interface is inundated by the numerous voids that may enhance water diffusion and hydration of the filler. These processes are better controlled in m-ACP

A simple BFS [22] screening of Bis-GMA-based ACP composites (**Figure 2**)

revealed the following order of mechanical stability upon extended (up to 3 months) exposure to aqueous environment: (m-ACP ≥ g-ACP ≥ Zn-ACP = Zr-ACP = silanized ACP) ≥ control (unmodified) ACP ≥ (Si-ACP ≥ surfactant-ACPs) ≥ (Al-ACP = PEO-ACP). The observed decrease in BFS of ACP composites can, generally, be attributed to either reduction in ACP's intactness caused by spatial changes that occur in parallel to ion release, the internal ACP to apatite conversion, or increased WS. The experimental finding that, after aqueous immersion, the BFS of Al-ACP and PEO-ACP composites deteriorated independently of their PSDs (dm of PEO-ACP was more than six times larger than that of Al-ACP, while their BFS was practically identical) suggests that fillers PSD has only a minor, if any, role in composite's ability to resist plasticization/degradation upon water exposure.

[20]. However, the exact controlling mechanism is yet to be determined.

**148**

composites [14].

Cations Ag+

Surfactants\*

Zonyl FSN Polymer\*

Treatments Silanization\*

*\**

**Table 1.**

Zr4+

, Fe2+, Zn2+, Al3+, Fe3+, Si4+,

*Effect of additives/treatments on PSD and water content of ACP solids and the biaxial flexure strength (BFS) following the aqueous immersion of composite specimens formulated with these fillers and bis-GMA resin. Indicated are mean values of minimum three independent runs and standard deviations (SD, represented by bars). Control ACP: as-ACP made without additives or a secondary treatment.*

Our findings contradict an earlier report [23] that PEO strengthened the interface with ACP interface. Better understanding of the observed behavior may require further in-depth mechanical testing and water sorption/desorption evaluations. It appears prudent to consider approaches apart from the undertaken filler treatments, such as choice of monomers utilized in resin formulations, to design composites with optimized performance.

#### **2.2 Resin matrix: fine-tuning a means of improving copolymer properties**

The experimental resins were formulated from the commercially available base, diluent, and adhesive monomers activated for light-, chemical-, and/or dual (light and chemical)-cure (LC, CC, and DC, respectively). Monomers utilized in resin formulations and the ranges of their concentrations are listed in **Table 2** (indicated acronyms will be used throughout this chapter). All commercial monomers were utilized as received, without any additional purification.

In LC formulations, all components were blended together and stirred at room temperature until reaching a uniform consistency. In CC formulations, monomers were first combined, their mixture homogenized and then divided into two equal parts by mass. The components of the CC initiating system are added separately to each monomer mixture and stirred magnetically until blended fully. In DC systems, LC and CC step were combined. LC and DC preparations required the use of yellow light to prevent the premature photopolymerization.

Copolymer specimens were prepared by packing resins into Teflon molds, covering each side of the mold with Mylar film and a glass slide, clamping assembly together, and, for LC and DC specimens, curing each side of the assembly with visible light for 2 minutes. Physicochemical tests routinely performed with copolymers involved assessment of degree of vinyl conversion (DVC; near-IR spectroscopy), WS (gravimetric measurements), and BFS.

Methacrylate-based dental resins typically entail high-viscosity base monomer (due to its larger molecular volume decreases polymerization shrinkage (PS) and enhances modulus of cured copolymer) and low-viscosity diluent monomer (due to its smaller molecular volume and greater flexibility enhances DVC and handling properties). Majority of contemporary dental resins is Bis-GMA-/ TEGDMA-based. Known drawbacks of Bis-GMA-based systems are relatively low DVCs, high PS, and susceptibility to hydrolytic and enzymatic degradation upon


#### **Table 2.**

*Methacrylate monomers and the components of polymerization-initiating systems utilized to fabricate the experimental resins.*

exposure to oral fluids. In our resin fine-tuning studies, we have been exploring the utility of EBPADMA and UDMA as the alternative base monomers and EHMA, HEMA, and PEG-U as the alternative diluent monomers. These studies are expected to yield valuable information on the interplay between the resin's hydrophilicity/hydrophobicity, DVC, and mechanical stability and their effects on thermodynamic stability and mechanical performance of ACP composites fabricated with these resins.

Results of DVC and BFS screenings of Bis-GMA-, EBPADMA-, and UDMAbased resins are compiled in **Figure 3**. The DVC values attained in our experimental formulations (ranging from 81.4 to 88.2%) with varying diluent and/or adhesive monomers (**Table 2**) significantly exceeded the typical DVC values reported for Bis-GMA/TEGDMA copolymers (60–70% [24]).

DVCs attained in Bis-GMA (81.4 ± 2.5%), EBPADMA (84.8 ± 5.5%), and UDMA (88.2 ± 2.2%) copolymers were significantly different (*P* < 0.05; Tukey post hoc test) only for UDMA vs. Bis-GMA systems. The BFS values, apparently increasing in going from Bis-GMA to EBPADMA to UDMA copolymers (R<sup>2</sup> = 0.9685), were significantly different (*P* < 0.05) only between UDMA and Bis-GMA copolymers. WS (data not shown) was not significantly affected by the resin's compositional changes. In all three groups, maximum WS (values ranged from 2.3 to 3.0 mass%) was achieved within 2 weeks of aqueous immersion. Slight differences between the measured WS values could be related primarily to the inclusion of more hydrophilic (HEMA, TEGDMA) or less hydrophilic (EHMA, PEG-U) monomers in the resin matrix. It is, however, highly significant that all systems maintained the WS profiles that support a sustained ion release from the composites (see Section 2.3).

**151**

**Figure 4.**

**Figure 3.**

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites*

**2.3 ACP composites: fabrication, physicochemical, mechanical, and biological** 

To fabricate composites, resin (60 mass%) and ACP filler (40 mass%) were combined by hand spatulation. The homogenized paste was kept overnight under vacuum to eliminate the entrapped air before being utilized for preparation of testing specimens by employing the procedures identical to those for the preparation of copolymer specimens. Besides DVC, BFS, WS, and PS testing, physicochemical and biological evaluation of composites also entailed polymerization shrinkage stress (PSS), shear bond strength (SBS) to dentin, ion release kinetics, leachability of unreacted species, remineralization efficacy, and in vitro cytotoxicity tests. DVCs (**Figure 4**) of Zr-ACP composites formulated with various resins significantly (ANOVA, Tukey test; *P* < 0.05) increased in going from Bis-GMA to EBPADMA to UDMA matrix. Higher DVC in UDMA systems could possibly be explained by the higher reactivity of UDMA than Bis-GMA and/or EBPADMA [25]. Resin composition, however, had no effect on the BFS of composites being ~50% lower than the BFS of their copolymer counterparts (compare BFS data in **Figure 4** vs. **Figure 3**). The reduction in BFS in UDMA-based composites (1.7%) was significantly lower (*P* < 0.05) than reduction in BFS in Bis-GMA- to EBPADMA-based composites (10.0 and 4.4%, respectively). This finding suggests that random

*DVC and BFS of LC binary, ternary, and quaternary bis-GMA-, EBPADMA-, and UDMA-based copolymers. Indicated are mean values + SD (represented by bars) for n* ≥ *24 (DVC) and n* ≥ *12 (BFS).*

*DVC and BFS of LC Zr-ACP bis-GMA-, EBPADMA-, and UDMA-based composites. Indicated are mean* 

*values + SD (represented by bars) for n* ≥ *24 (DVC) and n* ≥ *12 (BFS).*

*DOI: http://dx.doi.org/10.5772/intechopen.86640*

**assessments**

#### **2.3 ACP composites: fabrication, physicochemical, mechanical, and biological assessments**

To fabricate composites, resin (60 mass%) and ACP filler (40 mass%) were combined by hand spatulation. The homogenized paste was kept overnight under vacuum to eliminate the entrapped air before being utilized for preparation of testing specimens by employing the procedures identical to those for the preparation of copolymer specimens. Besides DVC, BFS, WS, and PS testing, physicochemical and biological evaluation of composites also entailed polymerization shrinkage stress (PSS), shear bond strength (SBS) to dentin, ion release kinetics, leachability of unreacted species, remineralization efficacy, and in vitro cytotoxicity tests.

DVCs (**Figure 4**) of Zr-ACP composites formulated with various resins significantly (ANOVA, Tukey test; *P* < 0.05) increased in going from Bis-GMA to EBPADMA to UDMA matrix. Higher DVC in UDMA systems could possibly be explained by the higher reactivity of UDMA than Bis-GMA and/or EBPADMA [25]. Resin composition, however, had no effect on the BFS of composites being ~50% lower than the BFS of their copolymer counterparts (compare BFS data in **Figure 4** vs. **Figure 3**). The reduction in BFS in UDMA-based composites (1.7%) was significantly lower (*P* < 0.05) than reduction in BFS in Bis-GMA- to EBPADMA-based composites (10.0 and 4.4%, respectively). This finding suggests that random

#### **Figure 3.**

*Contemporary Topics about Phosphorus in Biology and Materials*

Base monomers 2,2-Bis[p-(2′-hydroxy-3′-methacryloxypropoxy) phenyl] propane

**Component Chemical name Acronym Content** 

Ethoxylated bisphenol A dimethacrylate EBPADMA ≤68.4 Urethane dimethacrylate UDMA ≤92.4

Ethyl-α-hydroxymethyl acrylate EHMA ≤29.2 2-hydroxyethyl methacrylate HEMA ≤30.4 Poly(ethylene glycol)-extended UDMA PEG-U ≤29.1 Triethylene glycol dimethacrylate TEGDMA ≤50.2

Methacryloyloxyethyl phthalate MEP ≤5.0

Camphorquinone CQ 0.2 Ethyl-4 N,N-dimethylamino benzoate 4EDMAB 0.8

Benzoyl peroxide BPO 2 2,2-Dihydroxyethyl-p-toluidine DHEPT 1

**(mass %)**

Bis-GMA ≥68.4

exposure to oral fluids. In our resin fine-tuning studies, we have been exploring the utility of EBPADMA and UDMA as the alternative base monomers and EHMA, HEMA, and PEG-U as the alternative diluent monomers. These studies are expected to yield valuable information on the interplay between the resin's hydrophilicity/hydrophobicity, DVC, and mechanical stability and their effects on thermodynamic stability and mechanical performance of ACP composites

*Methacrylate monomers and the components of polymerization-initiating systems utilized to fabricate the* 

Results of DVC and BFS screenings of Bis-GMA-, EBPADMA-, and UDMAbased resins are compiled in **Figure 3**. The DVC values attained in our experimental formulations (ranging from 81.4 to 88.2%) with varying diluent and/or adhesive monomers (**Table 2**) significantly exceeded the typical DVC values reported for

DVCs attained in Bis-GMA (81.4 ± 2.5%), EBPADMA (84.8 ± 5.5%), and UDMA (88.2 ± 2.2%) copolymers were significantly different (*P* < 0.05; Tukey post hoc test) only for UDMA vs. Bis-GMA systems. The BFS values, apparently increasing in going from Bis-GMA to EBPADMA to UDMA copolymers

 = 0.9685), were significantly different (*P* < 0.05) only between UDMA and Bis-GMA copolymers. WS (data not shown) was not significantly affected by the resin's compositional changes. In all three groups, maximum WS (values ranged from 2.3 to 3.0 mass%) was achieved within 2 weeks of aqueous immersion. Slight differences between the measured WS values could be related primarily to the inclusion of more hydrophilic (HEMA, TEGDMA) or less hydrophilic (EHMA, PEG-U) monomers in the resin matrix. It is, however, highly significant that all systems maintained the WS profiles that support a sustained ion release from the

**150**

(R<sup>2</sup>

fabricated with these resins.

Diluent monomers

Adhesive monomers

**Table 2.**

*experimental resins.*

Initiators *Light cure*

composites (see Section 2.3).

Bis-GMA/TEGDMA copolymers (60–70% [24]).

*Chemical cure*

*DVC and BFS of LC binary, ternary, and quaternary bis-GMA-, EBPADMA-, and UDMA-based copolymers. Indicated are mean values + SD (represented by bars) for n* ≥ *24 (DVC) and n* ≥ *12 (BFS).*

#### **Figure 4.**

*DVC and BFS of LC Zr-ACP bis-GMA-, EBPADMA-, and UDMA-based composites. Indicated are mean values + SD (represented by bars) for n* ≥ *24 (DVC) and n* ≥ *12 (BFS).*

distribution of Zr-ACP agglomerates [21], rather than the resin matrix composition, controls the mechanical performance of composites.

Generally, WS {(3.3–3.8) mass%}, SBS {(15.3–17.5) MPa}, and PS {(6.5– 7.0) vol%} results revealed no distinguishable differences between the Bis-GMA-, EBPADMA-, and UDMA-based composites. The observed increase in WS of composites vs. their copolymer counterparts (27–43%) was due to ACP's affinity to the environmental water. It is significant that, in terms of SBS, ACP composites performed as well as Sr-glass filled composites [17], thus providing the remineralizing component to the primary restorative function without impediment of shortand midterm dentin bonding. High PS (undesirable) seen in all three experimental groups go hand in hand with the high DVCs (desirable) attained in these systems. These high PS values are likely due to the increased hydrogen bonding occurring in all experimental matrices leading to the densification of polymerization [26]. It is particularly important that, in UDMA-based composites, high PS can be offset by a significant hygroscopic expansion (HE; up to 13.6 vol%; data not shown). The compensating effect of HE on PS has been demonstrated in [27–29]. We were unable to establish any correlation between the PS and PSS in our experimental systems. Although contributions on this subject in dental literature are considerable [30–33], there is a question whether processing factors such as configuration factor (C-factor) besides the filler type and its load level, resin composition, and polymerization mode control the composite performance [34].

Comparative kinetic study of ion release from Bis-GMA- and EBPADMAbased composites showed a systematic increase in solution Ca and PO4 concentrations with the increased filler level in the composite and with time of aqueous immersion. Values of the ion activity product (IAP) used to calculate the relative supersaturation of the solutions with respect to the stoichiometric hydroxyapatite (details on these calculations are provided in [35]) are presented in **Figure 5**. The ion release kinetics was practically identical in both types of composites. It was shown that a minimum of 35 mass% of ACP filler in composite is needed to create a sustained solution supersaturation inducive to apatite formation. The overall ion release kinetics in these systems was most likely controlled by the level of hydrophilic HEMA monomer in the resin phase. By more easily absorbing the water needed for ion diffusion, HEMA regulated the internal mineral saturation levels.

#### **Figure 5.**

*Mean value + SD (indicated by bars) of the ion activity product (IAP) of the solutions containing Ca and PO4 ions released from bis-GMA- and EBPADMA-based composites with various ACP levels (30, 35, or 40 mass%) at different times of immersion (1, 3, or 6 months) in saline solutions (23°C; continuous magnetic stirring). Number of repetitive experiments n = 3/group.*

**153**

**Figure 6.**

*Leachables (mean value + SD (represented by bars); n = 3/group) detected by 1*

*normalized with respect to the initial resin amount).*

*in acetone extracts from LC UPHM copolymers and their ACP composites (in composite series; values are* 

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites*

copolymers and their ACP composites, we have employed 1

In our studies, we conveniently use the DVC as a predictor of leachability (the higher the DVC, the lower the likelihood of unreacted monomers to leach) from the restorations. To quantify leaching from the UDMA/PEG-U/HEMA/MEP (UPHM)

The accelerated leaching was mimicked by using acetone as extraction medium. Results are presented in **Figure 6**. Leaching (expressed as a mass fraction of the initial amount incorporated in the resin and normalized to the equivalent amount of the resin for composite specimens) decreased in the following order: 4EDMA B > > MEP > UDMA > PEG-U > HEMA > CQ (undetectable). Introducing ACP filler into UPHM matrix did not have an impact on the leachability profile. In highly cross-linked UPHM matrix, polymer chain mobility remained small in both copolymers and composites, thus limiting the pathways for unreacted monomers to leach out from the specimens. The disproportionate leachability of 4EDMAB, a component of the LC initiator system, is most likely due to its excess relative to CQ (CQ:4EDMAB mass ratio 1:4), which appears to be fully consumed during polymerization and is undetectable in the extracts. Since HEMA is known for its increased toxicity and adverse side effects due to metabolizing to methacrylic acid [36], it is of high significance that the level of leachable HEMA from UPHM copolymers and composites is drastically (several orders of magnitude) lower than typically seen in HEMA-containing resins and composites. In addition, levels of leachable UDMA from UPHM matrices are more than twice lower than the amounts of leached UDMA from the experimental UDMA/TEGDMA formulations [37]. In the complex, high-DVC UPHM network, the mobility of the unreacted low-molecularweight HEMA is limited or reduced, and its elution to the environment is prevented. In vitro cytotoxicity screening should be performed in conjunction with the leachability tests to better predict the material's impact at cellular level. To mimic early ACP composite/cell interactions, we have performed the extraction experiments and assessed cell morphology and cell viability according to the recommended standards [38–40]. The osteoblast-like MC3T3-E1cells were seeded on ethanol-sterilized material disks, incubated in the extracts for 24 h, then evaluated in situ by optical microscopy, and assessed for cytotoxicity by tetrazolium-based {3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide; MTT) assay. Detailed protocols are described in [41]. The results of the cytotoxicity evaluation of UPHM composites are presented in **Figure 7**. Cell morphology (**Figure 7**, inlets) changed from polygonal, spread (cells exposed to pure medium and extracts from UPHM copolymer and compressed ACP filler) to spherical (cells incubated in

H-NMR spectroscopy.

*H-NMR spectroscopic method* 

*DOI: http://dx.doi.org/10.5772/intechopen.86640*

#### *Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites DOI: http://dx.doi.org/10.5772/intechopen.86640*

In our studies, we conveniently use the DVC as a predictor of leachability (the higher the DVC, the lower the likelihood of unreacted monomers to leach) from the restorations. To quantify leaching from the UDMA/PEG-U/HEMA/MEP (UPHM) copolymers and their ACP composites, we have employed 1 H-NMR spectroscopy. The accelerated leaching was mimicked by using acetone as extraction medium. Results are presented in **Figure 6**. Leaching (expressed as a mass fraction of the initial amount incorporated in the resin and normalized to the equivalent amount of the resin for composite specimens) decreased in the following order: 4EDMA B > > MEP > UDMA > PEG-U > HEMA > CQ (undetectable). Introducing ACP filler into UPHM matrix did not have an impact on the leachability profile. In highly cross-linked UPHM matrix, polymer chain mobility remained small in both copolymers and composites, thus limiting the pathways for unreacted monomers to leach out from the specimens. The disproportionate leachability of 4EDMAB, a component of the LC initiator system, is most likely due to its excess relative to CQ (CQ:4EDMAB mass ratio 1:4), which appears to be fully consumed during polymerization and is undetectable in the extracts. Since HEMA is known for its increased toxicity and adverse side effects due to metabolizing to methacrylic acid [36], it is of high significance that the level of leachable HEMA from UPHM copolymers and composites is drastically (several orders of magnitude) lower than typically seen in HEMA-containing resins and composites. In addition, levels of leachable UDMA from UPHM matrices are more than twice lower than the amounts of leached UDMA from the experimental UDMA/TEGDMA formulations [37]. In the complex, high-DVC UPHM network, the mobility of the unreacted low-molecularweight HEMA is limited or reduced, and its elution to the environment is prevented.

In vitro cytotoxicity screening should be performed in conjunction with the leachability tests to better predict the material's impact at cellular level. To mimic early ACP composite/cell interactions, we have performed the extraction experiments and assessed cell morphology and cell viability according to the recommended standards [38–40]. The osteoblast-like MC3T3-E1cells were seeded on ethanol-sterilized material disks, incubated in the extracts for 24 h, then evaluated in situ by optical microscopy, and assessed for cytotoxicity by tetrazolium-based {3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide; MTT) assay. Detailed protocols are described in [41]. The results of the cytotoxicity evaluation of UPHM composites are presented in **Figure 7**. Cell morphology (**Figure 7**, inlets) changed from polygonal, spread (cells exposed to pure medium and extracts from UPHM copolymer and compressed ACP filler) to spherical (cells incubated in

#### **Figure 6.**

*Contemporary Topics about Phosphorus in Biology and Materials*

controls the mechanical performance of composites.

ization mode control the composite performance [34].

distribution of Zr-ACP agglomerates [21], rather than the resin matrix composition,

Comparative kinetic study of ion release from Bis-GMA- and EBPADMAbased composites showed a systematic increase in solution Ca and PO4 concentrations with the increased filler level in the composite and with time of aqueous immersion. Values of the ion activity product (IAP) used to calculate the relative supersaturation of the solutions with respect to the stoichiometric hydroxyapatite (details on these calculations are provided in [35]) are presented in **Figure 5**. The ion release kinetics was practically identical in both types of composites. It was shown that a minimum of 35 mass% of ACP filler in composite is needed to create a sustained solution supersaturation inducive to apatite formation. The overall ion release kinetics in these systems was most likely controlled by the level of hydrophilic HEMA monomer in the resin phase. By more easily absorbing the water needed for ion diffusion, HEMA regulated the internal mineral

*Mean value + SD (indicated by bars) of the ion activity product (IAP) of the solutions containing Ca and PO4 ions released from bis-GMA- and EBPADMA-based composites with various ACP levels (30, 35, or 40 mass%) at different times of immersion (1, 3, or 6 months) in saline solutions (23°C; continuous magnetic stirring).* 

Generally, WS {(3.3–3.8) mass%}, SBS {(15.3–17.5) MPa}, and PS {(6.5– 7.0) vol%} results revealed no distinguishable differences between the Bis-GMA-, EBPADMA-, and UDMA-based composites. The observed increase in WS of composites vs. their copolymer counterparts (27–43%) was due to ACP's affinity to the environmental water. It is significant that, in terms of SBS, ACP composites performed as well as Sr-glass filled composites [17], thus providing the remineralizing component to the primary restorative function without impediment of shortand midterm dentin bonding. High PS (undesirable) seen in all three experimental groups go hand in hand with the high DVCs (desirable) attained in these systems. These high PS values are likely due to the increased hydrogen bonding occurring in all experimental matrices leading to the densification of polymerization [26]. It is particularly important that, in UDMA-based composites, high PS can be offset by a significant hygroscopic expansion (HE; up to 13.6 vol%; data not shown). The compensating effect of HE on PS has been demonstrated in [27–29]. We were unable to establish any correlation between the PS and PSS in our experimental systems. Although contributions on this subject in dental literature are considerable [30–33], there is a question whether processing factors such as configuration factor (C-factor) besides the filler type and its load level, resin composition, and polymer-

**152**

**Figure 5.**

*Number of repetitive experiments n = 3/group.*

saturation levels.

*Leachables (mean value + SD (represented by bars); n = 3/group) detected by 1 H-NMR spectroscopic method in acetone extracts from LC UPHM copolymers and their ACP composites (in composite series; values are normalized with respect to the initial resin amount).*

#### **Figure 7.**

*Viability of osteoblast-like cells upon exposure to the extracts from ACP UPHM composite, commercial composite, medium + surfactant, and medium only. Shown are mean values = SD (indicated by bars) for three repetitive runs in each group. Optical micrographs indicating typical changes in cell morphology upon exposure to the extracts from composites (left side) and pure medium (right side).*

extracts from the experimental ACP UPHM composite and commercial control). Viability of contracted cells (i.e., those exposed to extracts from both composites) was reduced approximately twofold compared to the polygonal cell incubated in medium only. Similar morphological changes and slow cell proliferation observed in bone-regenerating hydrogels [42] are reportedly linked to the alterations in cell nucleus and mineralization of osteoprogenitors [43]. While fuller understanding of the cellular effects is certainly needed, it is encouraging that reduction in cell viability seen with our experimental composites was comparable with the commercial control.

Remineralization efficacy studies were undertaken to demonstrate the ability of ACP composites to regenerate demineralized tooth structures. Details on the microradiographic in vitro evaluations of the changes in mineral content of the demineralized bovine and/or human enamel subjected to aggressive acid attacks, representative prolonged exposure in oral milieu, are provided in [44, 45], respectively. These studies indicate that lost tooth mineral is indeed regenerated upon application of ACP composites (**Figure 8**). Mineral recovery with ACP Bis-GMA-based composites significantly exceeded the remineralization effect of

#### **Figure 8.**

*Mineral recovery (or loss) following the pH regimens representing pH cycling in oral environment. Shown are mean values ± SD (indicated by bars) obtained from a minimum of eight micrographic images/experimental group.*

**155**

materials [53–56].

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites*

**2.4 Adding antimicrobial functionality to dental restoratives**

mers, using *Streptococcus mutans* (planktonic and biofilm).

testing, specimens were degassed for ≥5 days.

the glass-filled control. On average, more than threefold increase of mineral content attained with ACP EBPADMA-based composites compared to F-releasing control was, however, statistically insignificant due to a large data scattering. It is particularly important that the regenerative action took place throughout the lesions in contrast to F-supported repair which is typically limited to a subsurface region.

To prolong service life, the next generation of polymeric dental restoratives is envisioned to possess bioactive properties. Today, developments of dental materials with embedded AM properties are relatively frequently reported. One of the first quaternary ammonium (QA) methacrylates integrated into Bis-GMA/TEGDMA matrix has been reported 20 years ago [46]. More

recently, two polymerizable ionic dimethacrylates (2-(methacryloyloxy)-N-(2- (methacryloyloxy)ethyl)-N,N-dimethylethan-1-aminium bromide (IDMA1) and N,N′-([1,1′-biphenyl]-2,2′-diylbis(methylene))bis(2-(methacryloyloxy)-N,Ndimethylethan-1-aminium) bromide (IDMA2)) were proposed for use in dental applications [47, 48]. Purity of the synthesized materials was not reported, and the structural characterizations appeared incomplete. The IDMAs have been validated by nuclear magnetic resonance, mass spectroscopy, and FTIR spectroscopies [49, 50]. These studies also included direct contact cytotoxicity testing at biologically relevant concentrations. To date, AM assessments of the fully charac-

For the advancement of Class V restoratives, our goal was to evaluate the AM properties of purified/validated IDMA1 and IDMA2 integrated in dental copoly-

IDMA copolymer disks (6 mm diameter) were fabricated by adding IDMA1 or IDMA2 (10 wt %) to light-activated UDMA/PEG-U/EHMA (hereafter UPE) resin. For both planktonic and biofilm testing, UPE disks were used as a negative control group. After fabrication, all copolymer disks were subjected to aqueous extraction (72 h, 37°C). After drying, specimens were sterilized using an Anprolene gas sterilization chamber (Andersen Products, Inc., Haw River, NC). Prior to bacterial

A bioluminescent *S. mutans* strain JM10 [51] (derivative of wild-type UA159) was used to assess the antibacterial properties, the IDMAs, employing a real-time bioluminescent assay developed by Florez et al. [52]. In UPE resin matrix, 10% IDMA1 reduced (P ≤ 0.005) the colonization of *S. mutans* biofilms threefold (**Figure 9**). The initial report describing the synthesis of IDMA1 and incorporation into Bis-GMA-TEGDMA resin conducted *S. mutans* testing with phosphate-buffered saline (i.e., no extrinsic proteins) [48]. Therein, they report reduced bacterial colonization for IDMA1 concentrations of 10, 20, or 30%. In the presence [53, 54] or absence [55] of ACP, others have demonstrated that IDMA1 fabricated with Bis-GMA/TEGDMA results in a modest (1.4–1.7-fold decrease in metabolic activity of biofilm) AM function in a protein-rich environment. To enhance AM activity, IDMA1 in combination with silver nanoparticles and/or nano-ACP was integrated into dental

AM activity of IDMA2 has not been previously reported. Compared to the UPE control resin, our experiments demonstrated that IDMA2-UPE copolymer reduced (*P* ≤ 0.005) the colonization of *S. mutans* biofilms 1.6-fold (**Figure 9**). However, the AM activity of IDMA2 was less (*P* ≤ 0.05) than that observed with IDMA1. This difference may be due to steric hindrance resulting from the presence of aromatic rings in IDMA2. Further, it is noteworthy that IDMA2 can exist as a mixture of conformers. Thus, it is conceivable that the conformer (with decreased steric

*DOI: http://dx.doi.org/10.5772/intechopen.86640*

terized IDMAs are remaining.

the glass-filled control. On average, more than threefold increase of mineral content attained with ACP EBPADMA-based composites compared to F-releasing control was, however, statistically insignificant due to a large data scattering. It is particularly important that the regenerative action took place throughout the lesions in contrast to F-supported repair which is typically limited to a subsurface region.

#### **2.4 Adding antimicrobial functionality to dental restoratives**

To prolong service life, the next generation of polymeric dental restoratives is envisioned to possess bioactive properties. Today, developments of dental materials with embedded AM properties are relatively frequently reported. One of the first quaternary ammonium (QA) methacrylates integrated into Bis-GMA/TEGDMA matrix has been reported 20 years ago [46]. More recently, two polymerizable ionic dimethacrylates (2-(methacryloyloxy)-N-(2- (methacryloyloxy)ethyl)-N,N-dimethylethan-1-aminium bromide (IDMA1) and N,N′-([1,1′-biphenyl]-2,2′-diylbis(methylene))bis(2-(methacryloyloxy)-N,Ndimethylethan-1-aminium) bromide (IDMA2)) were proposed for use in dental applications [47, 48]. Purity of the synthesized materials was not reported, and the structural characterizations appeared incomplete. The IDMAs have been validated by nuclear magnetic resonance, mass spectroscopy, and FTIR spectroscopies [49, 50]. These studies also included direct contact cytotoxicity testing at biologically relevant concentrations. To date, AM assessments of the fully characterized IDMAs are remaining.

For the advancement of Class V restoratives, our goal was to evaluate the AM properties of purified/validated IDMA1 and IDMA2 integrated in dental copolymers, using *Streptococcus mutans* (planktonic and biofilm).

IDMA copolymer disks (6 mm diameter) were fabricated by adding IDMA1 or IDMA2 (10 wt %) to light-activated UDMA/PEG-U/EHMA (hereafter UPE) resin. For both planktonic and biofilm testing, UPE disks were used as a negative control group. After fabrication, all copolymer disks were subjected to aqueous extraction (72 h, 37°C). After drying, specimens were sterilized using an Anprolene gas sterilization chamber (Andersen Products, Inc., Haw River, NC). Prior to bacterial testing, specimens were degassed for ≥5 days.

A bioluminescent *S. mutans* strain JM10 [51] (derivative of wild-type UA159) was used to assess the antibacterial properties, the IDMAs, employing a real-time bioluminescent assay developed by Florez et al. [52]. In UPE resin matrix, 10% IDMA1 reduced (P ≤ 0.005) the colonization of *S. mutans* biofilms threefold (**Figure 9**). The initial report describing the synthesis of IDMA1 and incorporation into Bis-GMA-TEGDMA resin conducted *S. mutans* testing with phosphate-buffered saline (i.e., no extrinsic proteins) [48]. Therein, they report reduced bacterial colonization for IDMA1 concentrations of 10, 20, or 30%. In the presence [53, 54] or absence [55] of ACP, others have demonstrated that IDMA1 fabricated with Bis-GMA/TEGDMA results in a modest (1.4–1.7-fold decrease in metabolic activity of biofilm) AM function in a protein-rich environment. To enhance AM activity, IDMA1 in combination with silver nanoparticles and/or nano-ACP was integrated into dental materials [53–56].

AM activity of IDMA2 has not been previously reported. Compared to the UPE control resin, our experiments demonstrated that IDMA2-UPE copolymer reduced (*P* ≤ 0.005) the colonization of *S. mutans* biofilms 1.6-fold (**Figure 9**). However, the AM activity of IDMA2 was less (*P* ≤ 0.05) than that observed with IDMA1. This difference may be due to steric hindrance resulting from the presence of aromatic rings in IDMA2. Further, it is noteworthy that IDMA2 can exist as a mixture of conformers. Thus, it is conceivable that the conformer (with decreased steric

*Contemporary Topics about Phosphorus in Biology and Materials*

*to the extracts from composites (left side) and pure medium (right side).*

extracts from the experimental ACP UPHM composite and commercial control). Viability of contracted cells (i.e., those exposed to extracts from both composites) was reduced approximately twofold compared to the polygonal cell incubated in medium only. Similar morphological changes and slow cell proliferation observed in bone-regenerating hydrogels [42] are reportedly linked to the alterations in cell nucleus and mineralization of osteoprogenitors [43]. While fuller understanding of the cellular effects is certainly needed, it is encouraging that reduction in cell viability seen with our experimental composites was comparable with the commer-

*Viability of osteoblast-like cells upon exposure to the extracts from ACP UPHM composite, commercial composite, medium + surfactant, and medium only. Shown are mean values = SD (indicated by bars) for three repetitive runs in each group. Optical micrographs indicating typical changes in cell morphology upon exposure* 

Remineralization efficacy studies were undertaken to demonstrate the ability of ACP composites to regenerate demineralized tooth structures. Details on the microradiographic in vitro evaluations of the changes in mineral content of the demineralized bovine and/or human enamel subjected to aggressive acid attacks, representative prolonged exposure in oral milieu, are provided in [44, 45], respectively. These studies indicate that lost tooth mineral is indeed regenerated upon application of ACP composites (**Figure 8**). Mineral recovery with ACP Bis-GMA-based composites significantly exceeded the remineralization effect of

*Mineral recovery (or loss) following the pH regimens representing pH cycling in oral environment. Shown are mean values ± SD (indicated by bars) obtained from a minimum of eight micrographic images/experimental* 

**154**

*group.*

**Figure 8.**

cial control.

**Figure 7.**

#### **Figure 9.**

*Streptococcus mutans biofilm growth by the experimental IDMAs-UPE (10 wt %) copolymers compared to UPE resin. Presented are mean values ± SD (indicated by bars) of 5 specimens/experimental group. a P ≤ 0.05 compared to IDMA2; <sup>b</sup> P ≤ 0.005 compared to UPE resin.*

hindrance) could exhibit greater AM activity. Future research that enables preparations, enriched for each conformer, is warranted to determine whether AM activity of IDMA2 can be enhanced.

For planktonic bacterial testing, *S. mutans* UA-159 (ATCC® 700610) cultures were established in Todd Hewitt broth (THB). Copolymer disks were seeded with *S. mutans* suspension (~3 × 107 colony-forming units/disk). To maximize the contact between the copolymer and bacteria, a second disk was placed atop. This assembly was incubated 2 h at 37°C in a 5% CO2 environment. The samples were then placed in 1 ml of THB and utilized to prepare 10-fold serial dilutions. Of the resulting suspensions, 100 μl were distributed onto the surface of THB agar plates. Colony-forming units were the unit of measure used to approximate the number of viable *S. mutans* cells after exposure to the IDMAs-UPE copolymers. Among the IDMA groups, the number of colony-forming units observed was not statistically different from one another or the UPE control resin (data not shown).

Altogether, the full AM potential of the IDMAs has not yet been realized. For example, the current IDMA restorative material formulations are likely to have charges randomly distributed throughout the material. It is conceivable that chemical engineering advances (i.e., material gradient or layers) could favor the orientation of the N+ charges while maintaining the material mechanical properties. Secondly, previous reports demonstrated that the presence of proteins dampens the AM capability of QA methacrylates (reviewed by [57]). To overcome this problem, a commercially available protein repellent, 2-methacryloyloxyethyl phosphorylcholine (MEPC), has been integrated into AM dental material [58]. However, because of possible binding of the remineralizing calcium ions by MEPC [59], MEPC may have a deleterious effect on the remineralization capacity of the restorative material. Generally, to date, there is a paucity of information regarding the nature of proteins binding to the surface of QA methacrylates. Elucidation of the protein-material interactions would yield valuable information to develop a strategy to maximize AM efficacy of materials with a chargebased AM mechanism of action. Further, these data could facilitate the advancement of materials, modified to target specific proteins (i.e., via molecular imprinting).

**157**

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites*

Bioactive, ACP-based polymeric dental composites ACP have tremendous appeal due to their intrinsic ability to regenerate lost tooth mineral. ACP composites release remineralizing Ca and PO4 ions in a sustained fashion, thus providing a long-term protection against acid challenges in the oral environment. However, with respect to the mechanical strength and toughness, ACP-based composites are inferior to silanized glass-filled counterparts and can only be used in nonstress-bearing applications. The uncontrolled ACP particle agglomeration typically leads to the poor ACP filler/resin interactions that destabilize filler/polymer interface. A desired, more homogeneous PSD of ACP filler is achievable via mechanical treatments, preferably high-energy milling. Surface modifications of ACP were proven less effective in that respect. With HEMA as diluent monomer, LC Bis-GMA-, EBPADMA-, and UDMA-based copolymers and their ACP composites typically attain high-DVC values that would suggest a minimal leaching of unreacted monomeric species from these materials. Expectedly, these high DVCs are accompanied with high PS. In UDMA-based systems, high PS is likely to be offset by a significant HE. High remineralization capabilities are achieved with EBPADMA or UDMA as base monomers and HEMA or EHMA plus PEG-U as diluent monomers. The most recent efforts in our group focus on design of bifunctional, remineralizing, and AM ACP composites that may find utility as new Class V restoratives. These materials are currently being assessed for their in vitro cytotoxicity and AM properties. The evaluations coupled with the leachability studies need to be completed before embarking on animal testing and/or human clinical trials involving AM ACP dental composites.

The work presented in this chapter was supported by the National Institute of Dental and Craniofacial Research (grants DE 13169 and DE 26122), American Dental Association (ADA), and ADA Foundation. Special thanks to collaborators (Sharukh S. Khajotia, Fernando L. Esteban Florez, and Rochelle D. Hiers) at the University of Oklahoma Health Sciences Center, College of Dentistry, for conducting biofilm testing. Donation of the monomers by Esstech, Essington, PA, is

*DOI: http://dx.doi.org/10.5772/intechopen.86640*

**3. Conclusions**

**Acknowledgements**

gratefully acknowledged.

**Conflict of interest**

Authors declare no conflict of interest.

ACP amorphous calcium phosphate ADA American Dental Association

APTMS 3-aminopropyltrimethoxysilane

Al-ACP aluminum-modified ACP

AM antimicrobial

as-ACP as synthesized ACP BFS biaxial flexure strength

ADAF American Dental Association Foundation

**A.Appendix: list of abbreviations**

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites DOI: http://dx.doi.org/10.5772/intechopen.86640*

#### **3. Conclusions**

*Contemporary Topics about Phosphorus in Biology and Materials*

of IDMA2 can be enhanced.

*S. mutans* suspension (~3 × 107

hindrance) could exhibit greater AM activity. Future research that enables preparations, enriched for each conformer, is warranted to determine whether AM activity

*Streptococcus mutans biofilm growth by the experimental IDMAs-UPE (10 wt %) copolymers compared to UPE resin. Presented are mean values ± SD (indicated by bars) of 5 specimens/experimental group. a*

*P ≤ 0.005 compared to UPE resin.*

For planktonic bacterial testing, *S. mutans* UA-159 (ATCC® 700610) cultures were established in Todd Hewitt broth (THB). Copolymer disks were seeded with

contact between the copolymer and bacteria, a second disk was placed atop. This assembly was incubated 2 h at 37°C in a 5% CO2 environment. The samples were then placed in 1 ml of THB and utilized to prepare 10-fold serial dilutions. Of the resulting suspensions, 100 μl were distributed onto the surface of THB agar plates. Colony-forming units were the unit of measure used to approximate the number of viable *S. mutans* cells after exposure to the IDMAs-UPE copolymers. Among the IDMA groups, the number of colony-forming units observed was not statistically

Altogether, the full AM potential of the IDMAs has not yet been realized. For example, the current IDMA restorative material formulations are likely to have charges randomly distributed throughout the material. It is conceivable that chemical engineering advances (i.e., material gradient or layers) could favor the orientation of

 charges while maintaining the material mechanical properties. Secondly, previous reports demonstrated that the presence of proteins dampens the AM capability of QA methacrylates (reviewed by [57]). To overcome this problem, a commercially available protein repellent, 2-methacryloyloxyethyl phosphorylcholine (MEPC), has been integrated into AM dental material [58]. However, because of possible binding of the remineralizing calcium ions by MEPC [59], MEPC may have a deleterious effect on the remineralization capacity of the restorative material. Generally, to date, there is a paucity of information regarding the nature of proteins binding to the surface of QA methacrylates. Elucidation of the protein-material interactions would yield valuable information to develop a strategy to maximize AM efficacy of materials with a chargebased AM mechanism of action. Further, these data could facilitate the advancement of materials, modified to target specific proteins (i.e., via molecular imprinting).

different from one another or the UPE control resin (data not shown).

colony-forming units/disk). To maximize the

*P ≤ 0.05* 

**156**

the N+

**Figure 9.**

*compared to IDMA2; <sup>b</sup>*

Bioactive, ACP-based polymeric dental composites ACP have tremendous appeal due to their intrinsic ability to regenerate lost tooth mineral. ACP composites release remineralizing Ca and PO4 ions in a sustained fashion, thus providing a long-term protection against acid challenges in the oral environment. However, with respect to the mechanical strength and toughness, ACP-based composites are inferior to silanized glass-filled counterparts and can only be used in nonstress-bearing applications. The uncontrolled ACP particle agglomeration typically leads to the poor ACP filler/resin interactions that destabilize filler/polymer interface. A desired, more homogeneous PSD of ACP filler is achievable via mechanical treatments, preferably high-energy milling. Surface modifications of ACP were proven less effective in that respect. With HEMA as diluent monomer, LC Bis-GMA-, EBPADMA-, and UDMA-based copolymers and their ACP composites typically attain high-DVC values that would suggest a minimal leaching of unreacted monomeric species from these materials. Expectedly, these high DVCs are accompanied with high PS. In UDMA-based systems, high PS is likely to be offset by a significant HE. High remineralization capabilities are achieved with EBPADMA or UDMA as base monomers and HEMA or EHMA plus PEG-U as diluent monomers. The most recent efforts in our group focus on design of bifunctional, remineralizing, and AM ACP composites that may find utility as new Class V restoratives. These materials are currently being assessed for their in vitro cytotoxicity and AM properties. The evaluations coupled with the leachability studies need to be completed before embarking on animal testing and/or human clinical trials involving AM ACP dental composites.

#### **Acknowledgements**

The work presented in this chapter was supported by the National Institute of Dental and Craniofacial Research (grants DE 13169 and DE 26122), American Dental Association (ADA), and ADA Foundation. Special thanks to collaborators (Sharukh S. Khajotia, Fernando L. Esteban Florez, and Rochelle D. Hiers) at the University of Oklahoma Health Sciences Center, College of Dentistry, for conducting biofilm testing. Donation of the monomers by Esstech, Essington, PA, is gratefully acknowledged.

#### **Conflict of interest**

Authors declare no conflict of interest.

#### **A.Appendix: list of abbreviations**



**159**

**Author details**

Diane R. Bienek, Anthony A. Giuseppetti and Drago Skrtic\*

\*Address all correspondence to: drago.skrtic@nist.gov

provided the original work is properly cited.

Research Division, American Dental Association Foundation, Frederick, MD, USA

© 2019 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium,

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites*

Zonyl FSN F(C2F4)*x*(C2H4O)1+*y*H; nonionic surfactant (x = 1–9; y = 0–25) Zonyl FSP F(C2F4)*x*(C2H4O)*y*HP(O)(ONH4)z; anionic surfactant (x = 1–7;

*DOI: http://dx.doi.org/10.5772/intechopen.86640*

Zr-ACP zirconia-modified ACP

(y + z) = 3)

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites DOI: http://dx.doi.org/10.5772/intechopen.86640*

Zr-ACP zirconia-modified ACP Zonyl FSN F(C2F4)*x*(C2H4O)1+*y*H; nonionic surfactant (x = 1–9; y = 0–25) Zonyl FSP F(C2F4)*x*(C2H4O)*y*HP(O)(ONH4)z; anionic surfactant (x = 1–7; (y + z) = 3)

#### **Author details**

*Contemporary Topics about Phosphorus in Biology and Materials*

DHEPT 2,2-dihydroxyethyl-p-toluidine

EBPADMA ethoxylated bisphenol A dimethacrylate 4EDMAB ethyl-4 N,N-dimethylamino benzoate EHMA ethyl-α-hydroxymethyl acrylate FTIR Fourier-transform infrared

ylethan-1-aminium bromide

MEPC 2-methacryloyloxyethyl phosphorylcholine MPTMS methacryloxypropyltrimethoxy silane

n number of specimens/experimental runs

MEP methacryloyloxyethyl phthalate

assay Mw molecular weight

PEG poly(ethylene glycol) PEG-U PEG-extended UDMA PEO poly(ethylene oxide) PS polymerization shrinkage PSD particle size distribution PSS polymerization shrinkage stress

QA quaternary ammonium SBS shear bond strength

SD standard deviation Si-ACP silica-modified ACP

THB Todd Hewitt broth

XRD X-ray diffraction WS water sorption Zn-ACP zinc-modified ACP

UDMA urethane dimethacrylate UPE UDMA/PEG-U/EHMA resin UPHM UDMA/PEG-U/HEMA/MEP resin

SEM scanning electron microscopy

TEGDMA triethylene glycol dimethacrylate TGA thermogravimetric analysis

Triton 100 C14H22O(C2H4O)10H; nonionic surfactant Tween 80 C24H43O6(C2H4O)20H; nonionic surfactant

IDMA2 N,N′-([1,1′-biphenyl]-2,2′-diylbis(methylene))bis(2-

DVC degree of vinyl conversion

HE hygroscopic expansion HEMA 2-hydroxyethyl methacrylate

IAP ion activity product IDMA ionic dimethacrylate

BPO benzoyl peroxide CaP calcium phosphate CC chemical-cure CQ camphorquinone

dm median diameter

g-ACP ground ACP

LC light-cure m-ACP milled ACP

DC dual-cure

Bis-GMA 2,2-bis[p-(2′-hydroxy-3′-methacryloxypropoxy) phenyl] propane

IDMA1 2-(methacryloyloxy)-N-(2-(methacryloyloxy)ethyl)-N,N-dimeth-

MTT 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide

(methacryloyloxy)-N,N-dimethylethan-1-aminium) bromide

**158**

Diane R. Bienek, Anthony A. Giuseppetti and Drago Skrtic\* Research Division, American Dental Association Foundation, Frederick, MD, USA

\*Address all correspondence to: drago.skrtic@nist.gov

© 2019 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

#### **References**

[1] Dorozhkin SV. Calcium orthophosphates (CaPO4): Occurrence and properties. Progress in Biomaterials. 2016;**5**:9-70. DOI: 10.1007/ s40204-015-0045-z

[2] Eliaz N, Metoki N. Calcium phosphate bioceramics: A review of their history, structure, properties, coating technologies and biomedical applications. Materials. 2017;**10**:334. DOI: 10.3390/ma10040334

[3] Dorozhkin SV. Calcium orthophosphates (Ca3PO4) and dentistry. Bioceramics Development and Applications. 2016;**6**(2). DOI: 10.4172/2090-5025: 1000096

[4] Bienek DR, Skrtic D. Utility of amorphous calcium phosphate-based scaffolds in dental/biomedical research. Biointerface Research in Applied Chemistry. 2017;**7**(1):1989-1994

[5] Dorozhkin SV. Amorphous calcium orthophosphates: Nature, chemistry and biomedical applications. International Journal of Materials and Chemistry. 2012;**2**(1):19-46. DOI: 10.5923/j. ijmc20120201.04

[6] Glimcher MJ, Bonar LC, Grynpas MD, Landis WJ, Roufosse AH. Recent studies of bone mineral: Is the amorphous calcium phosphate theory valid? Journal of Crystal Growth. 1981;**53**:100-119. DOI: 10.1016/9922-0248(81)90058-0

[7] Grynpas MD, Bonar LC, Glimcher MJ. Failure to detect an amorphous calcium phosphate solid phase in in bone mineral: A radial distribution function study. Calcified Tissue International. 1984;**36**:291-301. DOI: 10.1007/BF0240533

[8] Weiner S, Sagi I, Addadi L. Choosing the crystallization path less travelled.

Science. 2005;**309**:1027-1028. DOI: 10.1126.science.1114920

[9] Weiner S. Transient precursor strategy in mineral formation of bone. Bone. 2006;**39**:431-433. DOI: 10.1016/j. bone2006.02.058

[10] Mahamid J, Sharir A, Addadi L, Weiner S. Amorphous calcium phosphate is a major component in forming fin bones of zebrafish: Indications for an amorphous precursor phase. Proceedings of the National Academy of Sciences of the United States of America. 2008;**105**:12748-12753. DOI: 10.1073/ pnas.0806869105

[11] Beniash A, Metzler RA, Lam RSK, Gilbert PUPA. Transient amorphous calcium phosphate in forming enamel. Journal of Structural Biology. 2009;**166**:133-143. DOI: 10.1016/j. jsb2009.02.001

[12] Eanes ED. Amorphous calcium phosphate. In: Chow LC, Eanes ED, editors. Monographs in Oral Science. Vol. 18. Basel: Karger; 2001. pp. 130-147. ISBN: 9783318007046 3318007048

[13] Heughebaert JC, Montel G. Conversion of amorphous calcium phosphate into apatitic tricalcium phosphate. Calcified Tissue International. 1982;**34**:S103-S108. DOI: 10.1007/BF02426637

[14] O'Donnell JNR, Antonucci JM, Skrtic D. Illuminating the role of agglomerates on critical physicochemical properties of amorphous calcium phosphate composites. Journal of Composite Materials. 2008;**42**:2231-2246. DOI: 10.1177/0021998308094797

[15] Eanes ED, Gillessen IH, Posner AS. Intermediate states in the precipitation of hydroxyapatite. Nature.

**161**

cg034148g

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites*

[22] ASTMF394-78 (re-approved 1991). Standard test method for biaxial flexure

strength of ceramic substrates

10.1023/A:1024741426069

S142-9612(00)00067-3

2007;**21**(4):375-393. DOI: 10.1177/0885328206064823

10.1038/sj.bdj.4808379

[26] Skrtic D, Antonucci JM. Dental composites based on amorphous calcium phosphate–resin composition/ physicochemical properties study. Journal of Biomaterials Applications.

[27] Momoi Y, Mccabe JF. Hygroscopic expansion of resin based composites during 6 months water storage. British Dental Journal. 1994;**176**(3):91-96. DOI:

[28] Huang C, Tay FR, Cheung GSP, Kei LH, Wei SHY, Pashley DH. Hygroscopic

[29] Suiter EA, Watson LE, Tantbirojn C, Lou JSB, Versluis A. Effective expansion: Balance between shrinkage

expansion of compomer and a composite on artificial gap reduction. Journal of Dentistry. 2002;**30**(1):11-19. DOI: 10.1016/S0300-5712(01)00053-7

[23] Li Y, Weng W, Cheng K, Du P, Shen G, Wang J, et al. Preparation of amorphous calcium phosphate in the presence of poly (ethylene glycol). Journal of Materials Science Letters. 2003;**22**(14):1015-1016. DOI:

[24] Marovic D, Panduric V, Tarle Z, Ristic M, Sariri K, Demoli N, et al. Degree of conversion and microhardness of dental composite resin materials. Journal of Molecular Structure. 2013;**1044**:299-302. DOI: 10.1016/j.molstruc.2012.10.062

[25] Morgan DR, Kalachandra S, Shobha HK, Gunduz N, Stejskal EO. Analysis of a dimethacrylate copolymer (bis-GMA and TEGDMA) network by DSC and 13C solution and solid-state NMR spectroscopy. Biomaterials. 2000;**21**(18):1897-1903. DOI: 10.1016/

*DOI: http://dx.doi.org/10.5772/intechopen.86640*

[16] Antonucci JM, Skrtic D. Bioactive

[17] O'Donnell JNR, Skrtic D. Degree of vinyl conversion, polymerization shrinkage and stress development in experimental endodontic composite. Journal of Biomimetics Biomaterials and Tissue Engineering. 2009;**4**:1-12. DOI: 10.4028.www.scientific.net/JBBTE.4.1

[18] Skrtic D, Antonucci JM. Dental composites: Bioactive polymeric amorphous calcium phosphatebased. In: Encyclopedia of

Biomedical Polymers and Polymeric Biomaterials. New York: Taylor and Francis; 2015. pp. 2443-2262. DOI: 10.1081/E-EBPP-120051063

[19] Ofir PBY, Govrin-Lipman R, Garti N, Furedi-Milhofer H. The influence of polyelectrolytes on the formation and transformation of amorphous calcium phosphate. Crystal Growth & Design. 2004;**4**(1):177-183. DOI: 10.1021/

[20] Antonucci JM, Liu DW, Skrtic D. Amorphous calcium phosphatebased composites: Effect of

surfactants and poly (ethylene oxide) on filler and composite properties. Journal of Dispersion Science and Technology. 2007;**28**(5):819-824. DOI:

10.1080/01932690701346255

[21] Skrtic D, Antonucci JM, Eanes ED, Eidelman N. Dental composites based on hybrid and surface-modified amorphous calcium phosphates. Biomaterials. 2004;**25**:1141-1150. DOI: 10.1016/j.biomaterials.2003.08.001

1965;**208**(5008):365-367. DOI:

and biocompatible polymeric composites based on amorphous calcium phosphate. In: Ramalingam M, Tiwari A, Ramakrishna S, Kobayashi H, editors. Integrated Biomaterials for Medical Applications. Salem: Scrivener Publishing; 2012. pp. 67-119. DOI: 10.1002/9781118482513.ch3

10.1038/208365a0

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites DOI: http://dx.doi.org/10.5772/intechopen.86640*

1965;**208**(5008):365-367. DOI: 10.1038/208365a0

[16] Antonucci JM, Skrtic D. Bioactive and biocompatible polymeric composites based on amorphous calcium phosphate. In: Ramalingam M, Tiwari A, Ramakrishna S, Kobayashi H, editors. Integrated Biomaterials for Medical Applications. Salem: Scrivener Publishing; 2012. pp. 67-119. DOI: 10.1002/9781118482513.ch3

[17] O'Donnell JNR, Skrtic D. Degree of vinyl conversion, polymerization shrinkage and stress development in experimental endodontic composite. Journal of Biomimetics Biomaterials and Tissue Engineering. 2009;**4**:1-12. DOI: 10.4028.www.scientific.net/JBBTE.4.1

[18] Skrtic D, Antonucci JM. Dental composites: Bioactive polymeric amorphous calcium phosphatebased. In: Encyclopedia of Biomedical Polymers and Polymeric Biomaterials. New York: Taylor and Francis; 2015. pp. 2443-2262. DOI: 10.1081/E-EBPP-120051063

[19] Ofir PBY, Govrin-Lipman R, Garti N, Furedi-Milhofer H. The influence of polyelectrolytes on the formation and transformation of amorphous calcium phosphate. Crystal Growth & Design. 2004;**4**(1):177-183. DOI: 10.1021/ cg034148g

[20] Antonucci JM, Liu DW, Skrtic D. Amorphous calcium phosphatebased composites: Effect of surfactants and poly (ethylene oxide) on filler and composite properties. Journal of Dispersion Science and Technology. 2007;**28**(5):819-824. DOI: 10.1080/01932690701346255

[21] Skrtic D, Antonucci JM, Eanes ED, Eidelman N. Dental composites based on hybrid and surface-modified amorphous calcium phosphates. Biomaterials. 2004;**25**:1141-1150. DOI: 10.1016/j.biomaterials.2003.08.001

[22] ASTMF394-78 (re-approved 1991). Standard test method for biaxial flexure strength of ceramic substrates

[23] Li Y, Weng W, Cheng K, Du P, Shen G, Wang J, et al. Preparation of amorphous calcium phosphate in the presence of poly (ethylene glycol). Journal of Materials Science Letters. 2003;**22**(14):1015-1016. DOI: 10.1023/A:1024741426069

[24] Marovic D, Panduric V, Tarle Z, Ristic M, Sariri K, Demoli N, et al. Degree of conversion and microhardness of dental composite resin materials. Journal of Molecular Structure. 2013;**1044**:299-302. DOI: 10.1016/j.molstruc.2012.10.062

[25] Morgan DR, Kalachandra S, Shobha HK, Gunduz N, Stejskal EO. Analysis of a dimethacrylate copolymer (bis-GMA and TEGDMA) network by DSC and 13C solution and solid-state NMR spectroscopy. Biomaterials. 2000;**21**(18):1897-1903. DOI: 10.1016/ S142-9612(00)00067-3

[26] Skrtic D, Antonucci JM. Dental composites based on amorphous calcium phosphate–resin composition/ physicochemical properties study. Journal of Biomaterials Applications. 2007;**21**(4):375-393. DOI: 10.1177/0885328206064823

[27] Momoi Y, Mccabe JF. Hygroscopic expansion of resin based composites during 6 months water storage. British Dental Journal. 1994;**176**(3):91-96. DOI: 10.1038/sj.bdj.4808379

[28] Huang C, Tay FR, Cheung GSP, Kei LH, Wei SHY, Pashley DH. Hygroscopic expansion of compomer and a composite on artificial gap reduction. Journal of Dentistry. 2002;**30**(1):11-19. DOI: 10.1016/S0300-5712(01)00053-7

[29] Suiter EA, Watson LE, Tantbirojn C, Lou JSB, Versluis A. Effective expansion: Balance between shrinkage

**160**

*Contemporary Topics about Phosphorus in Biology and Materials*

Science. 2005;**309**:1027-1028. DOI:

[9] Weiner S. Transient precursor strategy in mineral formation of bone. Bone. 2006;**39**:431-433. DOI: 10.1016/j.

[10] Mahamid J, Sharir A, Addadi L, Weiner S. Amorphous calcium phosphate is a major component in forming fin bones of zebrafish: Indications for an amorphous precursor phase. Proceedings of the National Academy of Sciences of the United States of America. 2008;**105**:12748-12753. DOI: 10.1073/

[11] Beniash A, Metzler RA, Lam RSK, Gilbert PUPA. Transient amorphous calcium phosphate in forming

enamel. Journal of Structural Biology. 2009;**166**:133-143. DOI: 10.1016/j.

[12] Eanes ED. Amorphous calcium phosphate. In: Chow LC, Eanes ED, editors. Monographs in Oral Science. Vol. 18. Basel: Karger; 2001. pp. 130-147. ISBN: 9783318007046 3318007048

[13] Heughebaert JC, Montel G. Conversion of amorphous calcium phosphate into apatitic tricalcium phosphate. Calcified Tissue

10.1007/BF02426637

International. 1982;**34**:S103-S108. DOI:

[15] Eanes ED, Gillessen IH, Posner AS. Intermediate states in the precipitation

[14] O'Donnell JNR, Antonucci JM, Skrtic D. Illuminating the role of agglomerates on critical physicochemical properties of amorphous calcium phosphate composites. Journal of Composite Materials. 2008;**42**:2231-2246. DOI:

10.1177/0021998308094797

of hydroxyapatite. Nature.

10.1126.science.1114920

bone2006.02.058

pnas.0806869105

jsb2009.02.001

[1] Dorozhkin SV. Calcium

and properties. Progress in

[2] Eliaz N, Metoki N. Calcium phosphate bioceramics: A review of their history, structure, properties, coating technologies and biomedical applications. Materials. 2017;**10**:334.

DOI: 10.3390/ma10040334

[3] Dorozhkin SV. Calcium orthophosphates (Ca3PO4) and dentistry. Bioceramics Development and Applications. 2016;**6**(2). DOI: 10.4172/2090-5025: 1000096

[4] Bienek DR, Skrtic D. Utility of amorphous calcium phosphate-based scaffolds in dental/biomedical research. Biointerface Research in Applied Chemistry. 2017;**7**(1):1989-1994

[5] Dorozhkin SV. Amorphous calcium orthophosphates: Nature, chemistry and biomedical applications. International Journal of Materials and Chemistry. 2012;**2**(1):19-46. DOI: 10.5923/j.

[6] Glimcher MJ, Bonar LC, Grynpas MD, Landis WJ, Roufosse AH. Recent studies of bone mineral: Is the amorphous calcium phosphate theory valid? Journal of Crystal Growth. 1981;**53**:100-119. DOI: 10.1016/9922-0248(81)90058-0

[7] Grynpas MD, Bonar LC, Glimcher MJ. Failure to detect an amorphous calcium phosphate solid phase in in bone mineral: A radial distribution function study. Calcified Tissue International. 1984;**36**:291-301. DOI:

[8] Weiner S, Sagi I, Addadi L. Choosing the crystallization path less travelled.

ijmc20120201.04

10.1007/BF0240533

s40204-015-0045-z

**References**

orthophosphates (CaPO4): Occurrence

Biomaterials. 2016;**5**:9-70. DOI: 10.1007/

and hygroscopic expansion. Journal of Dental Research. 2016;**95**(5):543-549. DOI: 10.1177/0022034516633450

[30] Braga LL, Ferracane JL. Contraction stress related to degree of conversion and reaction kinetics. Journal of Dental Research. 2002;**81**(2):114-118. DOI: 10.1177/0810114

[31] Ferracane JL. Developing a more complex understanding of stresses produced in dental composites during polymerization. Dental Materials. 2005;**21**(1):36-42. DOI: 10.1016/j. dental.2004.10.004

[32] Kleverlaan CJ, Feizler AJ. Polymerization shrinkage and contraction stress of dental resin composites. Dental Materials. 2005;**21**(1):1150-1157. DOI: 10.1016/j. dental.2005.02.004

[33] Choi KK, Condon JR, Ferracane JL. The effects of adhesive thickness on polymerization contraction stress of composite. Journal of Dental Research. 2000;**79**(3):812-817. DOI: 10.1177/00220345000790030501

[34] Antonucci JM, Giuseppetti AA, O'Donnell JNR, Schumacher GE, Skrtic D. Polymerization stress development in dental composites: Effect of cavity design factor. Materials. 2009;**2**(1):169- 180. DOI: 10.3390/ma2010169

[35] O' Donnell JNR, Langhorst SE, Fow MD, Skrtic D. Light-cured dimethacrylate-based resins and their composites; Comparative study of mechanical strength, water sorption and ion release. Journal of Bioactive and Compatible Polymers. 2008;**23**:207-226. DOI: 10.1177/0883911508089932

[36] Durner J, Walther UI, Zaspel J, Hickel R, Reich FX. Metabolism of TEGDMA and HEMA in human cells. Biomaterials. 2010;**31**:818-823. DOI: 10.1016/j.biomaterials.2009.09.097

[37] Floyd CJE, Dickens SH. Network structure of bis-GMA- and UDMAbased resin systems. Dental Materials. 2006;**22**:1143-1149. DOI: 10.1016/j. dental.2005.10.009

[38] International Standard ISO10993- 5. Biological evaluation of medical devices- Part 5. Tests for cytotoxicity; In vitro methods. Geneva, Switzerland: International Organization for Standardization; 1992

[39] International Standard ISO7405. Dentistry–preclinical evaluation of biocompatibility of medical devices used in dentistry. Test methods for dental materials. Geneva, Switzerland: International Organization for Standardization; 1997

[40] ANSI/ADA Specification No. 41. Recommended standard practices for biological evaluation of dental materials. 2005

[41] Simon CG, Antonucci JM, Liu DW, Skrtic D. *In vitro* cytotoxicity of amorphous calcium phosphate composites. Journal of Bioactive and Compatible Polymers. 2005;**20**(3):279- 295. DOI: 10.1177/0883911505061854

[42] Chatterjee K, Lin-Gibson S, Wallace WE, Parekh SH, Lee YJ, Cicerone MT, et al. The effect of 3D hydrogel scaffold modulus on osteoblast differentiation and mineralization revealed by combinatorial screening. Biomaterials. 2010;**39**(19):5051-5062. DOI: 10.1016/j. biomaterials.2010.03.024

[43] Dalby MJ, Biggs MJP, Gadegaard N, Kalna G, Wilkinson CDW, Oreffo ROC. The control of human mesenchymal cell differentiation using nanoscale symmetry and disorder. Nature Materials. 2007;**6**(12):997-1003. DOI: 10.1038/nmat2013

[44] Skrtic D, Hailer AW, Takagi S, Antonucci JM, Eanes ED. Quantitative assessment of the efficacy of amorphous

**163**

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites*

viability of *Streptococcus mutans* cells in response to antimicrobial treatments. Journal of Microbiological Methods. 2005;**61**(2):161-170. DOI: 10.1016/j.

[52] Esteban Florez FL, Hiers RD, Smart K, Kreth J, Qi F, Merritt J, et al. Realtime assessment of *Streptococcus mutans* biofilm metabolism on resin composite. Dental Materials. 2016;**32**(10):1263- 1269. DOI: 10.1016/j.dental.2016.07.010

[53] Cheng L, Weir MD, Xu HH, Antonucci JM, Kraigsley AM, Lin NJ, et al. Antibacterial amorphous calcium phosphate nanocomposites with a quaternary ammonium dimethacrylate and silver nanoparticles. Dental Materials. 2012;**28**(5):561-572. DOI:

10.1016/j.dental.2012.01.005

ammonium to inhibit biofilms. International Journal of Oral Science. 2016;**8**(3):172-181. DOI: 10.1038/

[55] Li F, Weir MD, Chen J, Xu HH. Comparison of quaternary ammonium-containing with nano-silver-containing adhesive in antibacterial properties and cytotoxicity. Dental Materials. 2013;**29**(4):450-461. DOI: 10/1016/j.

[56] Melo MA, Guedes SF, Xu HH, Rodrigues LK. Nanotechnology-based restorative materials for dental caries management. Trends in Biotechnology. 2013;**31**(8):459-467. DOI: 10.1016/j.

[57] Imazato S. Bio-active restorative materials with antibacterial effects; new dimension of innovation in restorative dentistry. Dental Materials Journal. 2009;**28**(1):11-19. DOI: 10.4012/

ijos.2016.13

dental.2013.01.012

tibitech.2013.05.010

dmj.28.11

[54] Cheng L, Zhang K, Zhou CC, Weir MD, Zhou XD, Xu HH. One-year waterageing of calcium phosphate composite containing nano-silver and quaternary

mimet.2004.11.012

*DOI: http://dx.doi.org/10.5772/intechopen.86640*

[45] Langhorst SE, O'Donnell JNR, Skrtic D. *In vitro* remineralization effectiveness

Quantitative micro-radiographic study. Dental Materials. 2009;**25**:884-891. DOI:

[46] Imazato S, Tarumi H, Kato S, Ebisu S. Water sorption and color stability of composites containing the antibacterial monomer MDPB. Journal of Dentistry.

[47] Antonucci JM. Polymerizable biomedical composition. US patent

[48] Antonucci JM, Zeiger DN, Tang K, Lin-Gibson S, Fowler BO, Lin NJ. Synthesis and characterization of dimethacrylates containing quaternary ammonium functionalities for dental applications. Dental Materials. 2012;**28**(2):219-228. DOI: 10.1016/j.

of polymeric ACP composites:

10.1016/j.dental.2009.01.094

1999;**27**(4):279-283

8217081; July 2012

dental.2011.10.004

[49] Bienek DR, Frukhtbeyn SA, Giuseppetti AA, Okeke UC, Pires RM,

[50] Bienek DR, Frukhtbeyn SA, Giuseppetti AA, Okeke UC, Skrtic D. Antimicrobial monomers for polymeric dental restoratives: Cytotoxicity and physicochemical properties. Journal of Functional Biomaterials. 2018;**9**(1):20. DOI:

dimethacrylates for antimicrobial and remineralizing dental composites. Annals of Dentistry and Oral Disorders.

[51] Merritt J, Kreth J, Qi F, Sullivan R, Shi W. Non-disruptive, real-time analyses of the metabolic status and

Antonucci JM, et al. Ionic

2018:108. PMID:30854515

10.3390/jfb9010020

calcium phosphate/methacrylate composites in remineralizing carieslike lesions artificially produced in bovine enamel. Journal of Dental Research. 1996;**75**(9):1679-1686. DOI: 10.1177/00220345596075091001

*Amorphous Calcium Phosphate as Bioactive Filler in Polymeric Dental Composites DOI: http://dx.doi.org/10.5772/intechopen.86640*

calcium phosphate/methacrylate composites in remineralizing carieslike lesions artificially produced in bovine enamel. Journal of Dental Research. 1996;**75**(9):1679-1686. DOI: 10.1177/00220345596075091001

*Contemporary Topics about Phosphorus in Biology and Materials*

[37] Floyd CJE, Dickens SH. Network structure of bis-GMA- and UDMAbased resin systems. Dental Materials. 2006;**22**:1143-1149. DOI: 10.1016/j.

[38] International Standard ISO10993- 5. Biological evaluation of medical devices- Part 5. Tests for cytotoxicity; In vitro methods. Geneva, Switzerland:

[39] International Standard ISO7405. Dentistry–preclinical evaluation of biocompatibility of medical devices used in dentistry. Test methods for dental materials. Geneva, Switzerland:

[40] ANSI/ADA Specification No. 41. Recommended standard practices for biological evaluation of dental

[41] Simon CG, Antonucci JM, Liu DW, Skrtic D. *In vitro* cytotoxicity of amorphous calcium phosphate composites. Journal of Bioactive and Compatible Polymers. 2005;**20**(3):279- 295. DOI: 10.1177/0883911505061854

[42] Chatterjee K, Lin-Gibson S, Wallace WE, Parekh SH, Lee YJ, Cicerone MT, et al. The effect of 3D hydrogel scaffold modulus on osteoblast differentiation and mineralization revealed by

combinatorial screening. Biomaterials. 2010;**39**(19):5051-5062. DOI: 10.1016/j.

[43] Dalby MJ, Biggs MJP, Gadegaard N, Kalna G, Wilkinson CDW, Oreffo ROC. The control of human

mesenchymal cell differentiation using nanoscale symmetry and disorder. Nature Materials. 2007;**6**(12):997-1003.

[44] Skrtic D, Hailer AW, Takagi S, Antonucci JM, Eanes ED. Quantitative assessment of the efficacy of amorphous

biomaterials.2010.03.024

DOI: 10.1038/nmat2013

International Organization for

International Organization for

Standardization; 1992

Standardization; 1997

materials. 2005

dental.2005.10.009

and hygroscopic expansion. Journal of Dental Research. 2016;**95**(5):543-549. DOI: 10.1177/0022034516633450

[30] Braga LL, Ferracane JL. Contraction stress related to degree of conversion and reaction kinetics. Journal of Dental Research. 2002;**81**(2):114-118. DOI:

[31] Ferracane JL. Developing a more complex understanding of stresses produced in dental composites during polymerization. Dental Materials. 2005;**21**(1):36-42. DOI: 10.1016/j.

10.1177/0810114

dental.2004.10.004

dental.2005.02.004

[32] Kleverlaan CJ, Feizler AJ. Polymerization shrinkage and contraction stress of dental resin composites. Dental Materials. 2005;**21**(1):1150-1157. DOI: 10.1016/j.

[33] Choi KK, Condon JR, Ferracane JL. The effects of adhesive thickness on polymerization contraction stress of composite. Journal of Dental Research. 2000;**79**(3):812-817. DOI: 10.1177/00220345000790030501

[34] Antonucci JM, Giuseppetti AA, O'Donnell JNR, Schumacher GE, Skrtic D. Polymerization stress development in dental composites: Effect of cavity design factor. Materials. 2009;**2**(1):169-

[35] O' Donnell JNR, Langhorst SE, Fow MD, Skrtic D. Light-cured dimethacrylate-based resins and their composites; Comparative study of mechanical strength, water sorption and ion release. Journal of Bioactive and Compatible Polymers. 2008;**23**:207-226. DOI: 10.1177/0883911508089932

[36] Durner J, Walther UI, Zaspel J, Hickel R, Reich FX. Metabolism of TEGDMA and HEMA in human cells. Biomaterials. 2010;**31**:818-823. DOI: 10.1016/j.biomaterials.2009.09.097

180. DOI: 10.3390/ma2010169

**162**

[45] Langhorst SE, O'Donnell JNR, Skrtic D. *In vitro* remineralization effectiveness of polymeric ACP composites: Quantitative micro-radiographic study. Dental Materials. 2009;**25**:884-891. DOI: 10.1016/j.dental.2009.01.094

[46] Imazato S, Tarumi H, Kato S, Ebisu S. Water sorption and color stability of composites containing the antibacterial monomer MDPB. Journal of Dentistry. 1999;**27**(4):279-283

[47] Antonucci JM. Polymerizable biomedical composition. US patent 8217081; July 2012

[48] Antonucci JM, Zeiger DN, Tang K, Lin-Gibson S, Fowler BO, Lin NJ. Synthesis and characterization of dimethacrylates containing quaternary ammonium functionalities for dental applications. Dental Materials. 2012;**28**(2):219-228. DOI: 10.1016/j. dental.2011.10.004

[49] Bienek DR, Frukhtbeyn SA, Giuseppetti AA, Okeke UC, Pires RM, Antonucci JM, et al. Ionic dimethacrylates for antimicrobial and remineralizing dental composites. Annals of Dentistry and Oral Disorders. 2018:108. PMID:30854515

[50] Bienek DR, Frukhtbeyn SA, Giuseppetti AA, Okeke UC, Skrtic D. Antimicrobial monomers for polymeric dental restoratives: Cytotoxicity and physicochemical properties. Journal of Functional Biomaterials. 2018;**9**(1):20. DOI: 10.3390/jfb9010020

[51] Merritt J, Kreth J, Qi F, Sullivan R, Shi W. Non-disruptive, real-time analyses of the metabolic status and

viability of *Streptococcus mutans* cells in response to antimicrobial treatments. Journal of Microbiological Methods. 2005;**61**(2):161-170. DOI: 10.1016/j. mimet.2004.11.012

[52] Esteban Florez FL, Hiers RD, Smart K, Kreth J, Qi F, Merritt J, et al. Realtime assessment of *Streptococcus mutans* biofilm metabolism on resin composite. Dental Materials. 2016;**32**(10):1263- 1269. DOI: 10.1016/j.dental.2016.07.010

[53] Cheng L, Weir MD, Xu HH, Antonucci JM, Kraigsley AM, Lin NJ, et al. Antibacterial amorphous calcium phosphate nanocomposites with a quaternary ammonium dimethacrylate and silver nanoparticles. Dental Materials. 2012;**28**(5):561-572. DOI: 10.1016/j.dental.2012.01.005

[54] Cheng L, Zhang K, Zhou CC, Weir MD, Zhou XD, Xu HH. One-year waterageing of calcium phosphate composite containing nano-silver and quaternary ammonium to inhibit biofilms. International Journal of Oral Science. 2016;**8**(3):172-181. DOI: 10.1038/ ijos.2016.13

[55] Li F, Weir MD, Chen J, Xu HH. Comparison of quaternary ammonium-containing with nano-silver-containing adhesive in antibacterial properties and cytotoxicity. Dental Materials. 2013;**29**(4):450-461. DOI: 10/1016/j. dental.2013.01.012

[56] Melo MA, Guedes SF, Xu HH, Rodrigues LK. Nanotechnology-based restorative materials for dental caries management. Trends in Biotechnology. 2013;**31**(8):459-467. DOI: 10.1016/j. tibitech.2013.05.010

[57] Imazato S. Bio-active restorative materials with antibacterial effects; new dimension of innovation in restorative dentistry. Dental Materials Journal. 2009;**28**(1):11-19. DOI: 10.4012/ dmj.28.11

**Chapter 9**

**Abstract**

Structural and Calorimetric

Studies of Zinc, Magnesium

revealed the depolymerization of metaphosphate chains (Q2

enthalpy increases with the incorporation of ZnO oxide.

advantages over conventional silicate and borate glasses.

is endothermic for lower MO content and become exothermic when x rises. For (50-x/2)Na2O-xZnO-(50-x/2)P2O5 (0 ≤ x ≤ 33 mol%) glasses, the formation

**Keywords:** phosphate glasses, phosphate-silicate glasses, depolymerization,

modifier oxide, calorimetric dissolution, formation enthalpy, thermochemical cycle

Over the past several decades, great interests have been considered for phosphate glasses due to their superior physical properties which impart them a several

Phosphate-based glasses are an interesting class in the world of glass and glass ceramics owing to their higher thermal expansion, lower melting and softening

tion of phosphate dimers (Q1

**1. Introduction**

**165**

*and Ahmed Hichem Hamzaoui*

and Manganese Based Phosphate

and Phosphate-Silicate Glasses

*Refka Oueslati Omrani, Mohamed Jemal, Ismail Khattech*

Glasses of the (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mg or Mn) (0 ≤ x ≤ 33 mol%), (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) (0 ≤ x ≤ 33 mol%), and (0.9-x) NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) were prepared by the melt quenching technique. Samples were investigated by means of X-ray diffraction, Archimede's method, ellipsometry, Fourier-transformed infrared (FTIR), Raman, 31P solid state magic angle spinning nuclear magnetic resonance (MAS-NMR), UV-visible spectroscopy and calorimetry. For zinc, manganese and magnesium phosphate glasses, the increase in density with the addition of MO oxide suggests the compactness of the vitreous network. For zinc phosphate silicate glasses, the variations of density and refractive index were attributed to the structural changes when SiO2 oxide is progressively introduced. The increase in the glass transition temperature (Tg) reflects an increase in the cross-link strength of the structure as MO and SiO2 oxides are gradually incorporated. For all glass composition, spectroscopic investigations

) allowing the forma-

). Calorimetric dissolution shows that the dissolution

[58] Wang JJ, Liu F. Photoinduced graft polymerization of 2-methacryloyloxyethyl phosphorylcholine on silicone hydrogels for reducing protein adsorption. Journal of Materials Science. Materials in Medicine. 2011;**22**(12):2651-2657. DOI: 10.1007/s10856-011-4452-y

[59] Zhang N, Zhang K, Xie X, Dai Z, Zhao Z, Imazato S, et al. Nanostructured polymeric materials with protein-repellent and anti-caries properties for dental applications. Nanomaterials. 2018;**8**:393. DOI: 10.3390/nano8060393

#### **Chapter 9**

*Contemporary Topics about Phosphorus in Biology and Materials*

[58] Wang JJ, Liu F. Photoinduced

phosphorylcholine on silicone hydrogels for reducing protein adsorption. Journal of Materials Science. Materials in Medicine. 2011;**22**(12):2651-2657. DOI:

graft polymerization of 2-methacryloyloxyethyl

10.1007/s10856-011-4452-y

10.3390/nano8060393

[59] Zhang N, Zhang K, Xie X, Dai Z, Zhao Z, Imazato S, et al. Nanostructured polymeric materials with protein-repellent and anti-caries properties for dental applications. Nanomaterials. 2018;**8**:393. DOI:

**164**

## Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate and Phosphate-Silicate Glasses

*Refka Oueslati Omrani, Mohamed Jemal, Ismail Khattech and Ahmed Hichem Hamzaoui*

#### **Abstract**

Glasses of the (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mg or Mn) (0 ≤ x ≤ 33 mol%), (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) (0 ≤ x ≤ 33 mol%), and (0.9-x) NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) were prepared by the melt quenching technique. Samples were investigated by means of X-ray diffraction, Archimede's method, ellipsometry, Fourier-transformed infrared (FTIR), Raman, 31P solid state magic angle spinning nuclear magnetic resonance (MAS-NMR), UV-visible spectroscopy and calorimetry. For zinc, manganese and magnesium phosphate glasses, the increase in density with the addition of MO oxide suggests the compactness of the vitreous network. For zinc phosphate silicate glasses, the variations of density and refractive index were attributed to the structural changes when SiO2 oxide is progressively introduced. The increase in the glass transition temperature (Tg) reflects an increase in the cross-link strength of the structure as MO and SiO2 oxides are gradually incorporated. For all glass composition, spectroscopic investigations revealed the depolymerization of metaphosphate chains (Q2 ) allowing the formation of phosphate dimers (Q1 ). Calorimetric dissolution shows that the dissolution is endothermic for lower MO content and become exothermic when x rises. For (50-x/2)Na2O-xZnO-(50-x/2)P2O5 (0 ≤ x ≤ 33 mol%) glasses, the formation enthalpy increases with the incorporation of ZnO oxide.

**Keywords:** phosphate glasses, phosphate-silicate glasses, depolymerization, modifier oxide, calorimetric dissolution, formation enthalpy, thermochemical cycle

#### **1. Introduction**

Over the past several decades, great interests have been considered for phosphate glasses due to their superior physical properties which impart them a several advantages over conventional silicate and borate glasses.

Phosphate-based glasses are an interesting class in the world of glass and glass ceramics owing to their higher thermal expansion, lower melting and softening

temperature, higher ultra-violet transmission and optical characteristic. Phosphate glasses have potential of applications in medicine, biology, batteries, laser technology, electronic, telecommunication [1–44].

are intermediate structures [11, 13, 16, 19, 20]. **Figure 1a** shows the nomenclature of phosphate groups as Qn tetrahedral sites also with the variation of an O/P ratio [45]. For silicon-oxygen networks, n varies between 0 and 4, where Q0 represents

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

Glasses results from many possible combinations of network-forming oxides together with one or several modifier or intermediate oxides which lead to a special

*a. Infrared spectra of (50-x)Na2O-xZnO-50P2O5 glasses: (a) 0 mol% ZnO, (b) 5 mol% ZnO, (c) 10 mol% ZnO, (d) 15 mol% ZnO, (e) 20 mol% ZnO, (f) 25 mol% ZnO, (g) 30 mol% ZnO, (h) 33 mol% ZnO. b. Infrared spectra of (0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) glasses: (a) 0 mol SiO2, (b) 0.02 mol*

*SiO2, (c) 0.04 mol SiO2, (d) 0.06 mol SiO2, (e) 0.08 mol SiO2, (f) 0.10 mol SiO2.*

, Q2 and Q<sup>1</sup> represents the interme-

<sup>4</sup>), Q<sup>4</sup> is pure SiO2 and Q<sup>3</sup>

orthosilicates (SiO4

**Figure 1.**

**167**

physical properties [22].

diate silicate structures [2, 20].

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

In recent times, phosphate glasses have received considerable interest as a result of the synthesis of new glass composition with high chemical stability. The improvement of this quality induces the application of phosphate glasses in numerous fields of materials science, such as fast ionic conductors, semiconductors, photonic materials, hermetic seals, rare-earth in host solid state lasers and biomedical materials [1–44].

Thus, phosphate and silicate glasses are the most important materials which can extensively be used for laser sources and fiber amplifiers [3].

The basic structure unit of the phosphate network is based on corner-sharing PO4 units which form chains and rings or isolated groups PO4 [6].

The covalence of the bridging oxygen atoms is responsible for the formation of phosphate groups [18, 27].

The structure of the three dimensional network is obtained by linking three oxygen atoms with others PO4 tetrahedrons. The metaphosphate groups contain two covalent bridging oxygen atoms. Whereas, the pyrophosphate groups are formed by bending only single oxygen atom with other tetrahedral site.

Recently, the structure of phosphate glass is describes using the O/P ratio [13, 18, 19, 27]. Furthermore, many investigators used the O/P ratio in order to classify the distribution of phosphate groups in the vitreous network. **Table 1** reports the classification of phosphate glasses as a function of O/P ratio [45].

For 2.5 < O/P < 3, the glass network is described by the distribution of ultraphosphate groups [19].

The metaphosphate groups are obtained for O/P ratio equal to 3. The glass network is described by the connection of PO4 tetrahedral anions with neighbors in order to form chains and rings.

For polyphosphate glasses, the O/P ratio is between 3 and 3.5. The structure is described by chains formed by PO4 tetrahedral anions joined with others.

For O/P = 3.5, the structure is obtained by forming the phosphate dimers such as pyrophosphate groups in which two PO4 tetrahedral shared one bridging oxygen [19]. For 3.5 < O/P < 4, the isolated PO4 <sup>3</sup> units are formed such as orthophosphate.

The increase of O/P ratio induces the depolymerization of phosphate groups which suggests the shortening of the average chain length [19].

The network connectivity of phosphate compound is conventionally expressed as Qn tetrahedral sites (n = 0…3), when n is the number of bridging oxygen (BO) per Q unit to neighbor phosphate tetrahedral [2, 23, 28]. Q0 represents orthophosphates (PO4 <sup>3</sup>), Q3 is pure P2O5 and Q2 (metaphosphates) and Q1 (pyrophosphates)


#### **Table 1.**

*Classification of phosphate glasses as a function of O/P ratio and Qn tetrahedral sites.*

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

are intermediate structures [11, 13, 16, 19, 20]. **Figure 1a** shows the nomenclature of phosphate groups as Qn tetrahedral sites also with the variation of an O/P ratio [45].

For silicon-oxygen networks, n varies between 0 and 4, where Q0 represents orthosilicates (SiO4 <sup>4</sup>), Q<sup>4</sup> is pure SiO2 and Q<sup>3</sup> , Q2 and Q<sup>1</sup> represents the intermediate silicate structures [2, 20].

Glasses results from many possible combinations of network-forming oxides together with one or several modifier or intermediate oxides which lead to a special physical properties [22].

#### **Figure 1.**

temperature, higher ultra-violet transmission and optical characteristic. Phosphate glasses have potential of applications in medicine, biology, batteries, laser technol-

of the synthesis of new glass composition with high chemical stability. The

extensively be used for laser sources and fiber amplifiers [3].

PO4 units which form chains and rings or isolated groups PO4 [6].

classification of phosphate glasses as a function of O/P ratio [45].

which suggests the shortening of the average chain length [19].

**O/P Classification Q<sup>n</sup>** 2.5-3 Ultraphosphates Q<sup>3</sup> 3 Metaphosphates Q<sup>2</sup> >3 Polyphosphates Q<sup>2</sup> + Q<sup>1</sup> 3.5 Pyrophosphates Q<sup>2</sup> >3.5 Orthophosphates Q<sup>0</sup>

*Classification of phosphate glasses as a function of O/P ratio and Qn tetrahedral sites.*

In recent times, phosphate glasses have received considerable interest as a result

Thus, phosphate and silicate glasses are the most important materials which can

The basic structure unit of the phosphate network is based on corner-sharing

The covalence of the bridging oxygen atoms is responsible for the formation of

Recently, the structure of phosphate glass is describes using the O/P ratio [13, 18, 19, 27]. Furthermore, many investigators used the O/P ratio in order to classify the distribution of phosphate groups in the vitreous network. **Table 1** reports the

The structure of the three dimensional network is obtained by linking three oxygen atoms with others PO4 tetrahedrons. The metaphosphate groups contain two covalent bridging oxygen atoms. Whereas, the pyrophosphate groups are formed by bending only single oxygen atom with other tetrahedral site.

For 2.5 < O/P < 3, the glass network is described by the distribution of

described by chains formed by PO4 tetrahedral anions joined with others.

The metaphosphate groups are obtained for O/P ratio equal to 3. The glass network is described by the connection of PO4 tetrahedral anions with neighbors in

For polyphosphate glasses, the O/P ratio is between 3 and 3.5. The structure is

For O/P = 3.5, the structure is obtained by forming the phosphate dimers such as pyrophosphate groups in which two PO4 tetrahedral shared one bridging oxygen

The increase of O/P ratio induces the depolymerization of phosphate groups

**O**: oxygen atom **P**: Phosphorus atom **Q**: Phosphorus tetrahedral sites

The network connectivity of phosphate compound is conventionally expressed as Qn tetrahedral sites (n = 0…3), when n is the number of bridging oxygen (BO) per Q unit to neighbor phosphate tetrahedral [2, 23, 28]. Q0 represents orthophos-

<sup>3</sup>), Q3 is pure P2O5 and Q2 (metaphosphates) and Q1 (pyrophosphates)

<sup>3</sup> units are formed such as orthophosphate.

improvement of this quality induces the application of phosphate glasses in numerous fields of materials science, such as fast ionic conductors, semiconductors, photonic materials, hermetic seals, rare-earth in host solid state lasers and biomedical

ogy, electronic, telecommunication [1–44].

*Contemporary Topics about Phosphorus in Biology and Materials*

materials [1–44].

phosphate groups [18, 27].

ultraphosphate groups [19].

order to form chains and rings.

phates (PO4

**Table 1.**

**166**

[19]. For 3.5 < O/P < 4, the isolated PO4

*a. Infrared spectra of (50-x)Na2O-xZnO-50P2O5 glasses: (a) 0 mol% ZnO, (b) 5 mol% ZnO, (c) 10 mol% ZnO, (d) 15 mol% ZnO, (e) 20 mol% ZnO, (f) 25 mol% ZnO, (g) 30 mol% ZnO, (h) 33 mol% ZnO. b. Infrared spectra of (0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) glasses: (a) 0 mol SiO2, (b) 0.02 mol SiO2, (c) 0.04 mol SiO2, (d) 0.06 mol SiO2, (e) 0.08 mol SiO2, (f) 0.10 mol SiO2.*

Introducing alkali metal oxide or divalent metal oxide to the glass network induces the fundamental optical absorption edge falls in the UV region bellow 400 nm which meets with the requirement for desirable applications in optical systems [22]. These additions not only enhance the chemical durability of the phosphate glasses but also can impart special functions to the glasses and expand the glass application fields.

Furthermore, Szumera et al. have detailed the effect of MoO3 addition on silicate phosphate glasses using spectroscopic analysis such as FTIR, Raman and 31P MAS-NMR. The molar content of SiO2 decreases from 41.6 to 39.6 mol% and the P2O5 proportion increases from 5.7 to 7.8 mol% [44]. The obtained results revealed the cleavage of oxygen bridges which suggests that acts as a network modifier [44]. In our knowledge, the thermochemical data of zinc-, magnesium-, and manganese-based phosphate and phosphate-silicate glasses are rare in literature. For this purpose, the aims of this research were to study the correlations between structural changes, thermal investigations, optical properties and thermochemical behaviors of these glassy compounds with the incorporation of zinc, magnesium

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

However, it is well known that phosphate glasses have a poor chemical durability, volatile nature and hygroscopic character. These disadvantages decrease their

The addition of alkali and alkaline earth cations with the decease of phosphorus

The addition of transition metal oxides (CuO, MgO, ZnO, MnO, CaO…) into the vitreous network (TMO) disrupted the P▬O▬P bridges leading to the structure depolymerization and the formation of non-bridging oxygen atoms (NBO) [2, 5, 6, 8, 9, 13, 21, 25, 32, 35]. Which induces the formation of P▬O▬M bonds replacing the easily hydrosoluble P▬O▬P bridge that improve the chemical durability of the

Among several oxides mentioned above, zinc oxide gained considerable attention because Zinc doped glasses find numerous applications in optic field can be used as LED light sources and substrates for optical waveguides. It can also play an

Zinc phosphate compositions are chemically durable, have processing tempera-

In recent years, there have also appeared some publications on the influence of the addition of ZnO on the structure of glasses with a mixed phosphate-silicate

Among the wide class of phosphate glasses, ZnO-based glasses have low glass transition temperature in the region of 280–380°C and significantly high chemical

Furthermore, the addition of ZnO to the glass network is expected to improve the chemical stability of the structure. It can also ameliorate the electrical, optical and magnetic properties of glasses due to the appearance of P▬O▬Zn ionic bond which induces the increase in the compactness and the rigidity of the glass network

In fact, in crystalline solid compound, the structure was described by a repetitive

arrangement of a large scale patterns contrary to amorphous structure which

Additionally, ZnO is an intermediate oxide. It can act as former or modifier network depending on his content in phosphate network. When it occupies the tetrahedral sites by forming ZnO4 structural units, ZnO oxide plays the role of glass former. But, when it occupies the octahedral sites coordinated, ZnO oxide acts as

tures under 400°C and can be co-formed with high temperature under 400°C

, K<sup>+</sup>

, Mg<sup>+</sup>

, Ca2+…)

content can depolymerize the glass network which suggests the cleavage of

disrupts the glassy network, leading to the structure depolymerization and the formation of non-bridging oxygen atoms (NBO), also named terminal

stability which limits their use in technological applications [5].

The incorporation of certain network modifier cations (Na<sup>+</sup>

important role in bone formation and mineralization [3].

polymers to produce unusual organic/inorganic composites [19].

and manganese oxides.

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

P▬O▬P bridges [28].

oxygen (OT) [2, 4].

phosphate network [21].

structure [7, 34].

durability [14].

[5, 13, 32, 35].

exhibits a short range order.

glass modifier [7, 25].

**169**

Furthermore, alkali phosphate glasses have attracted more attention due to their mixed electronic ionic conductivity, low melting point and strong glass-forming character [4]. Among the phosphate-based glasses those containing calcium, magnesium, sodium and zinc have received great attention due to their excellent optical properties, high refractive index, low dispersion and good transparency in the UV and IR region [1]. With the decrease of P2O5 content, the glass become more resistant to moisture attack but restricts the glass formation areas. Thus, MgO oxide was incorporated in order to overcome these problems [23].

Nevertheless, phosphate-based glasses containing transition metal ions are scientifically interesting materials due to their attractive properties which can be used in many technological applications including electronic and electro-optical devices [21].

In fact, transition metal oxides can be dissolved easily in phosphate glasses which exhibit one than more oxidation sates [21, 25, 30].

For glasses doped with manganese ions. These latter are presented in the +2 or +3 oxidation states. The content of Mn3+ ions in the glass leads to the staining glass in the color range from light to dark purple depending on the concentration [30]. This characteristic coloration could be explained by the d-d electronic transitions. This color can be associated with the broad absorption band in the visible region at 520 nm, which pertains to the Mn3+ ions [25, 30]. This behavior allows obtaining additional luminescence bands in the red spectral region that shift LED emission from cool white to warm light [30].

Contrary to Zn2+ and Mg2+ ions presented in one oxidation state, spectralluminescent properties of manganese ions in phosphate glasses allow them to be good candidates for interesting optical applications [21, 25, 30].

In borate glasses, manganese exists mainly as Mn3+ ions with octahedral coordination in glass networks whereas in silicate and germinate glasses, it identified as Mn2+ ions with both tetrahedral and octahedral coordination [21].

Referring to literature, Montagne et al. have been studied the zinc phosphate glasses with a general formula (100-x)NaPO3-xZnO with 0 < x < 33.3 mol% using 31P MAS-NMR, 31P NMR of liquid sample, visible spectroscopy, refractive index measurements, density evolutions, Tg variations, activation energy, chemical durability and chemical analysis [1, 2, 46, 47].

The obtained results revealed the distortion of metaphosphate chains (Q2 ) which suggests the formation of phosphate dimmers (Q<sup>1</sup> ) [46, 47].

Moreover, Zotov et al. have been studied the manganese phosphate glasses with a general formula (MnO)x(NaPO3)1-x when x = 0.0, 0.024, 0.048 and 0.167 mol [48].

These investigations have been performed using X ray diffraction, EXAFS and Raman spectroscopy. The increase of MnO content causes the depolymerization of metaphosphate chains leading to the decrease of the average chain length [48].

For zinc phosphate-silicate glasses, chemical compositions of the prepared glasses, picked from literature, have been chosen with higher level of SiO2 and lower P2O5 content [2, 20, 24, 34].

Aguiar et al. have been studied the Na2O-MgO-CaO-P2O5-SiO2 bioactive glasses using Raman, 31P MAS-NMR and 29Si MAS-NMR spectroscopies. The glass compositions were prepared with varying SiO2 content from 25 to 54 mol% and the P2O5 proportion from 2 to 11 mol% [2, 20, 24].

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

Furthermore, Szumera et al. have detailed the effect of MoO3 addition on silicate phosphate glasses using spectroscopic analysis such as FTIR, Raman and 31P MAS-NMR. The molar content of SiO2 decreases from 41.6 to 39.6 mol% and the P2O5 proportion increases from 5.7 to 7.8 mol% [44]. The obtained results revealed the cleavage of oxygen bridges which suggests that acts as a network modifier [44].

In our knowledge, the thermochemical data of zinc-, magnesium-, and manganese-based phosphate and phosphate-silicate glasses are rare in literature. For this purpose, the aims of this research were to study the correlations between structural changes, thermal investigations, optical properties and thermochemical behaviors of these glassy compounds with the incorporation of zinc, magnesium and manganese oxides.

However, it is well known that phosphate glasses have a poor chemical durability, volatile nature and hygroscopic character. These disadvantages decrease their stability which limits their use in technological applications [5].

The addition of alkali and alkaline earth cations with the decease of phosphorus content can depolymerize the glass network which suggests the cleavage of P▬O▬P bridges [28].

The incorporation of certain network modifier cations (Na<sup>+</sup> , K<sup>+</sup> , Mg<sup>+</sup> , Ca2+…) disrupts the glassy network, leading to the structure depolymerization and the formation of non-bridging oxygen atoms (NBO), also named terminal oxygen (OT) [2, 4].

The addition of transition metal oxides (CuO, MgO, ZnO, MnO, CaO…) into the vitreous network (TMO) disrupted the P▬O▬P bridges leading to the structure depolymerization and the formation of non-bridging oxygen atoms (NBO) [2, 5, 6, 8, 9, 13, 21, 25, 32, 35]. Which induces the formation of P▬O▬M bonds replacing the easily hydrosoluble P▬O▬P bridge that improve the chemical durability of the phosphate network [21].

Among several oxides mentioned above, zinc oxide gained considerable attention because Zinc doped glasses find numerous applications in optic field can be used as LED light sources and substrates for optical waveguides. It can also play an important role in bone formation and mineralization [3].

Zinc phosphate compositions are chemically durable, have processing temperatures under 400°C and can be co-formed with high temperature under 400°C polymers to produce unusual organic/inorganic composites [19].

In recent years, there have also appeared some publications on the influence of the addition of ZnO on the structure of glasses with a mixed phosphate-silicate structure [7, 34].

Among the wide class of phosphate glasses, ZnO-based glasses have low glass transition temperature in the region of 280–380°C and significantly high chemical durability [14].

Furthermore, the addition of ZnO to the glass network is expected to improve the chemical stability of the structure. It can also ameliorate the electrical, optical and magnetic properties of glasses due to the appearance of P▬O▬Zn ionic bond which induces the increase in the compactness and the rigidity of the glass network [5, 13, 32, 35].

In fact, in crystalline solid compound, the structure was described by a repetitive arrangement of a large scale patterns contrary to amorphous structure which exhibits a short range order.

Additionally, ZnO is an intermediate oxide. It can act as former or modifier network depending on his content in phosphate network. When it occupies the tetrahedral sites by forming ZnO4 structural units, ZnO oxide plays the role of glass former. But, when it occupies the octahedral sites coordinated, ZnO oxide acts as glass modifier [7, 25].

Introducing alkali metal oxide or divalent metal oxide to the glass network induces the fundamental optical absorption edge falls in the UV region bellow 400 nm which meets with the requirement for desirable applications in optical systems [22]. These additions not only enhance the chemical durability of the phosphate glasses but also can impart special functions to the glasses and expand

Furthermore, alkali phosphate glasses have attracted more attention due to their mixed electronic ionic conductivity, low melting point and strong glass-forming character [4]. Among the phosphate-based glasses those containing calcium, magnesium, sodium and zinc have received great attention due to their excellent optical properties, high refractive index, low dispersion and good transparency in the UV and IR region [1]. With the decrease of P2O5 content, the glass become more resistant to moisture attack but restricts the glass formation areas. Thus, MgO oxide

Nevertheless, phosphate-based glasses containing transition metal ions are scientifically interesting materials due to their attractive properties which can be used in many technological applications including electronic and electro-optical

In fact, transition metal oxides can be dissolved easily in phosphate glasses

Contrary to Zn2+ and Mg2+ ions presented in one oxidation state, spectralluminescent properties of manganese ions in phosphate glasses allow them to be

In borate glasses, manganese exists mainly as Mn3+ ions with octahedral coordination in glass networks whereas in silicate and germinate glasses, it identified as

Referring to literature, Montagne et al. have been studied the zinc phosphate glasses with a general formula (100-x)NaPO3-xZnO with 0 < x < 33.3 mol% using 31P MAS-NMR, 31P NMR of liquid sample, visible spectroscopy, refractive index measurements, density evolutions, Tg variations, activation energy, chemical dura-

The obtained results revealed the distortion of metaphosphate chains (Q2

Moreover, Zotov et al. have been studied the manganese phosphate glasses with a general formula (MnO)x(NaPO3)1-x when x = 0.0, 0.024, 0.048 and 0.167 mol [48]. These investigations have been performed using X ray diffraction, EXAFS and Raman spectroscopy. The increase of MnO content causes the depolymerization of metaphosphate chains leading to the decrease of the average chain length [48]. For zinc phosphate-silicate glasses, chemical compositions of the prepared glasses, picked from literature, have been chosen with higher level of SiO2 and

Aguiar et al. have been studied the Na2O-MgO-CaO-P2O5-SiO2 bioactive glasses using Raman, 31P MAS-NMR and 29Si MAS-NMR spectroscopies. The glass compositions were prepared with varying SiO2 content from 25 to 54 mol% and the P2O5

) [46, 47].

) which

For glasses doped with manganese ions. These latter are presented in the +2 or +3 oxidation states. The content of Mn3+ ions in the glass leads to the staining glass in the color range from light to dark purple depending on the concentration [30]. This characteristic coloration could be explained by the d-d electronic transitions. This color can be associated with the broad absorption band in the visible region at 520 nm, which pertains to the Mn3+ ions [25, 30]. This behavior allows obtaining additional luminescence bands in the red spectral region that shift LED emission

was incorporated in order to overcome these problems [23].

*Contemporary Topics about Phosphorus in Biology and Materials*

which exhibit one than more oxidation sates [21, 25, 30].

good candidates for interesting optical applications [21, 25, 30].

Mn2+ ions with both tetrahedral and octahedral coordination [21].

from cool white to warm light [30].

bility and chemical analysis [1, 2, 46, 47].

lower P2O5 content [2, 20, 24, 34].

**168**

proportion from 2 to 11 mol% [2, 20, 24].

suggests the formation of phosphate dimmers (Q<sup>1</sup>

the glass application fields.

devices [21].

The literature concerning zinc in various mixed oxide compounds revealed that with the exception of few structures, zinc has tetrahedral oxygen ligancy and the zinc-oxygen distance varies only slightly [25, 34]. It was concluded that zinc had a coordination number of six in borate and silicate glasses [34].

(0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) compositions have been prepared

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

A series of glasses were prepared by varying the MO (M = Zn, Mn, Mg) content from 0 to 33 mol% using reagent grade compounds, NaH2PO4, NH4H2PO4, MgO, ZnO, MnCO3 with a high purity (99% purity), in the suitable proportions.

The mixture corresponding to the desired compositions was heated in platinum crucible at 400°C in order to evaporate water and start the condensation of phosphate groups. The temperature was then progressively increased to 750–900°C, depending on glass composition, and held constant for 30 min. The batch was finally quenched to room temperature under air atmosphere in order to produce

Using the same technique, phosphate-based silicate glasses with a general formula (0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) have been synthesized using reagent grade compounds, NaH2PO4.H2O ZnO and SiO2 with a high purity (99%

The mixture was then putted in platinum crucible at 400°C for 1 hour in order to eliminate residual water. The temperature was raised progressively to 1200°C for 30 min in order to homogenize the melting mixture. Finally, the batch was quenched to room temperature under air atmosphere in order to obtain glasses. The amorphous state was confirmed by X-ray diffraction. All the products were annealed at about 20°C below their glass transition temperature for 2 hours in order

Phosphorus, sodium, magnesium, zinc, manganese and silica were analyzed by inductively coupled plasma atomic emission spectroscopy (Jobin Yvon Ultra C).

Density of glass is a strong function of its composition and its intrinsic property which shed light on the short range structure of the glassy material [4, 37]. This work presents a series of glasses with various amounts of modifier oxides. These modifies, depending on their polarity and size, tend to occupy the interstices within the network and form new bonds resulting in a change in the structure and proper-

Glass density measurements have been determined using the standard Archimedes method using diethyl orthophthalate as immersion fluid. The relative error

The molar volume of glasses has been calculated from the density (Vm = M/ρ)

For (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mn, Mg) where 3 ≤ O/P ≤ 3.49 and (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) with O/P = 3 (0 ≤ x ≤ 33 mol%) series glasses, the density increases gradually with the incorporation of MO oxide. The increase in density indicates that the MO oxide reticulate the vitreous network because P▬O▬M bond are more ionic than P▬O▬P [11–13, 15, 29, 32, 33, 42]. **Table 2** shows that the molar volume deceases monotonically with the increase

The decrease in the molar volume for (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mn, Mg) where 3 ≤ O/P ≤ 3.49 and (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) with O/P = 3 (0 ≤ x ≤ 33 mol%) series glasses could be explained by the higher

to eliminate internal tensions and get a more homogenized sample.

using a melt quenching technique.

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

purity) with the desired compositions.

vitreous compounds.

**2.2 ICP analysis**

**2.3 Physical properties**

ties of the glass [4, 37].

and the molar weight.

**171**

of these measurements is 3%.

of ZnO content for phosphate glasses.

*2.3.1 Density and molar volume*

Interestingly, the heat treatment of glasses at temperature higher than their glass transition (Tg) or crystallization temperature (Tc) improves the electrical conductivity of the glassy compound due to the "structural relaxation" of the glass network [36].

ZnO is widely used in glass production because it improves the glass quality by enhancing mechanical properties and chemical durability and by reducing the thermal expansion [7]. Zinc is a microelement that plays an important role in the bone formation and mineralization. Zinc containing glasses and glass ceramics has been developed for bone engineering applications [7]. The small size of Zn ion (0.74 Å) helps it to locate itself into smaller cavities of the network [37].

Moreover, MgO oxide is of interest from a biological viewpoint because Mg2+ is known to play a physiological role in positively influencing bone strength which can be substituted into apatites [13, 18].

The bioactive behavior of magnesium rich glasses is identified as their ability to react chemically with living tissues, forming with them mechanically strong and lasting bonds. These bone bondings are attributed to the formation of an apatite-like layer on the glass surface, with composition and structure equivalent to the mineral phase of bone.

In fact, bioactive glasses have received special attention due to their better bone bonding ability in vivo. Due to their good bioactive and tailorable degradation properties, these glasses can be used for various biomedical applications such as bone graft, filler, dental, implant coating.

Furthermore, by increasing the concentration of modifier oxides, electrical conductivity of the glass increases. This property was probably influenced by the structural changes resulting from the disruption of the glass network which affected the mobility of the cations and anions when the modifying oxide was progressively introduced [38].

Glasses containing transition metal oxides possess interesting electronic, optic and magnetic properties due to the ability to exist in more than one valence state. However, the electronic conduction of these glassy compounds is resulting from the electronic transfer of cation that exists in different valence sates [4, 38].

Compared with phosphate glasses, silicate glasses exhibit superior chemical resistance which makes them compatible with the fabrication process in the development of optical devices [3].

Silicate glasses are an attractive host matrix for transition metal ions due to their excellent optical and mechanical properties, good chemical durability, good chemical stability and low thermal expansion coefficient leading to strong thermal resistance [6]. Silicate glasses have many advantages rather than phosphate glasses. Silicatebased glasses are chemically durable, thermally stable and optically transparent at excitation and lasing wavelength. However, the higher viscosity of these glasses allows the glass to be formed without crystallization process. In addition, these amorphous materials are useful in optics as lenses or beam splitters in optical telecommunications, micro- and optoelectronics and in near IR-windows [6, 39, 40].

#### **2. Experimental procedures**

#### **2.1 Glass preparation**

Glasses of the (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mn, Mg) (0 ≤ x ≤ 33 mol%), (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) (0 ≤ x ≤ 33 mol%), and *Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

(0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) compositions have been prepared using a melt quenching technique.

A series of glasses were prepared by varying the MO (M = Zn, Mn, Mg) content from 0 to 33 mol% using reagent grade compounds, NaH2PO4, NH4H2PO4, MgO, ZnO, MnCO3 with a high purity (99% purity), in the suitable proportions.

The mixture corresponding to the desired compositions was heated in platinum crucible at 400°C in order to evaporate water and start the condensation of phosphate groups. The temperature was then progressively increased to 750–900°C, depending on glass composition, and held constant for 30 min. The batch was finally quenched to room temperature under air atmosphere in order to produce vitreous compounds.

Using the same technique, phosphate-based silicate glasses with a general formula (0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) have been synthesized using reagent grade compounds, NaH2PO4.H2O ZnO and SiO2 with a high purity (99% purity) with the desired compositions.

The mixture was then putted in platinum crucible at 400°C for 1 hour in order to eliminate residual water. The temperature was raised progressively to 1200°C for 30 min in order to homogenize the melting mixture. Finally, the batch was quenched to room temperature under air atmosphere in order to obtain glasses.

The amorphous state was confirmed by X-ray diffraction. All the products were annealed at about 20°C below their glass transition temperature for 2 hours in order to eliminate internal tensions and get a more homogenized sample.

#### **2.2 ICP analysis**

The literature concerning zinc in various mixed oxide compounds revealed that with the exception of few structures, zinc has tetrahedral oxygen ligancy and the zinc-oxygen distance varies only slightly [25, 34]. It was concluded that zinc had a

Interestingly, the heat treatment of glasses at temperature higher than their glass transition (Tg) or crystallization temperature (Tc) improves the electrical conductivity of the glassy compound due to the "structural relaxation" of the glass network [36]. ZnO is widely used in glass production because it improves the glass quality by enhancing mechanical properties and chemical durability and by reducing the thermal expansion [7]. Zinc is a microelement that plays an important role in the bone formation and mineralization. Zinc containing glasses and glass ceramics has been developed for bone engineering applications [7]. The small size of Zn ion (0.74 Å)

Moreover, MgO oxide is of interest from a biological viewpoint because Mg2+ is known to play a physiological role in positively influencing bone strength which can

The bioactive behavior of magnesium rich glasses is identified as their ability to react chemically with living tissues, forming with them mechanically strong and lasting bonds. These bone bondings are attributed to the formation of an apatite-like layer on the glass surface, with composition and structure equivalent to the mineral

In fact, bioactive glasses have received special attention due to their better bone

Furthermore, by increasing the concentration of modifier oxides, electrical con-

Glasses containing transition metal oxides possess interesting electronic, optic and magnetic properties due to the ability to exist in more than one valence state. However, the electronic conduction of these glassy compounds is resulting from the

Compared with phosphate glasses, silicate glasses exhibit superior chemical resistance which makes them compatible with the fabrication process in the devel-

Silicate glasses are an attractive host matrix for transition metal ions due to their excellent optical and mechanical properties, good chemical durability, good chemical stability and low thermal expansion coefficient leading to strong thermal resistance [6]. Silicate glasses have many advantages rather than phosphate glasses. Silicatebased glasses are chemically durable, thermally stable and optically transparent at excitation and lasing wavelength. However, the higher viscosity of these glasses allows the glass to be formed without crystallization process. In addition, these amorphous materials are useful in optics as lenses or beam splitters in optical telecommunications, micro- and optoelectronics and in near IR-windows [6, 39, 40].

bonding ability in vivo. Due to their good bioactive and tailorable degradation properties, these glasses can be used for various biomedical applications such as

ductivity of the glass increases. This property was probably influenced by the structural changes resulting from the disruption of the glass network which affected the mobility of the cations and anions when the modifying oxide was progressively

electronic transfer of cation that exists in different valence sates [4, 38].

Glasses of the (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mn, Mg)

(0 ≤ x ≤ 33 mol%), (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) (0 ≤ x ≤ 33 mol%), and

coordination number of six in borate and silicate glasses [34].

*Contemporary Topics about Phosphorus in Biology and Materials*

helps it to locate itself into smaller cavities of the network [37].

be substituted into apatites [13, 18].

bone graft, filler, dental, implant coating.

phase of bone.

introduced [38].

opment of optical devices [3].

**2. Experimental procedures**

**2.1 Glass preparation**

**170**

Phosphorus, sodium, magnesium, zinc, manganese and silica were analyzed by inductively coupled plasma atomic emission spectroscopy (Jobin Yvon Ultra C).

#### **2.3 Physical properties**

#### *2.3.1 Density and molar volume*

Density of glass is a strong function of its composition and its intrinsic property which shed light on the short range structure of the glassy material [4, 37]. This work presents a series of glasses with various amounts of modifier oxides. These modifies, depending on their polarity and size, tend to occupy the interstices within the network and form new bonds resulting in a change in the structure and properties of the glass [4, 37].

Glass density measurements have been determined using the standard Archimedes method using diethyl orthophthalate as immersion fluid. The relative error of these measurements is 3%.

The molar volume of glasses has been calculated from the density (Vm = M/ρ) and the molar weight.

For (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mn, Mg) where 3 ≤ O/P ≤ 3.49 and (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) with O/P = 3 (0 ≤ x ≤ 33 mol%) series glasses, the density increases gradually with the incorporation of MO oxide. The increase in density indicates that the MO oxide reticulate the vitreous network because P▬O▬M bond are more ionic than P▬O▬P [11–13, 15, 29, 32, 33, 42].

**Table 2** shows that the molar volume deceases monotonically with the increase of ZnO content for phosphate glasses.

The decrease in the molar volume for (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mn, Mg) where 3 ≤ O/P ≤ 3.49 and (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) with O/P = 3 (0 ≤ x ≤ 33 mol%) series glasses could be explained by the higher field ΔF (ΔF = Z/r2 ; with z is the valence cation and r is the ionic radius) of M2+ compared to that of Na<sup>+</sup> [11–13, 15, 29, 32, 33, 42].

The decrease in the molar volume is extensively related to structural changes due to the incorporation of MO oxide that disrupted the average chain length of metaphosphate resulting from the following reaction:

$$\text{P} \blacksquare \text{O} \blacksquare \text{P} + \text{MO} \rightarrow (\text{2PO})\text{M} \tag{1}$$

The glass transition temperatures were determined on 40–50 mg of samples using DSC-ATD Netzsch 404 PC with a 10°C/min heating rate (accuracy �5°C). With increasing MO content, glass transition temperature, Tg, increases linearly

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

This behavior is undoubtedly corresponding to some changes in the nature of bonding in the structural network. This parameters is strictly related to the bond strength of the glass network which can be explained in terms of bond length (which is the charge divided by the square of the cation-oxygen distance) affected by the cation field strength resulting in a higher of Tg values [11–13, 15, 29, 32, 33, 42]. These variations indicate the progressive increase of the reticulation and the rigidity of the glass network by gathering the non-bridging oxygen atoms (NBO) with the increase of MO proportion. As a result the formation of P▬O▬M bonds suggesting the increase in the rigidity and the compactness of the structure that

A similar behavior has been observed for zinc phosphate-silicate glasses

According to Dietezel, the thermal stability of glasses (ΔT) can be expressed by the temperature difference between Tg and Tc, ΔT=Tc–Tg, in which Tg and Tc are the glass transition and crystallization temperature. Increasing ΔT delays the nucle-

Inspecting these data, one can note that the undoped ZnO and SiO2 oxide glass matrix has the lowest thermal stability indicating a tendency towards crystallization

ΔT increases when SiO2 oxide is progressively added, indicating a better stability of the glass. The larger value of ΔT, the stronger is the inhibition to nucleation and

The ability to control the physical properties of glasses, e.g., the refractive index, by variation in glass composition suggests the feasibility of chemically controlling

For glassy compounds, refractive index is a fundamental parameter that strongly relevant to optical devices performance and reliability in the basic elements in all optical instruments. Hence, a large number of researchers have carried out investigations to ascertain the relation between refractive index and glass composition [10]. This parameter is one of the fundamental properties of materials, because it is closely related to the electric polarizability of ions and the local field inside the

The carat eristic feature of phosphate glasses is the low value of the refractive index that is in the order of 1.49. The variation of this quantity for zinc phosphatebased silicate glasses is presented in **Table 4**. Inspecting these data, one can note that n increases from 1.44 to 1.49 when x rises from 0 to 10 mol% of SiO2 oxide which suggests that the refractive index of glassy compounds depends essentially on

In addition to density, many parameters can prevails the refractive index such as density, polarizability of the first neighbor ions coordinated with it (anion), coordination number of ion, electronic polarizability of the oxide ion and optical basicity [3, 10, 11]. The molar refractivity (Rm) was estimated from the refractive index and

*<sup>n</sup>*ð Þ <sup>2</sup> <sup>þ</sup> <sup>2</sup> *Vm* (2)

the molar volume (Vm) using the Lorenz-Lorenz Equation [3, 10, 11]:

*Rm* <sup>¼</sup> *<sup>n</sup>*ð Þ <sup>2</sup> � <sup>1</sup>

for all glass compositions as mentioned **Table 2**.

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

ameliorate the chemical durability of glasses.

ation process, indicating a better stability of the glass [29].

crystallization process as mentioned **Table 3** [12, 13, 15, 29, 32, 33].

the materials according to the needs of a given application [3].

[11–13, 15, 29, 32, 33, 42].

as shown **Tables 2** and **3**.

material [3, 10, 11].

**173**

*2.3.3 Refractive index measurements*

the density of glass network [3, 10, 11].

The variation of these properties is closely related to the structural modification when M2+ ion is progressively introduced.

The effect of composition on the density and molar volume for zinc phosphatesilicate glass, having a general formula: (0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol), shows that the replacement of NaPO3 by SiO2 oxide induces a decrease of density as mentioned **Table 3**. This is due to the lower molecular weight of SiO2 than that of NaPO3 (MSiO2 <sup>¼</sup> 60 gmol�<sup>1</sup> , MNaPO3 <sup>¼</sup> 102 gmol�<sup>1</sup> ) [11–13, 15, 29, 32, 33, 42].

As the same for the molar volume, this quantity decreases monotonically with the incorporation of SiO2 oxide (**Table 3**). This variation indicates that SiO2 oxide reticulates the vitreous network suggesting the increase in the rigidity of the structure.

Furthermore, the regular decrease in the molar volume is closely related to the nature of bending in the glass structure, because P▬O▬Si are more ionic than P▬O▬P bridges, suggesting the compactness of the vitreous network [11–13, 15, 29, 32, 33, 42].

#### *2.3.2 DSC investigations*

Generally, the glass transition phenomenon occurs due to the increasing viscosity of the overcooled liquids so Tg strongly depends on the polymerization ratio of the network [40].


**Table 2.**

*Density, molar volume, glass composition, glass transition temperature Tg, Tc, ΔT of (50-x/2)Na2O-xZnO- (50-x/2)P2O5 (0* ≤ *x* ≤ *33 mol%) phosphate glasses.*


**Table 3.**

*Density, molar volume, refractive index, glass composition, glass transition temperature Tg, Tc, ΔT of (0.9-x) NaPO3-xSiO2-0.1 ZnO (0* ≤ *x* ≤ *0.1 mol) glass series.*

#### *Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

The glass transition temperatures were determined on 40–50 mg of samples using DSC-ATD Netzsch 404 PC with a 10°C/min heating rate (accuracy �5°C).

With increasing MO content, glass transition temperature, Tg, increases linearly for all glass compositions as mentioned **Table 2**.

This behavior is undoubtedly corresponding to some changes in the nature of bonding in the structural network. This parameters is strictly related to the bond strength of the glass network which can be explained in terms of bond length (which is the charge divided by the square of the cation-oxygen distance) affected by the cation field strength resulting in a higher of Tg values [11–13, 15, 29, 32, 33, 42].

These variations indicate the progressive increase of the reticulation and the rigidity of the glass network by gathering the non-bridging oxygen atoms (NBO) with the increase of MO proportion. As a result the formation of P▬O▬M bonds suggesting the increase in the rigidity and the compactness of the structure that ameliorate the chemical durability of glasses.

A similar behavior has been observed for zinc phosphate-silicate glasses [11–13, 15, 29, 32, 33, 42].

According to Dietezel, the thermal stability of glasses (ΔT) can be expressed by the temperature difference between Tg and Tc, ΔT=Tc–Tg, in which Tg and Tc are the glass transition and crystallization temperature. Increasing ΔT delays the nucleation process, indicating a better stability of the glass [29].

Inspecting these data, one can note that the undoped ZnO and SiO2 oxide glass matrix has the lowest thermal stability indicating a tendency towards crystallization as shown **Tables 2** and **3**.

ΔT increases when SiO2 oxide is progressively added, indicating a better stability of the glass. The larger value of ΔT, the stronger is the inhibition to nucleation and crystallization process as mentioned **Table 3** [12, 13, 15, 29, 32, 33].

#### *2.3.3 Refractive index measurements*

The ability to control the physical properties of glasses, e.g., the refractive index, by variation in glass composition suggests the feasibility of chemically controlling the materials according to the needs of a given application [3].

For glassy compounds, refractive index is a fundamental parameter that strongly relevant to optical devices performance and reliability in the basic elements in all optical instruments. Hence, a large number of researchers have carried out investigations to ascertain the relation between refractive index and glass composition [10].

This parameter is one of the fundamental properties of materials, because it is closely related to the electric polarizability of ions and the local field inside the material [3, 10, 11].

The carat eristic feature of phosphate glasses is the low value of the refractive index that is in the order of 1.49. The variation of this quantity for zinc phosphatebased silicate glasses is presented in **Table 4**. Inspecting these data, one can note that n increases from 1.44 to 1.49 when x rises from 0 to 10 mol% of SiO2 oxide which suggests that the refractive index of glassy compounds depends essentially on the density of glass network [3, 10, 11].

In addition to density, many parameters can prevails the refractive index such as density, polarizability of the first neighbor ions coordinated with it (anion), coordination number of ion, electronic polarizability of the oxide ion and optical basicity [3, 10, 11]. The molar refractivity (Rm) was estimated from the refractive index and the molar volume (Vm) using the Lorenz-Lorenz Equation [3, 10, 11]:

$$R\_m = \frac{(n^2 - 1)}{(n^2 + 2)} V\_m \tag{2}$$

field ΔF (ΔF = Z/r2

compared to that of Na<sup>+</sup> [11–13, 15, 29, 32, 33, 42].

*Contemporary Topics about Phosphorus in Biology and Materials*

metaphosphate resulting from the following reaction:

when M2+ ion is progressively introduced.

**Glass composition X Density**

*(50-x/2)P2O5 (0* ≤ *x* ≤ *33 mol%) phosphate glasses.*

0 0.02 0.04 0.06 0.08 0.1

*NaPO3-xSiO2-0.1 ZnO (0* ≤ *x* ≤ *0.1 mol) glass series.*

**Glass composition x Density**

(0.9-x)NaPO3-xSiO2-

0.1 ZnO

**Table 3.**

**172**

**(gcm**�**<sup>3</sup> )**

2.60 � 0.10 2.58 � 0.10 2.57 � 0.10 2.55 � 0.10 2.55 � 0.10 2.53 � 0.10

NaPO3 (MSiO2 <sup>¼</sup> 60 gmol�<sup>1</sup>

*2.3.2 DSC investigations*

(50-x/2)Na2O-xZnO-(50-x/2)

the network [40].

P2O5

**Table 2.**

; with z is the valence cation and r is the ionic radius) of M2+

P▬O▬P þ MO ! ð Þ 2PO M (1)

) [11–13, 15, 29, 32, 33, 42].

The decrease in the molar volume is extensively related to structural changes due to the incorporation of MO oxide that disrupted the average chain length of

The variation of these properties is closely related to the structural modification

As the same for the molar volume, this quantity decreases monotonically with the incorporation of SiO2 oxide (**Table 3**). This variation indicates that SiO2 oxide reticulates the vitreous network suggesting the increase in the rigidity of the structure. Furthermore, the regular decrease in the molar volume is closely related to the nature of bending in the glass structure, because P▬O▬Si are more ionic than P▬O▬P bridges, suggesting the compactness of the vitreous network [11–13, 15, 29, 32, 33, 42].

Generally, the glass transition phenomenon occurs due to the increasing viscosity of the overcooled liquids so Tg strongly depends on the polymerization ratio of

> **Vm (cm<sup>3</sup> mol**�**<sup>1</sup>**

42.00 � 1.30 41.00 � 1.23 38.15 � 1.14 36.63 � 1.10 35.60 � 1.10 34.00 � 1.02 32.70 � 1.00 32.00 � 1.00

**)**

280 � 5 - 285 � 5 - 287 � 5 294 � 5 306 � 5 314 � 5

**Tg (°C) Tc (°C) ΔT (°C)**

290 � 5 - 371 � 5 - 368 � 5 439 � 5 456 � 5 446 � 5

**n Tg (°C) Tc (°C) ΔT (°C)**

**(g cm**�**<sup>3</sup> )**

2.43 � 0.07 2.47 � 0.07 2.62 � 0.08 2.70 � 0.08 2.75 � 0.08 2.85 � 0.09 2.93 � 0.09 2.98 � 0.09

*Density, molar volume, glass composition, glass transition temperature Tg, Tc, ΔT of (50-x/2)Na2O-xZnO-*

**Vm (cm<sup>3</sup> mol**�**<sup>1</sup>**

38.50 � 1.20 38.35 � 1.20 38.24 � 1.20 38.13 � 1.10 38.00 � 1.10 37.80 � 1.10

*Density, molar volume, refractive index, glass composition, glass transition temperature Tg, Tc, ΔT of (0.9-x)*

**)**

1.44 � 0.05 1.45 � 0.05 1.46 � 0.05 1.47 � 0.05 1.48 � 0.05 1.49 � 0.05

The effect of composition on the density and molar volume for zinc phosphatesilicate glass, having a general formula: (0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol), shows that the replacement of NaPO3 by SiO2 oxide induces a decrease of density as mentioned **Table 3**. This is due to the lower molecular weight of SiO2 than that of

, MNaPO3 <sup>¼</sup> 102 gmol�<sup>1</sup>


**Table 4.**

*Refractive index, molar refractivity (Rm), molar electronic polarizability (αm) and band gap energy of (0.9-x) NaPO3-xSiO2-0.1ZnO (0* ≤ *x* ≤ *0.1 Mol) glass series.*

The molar electronic polarizability α<sup>m</sup> was calculated using the relation of Clasius-Mosotti as follows [3, 10, 11]:

$$a\_m = \frac{3}{4} \Pi \text{INR}\_m \tag{3}$$

Furthermore, the asymmetric and symmetric stretching vibrations of P▬O▬P

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

FTIR spectra of (50-x/2)Na2O-xZnO-(50-x/2)P2O5 and (0.9-x)NaPO3-xSiO2-

As MO oxide is introduced, the asymmetric band of PO2 shifts from 1280 cm<sup>1</sup> shift to lower frequency as showed **Figures 1a** and **b** indicating the depolymeriza-

For higher ZnO content, FTIR spectra revealed the displacement of the asymmetric stretching mode vibration of the P▬O▬P band from 880 to 920 cm<sup>1</sup> when x rises from 0 to 33 mol%. This result can be correlated to the increase in the covalence character of P▬O▬P bridges when monovalent cation Na<sup>+</sup> was replaced

It may be also attributed to the shortening of phosphate chain length due to the

The FTIR spectra of NaPO3 glasses revealed also two bands in the frequency range 780–720 cm<sup>1</sup> which are attributed to the presence of two P▬O▬P bridges in

32, 33]. However, it is interesting to note that for the series glasses containing 30 and 33 mol% ZnO FTIR spectra exhibit only a single band at 740 cm<sup>1</sup> assigned to

On the other hand, this result could be explained by disruption of the infinite metaphosphate chains when MO oxide is gradually incorporated suggesting the

Similar FTIR spectra have been recorded for zinc phosphate-based silicate glasses, with a general formula: (0.9-x)NaPO3-xSiO2-0.1ZnO, (0 ≤ x ≤ 0.1 mol) as

some bands assigned to phosphate-silicate glasses in the range 1000–1300 cm<sup>1</sup> as shown **Figure 1b**. The asymmetric stretching vibration bands of silicate and phosphate

attributed to the symmetric stretching vibration of O▬Si▬O in metasilicate (Q2

tion of the infinite metaphosphate chains with the addition of SiO2 oxide.

could be attributed to the reduction of infinite phosphate groups (P2O6

SiO4 tetrahedron shared two oxygen with their neighbor (Si▬O▬2NBO). The sym-

vibration of Si▬O▬Si and O▬Si▬O bonds is around 460 cm<sup>1</sup> [12, 13, 15, 29, 32, 33]. When SiO2 is incorporated, FTIR spectra revealed the displacement of the asymmetric stretching mode vibration of PO2 band from 1280 to 1250 cm<sup>1</sup> when x increases from 0 to 10 mol%. This result can be probably due to the depolymeriza-

Furthermore, for 0.8 NaPO3-0.1 SiO2-0.1 ZnO glass composition, FTIR spectrum revealed the appearance of a only a single band at 760 cm<sup>1</sup> assigned to P▬O▬P

<sup>4</sup> and PO4

changes have been observed for magnesium and manganese phosphate glasses that

Raman spectroscopy is an adequate technique for the analysis of glass matrix structure. Raman bands are generally characteristics of structures involving chains

The FTIR spectra of zinc phosphate-based silicate glasses revealed the appearance of

32, 33]. These spectral changes depend essentially on the glass composition.

higher field strength and the size of the metallic cation, when the ratio O/P increases for (50-x/2)Na2O-xZnO-(50-x/2)P2O5 [12, 13, 15, 29, 32, 33].

. The deformation mode of P▬O▬(PO4

<sup>2</sup> groups (**Figure 1a** and **b**) [12, 13, 15, 29,

<sup>2</sup>)<sup>∞</sup> into short phosphate groups such as:

. It seems that a smaller band at 840 cm<sup>1</sup> is

<sup>4</sup>) as mentioned **Figure 1b**. These spectral

<sup>3</sup> [12, 13, 15, 29, 32, 33].

<sup>4</sup> (**Figure 1a**) [12, 13, 15, 29,

) when

. In addition, the bending

<sup>2</sup>) into

<sup>3</sup>)

bands are around 880, 780 and 720 cm<sup>1</sup>

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

metaphosphate chains based on (P2O6)

depolymerization of the skeleton of (P2O6

P2O7

<sup>4</sup> and PO4

mentioned **Figure 1b**.

tetrahedron is located at 1100 cm<sup>1</sup>

bands in phosphate dimers (P2O7

*2.4.2 Raman spectroscopy*

**175**

shorter phosphate groups such as: P2O7

groups at 535 and 480 cm<sup>1</sup> [12, 13, 15, 29, 32, 33].

tion of phosphate chains when x increases [12, 13, 15, 29, 32, 33].

by divalent cation (such as Zn2+) [12, 13, 15, 29, 32, 33].

the P▬O▬P linkage in pyrophosphate group (P2O7)

<sup>3</sup> [12, 13, 15, 29, 32, 33].

metric stretching vibration band Si▬O▬Si is about 1080 cm<sup>1</sup>

0.1ZnO glasses are shown in **Figures 1a** and **b**.

where the value of <sup>3</sup> 4 QN is known as the Lorentz function and N is the Avogadro number. **Table 4** reports the values of Rm and αm. These parameters increase gradually with the incorporation of SiO2 oxide.

**Table 4** shows that Rm increases from 10.23 to 10.60 m3 mol�<sup>1</sup> and α<sup>m</sup> are between 4.06 and 4.20 Å. These variations indicate that the refractive index as a function of both density and molar electric polarizability of glassy compounds [3, 10, 11].

In the present work, we found that the refractive index (n) depends on the ratio (*<sup>α</sup><sup>m</sup> Vm*). This quantity shows that the refractive index (n) of the studied glasses increases linearly versus (*<sup>α</sup><sup>m</sup> Vm*) ratio. This variation can be probably due to the electronic polarizability of oxide ions.

For Na2O ionic-based glasses, the polarizability of oxygen ions has the smaller value (*αO*2� = 2.45 Å) compared to copper rich glasses [3, 10, 11].

Duffy et al. suggested that increasing the optical basicity (<sup>Λ</sup> <sup>¼</sup> <sup>1</sup>*:*67 1 � <sup>1</sup> *αO*2� � �) indicates an increase in the effective electronic density of the oxide ions and

accordingly, increasing covalency in the oxygen-cation bonding [3, 10, 11].

The decrease in the molar volume for zinc-based phosphate-silicate glasses induces an increase in the rigidity and the compactness of the vitreous network, when SiO2 oxide is progressively introduced, because Si▬O▬P bonds are more ionic that P▬O▬P.

#### **2.4 Spectroscopic analysis**

#### *2.4.1 FTIR spectroscopy*

Infrared spectra of the glass series have been recorded by Perkin-Elmer (FTIR 2000) spectrometer using KBr pellets in the frequency range 400–4000 cm�<sup>1</sup> at room temperature. The samples were prepared by grinding about 9 mg of glass powder with 300 mg of spectroscopic grade dried KBr.

For undoped zinc phosphate glasses, NaPO3, FTIR spectrum revealed an asymmetric and symmetric stretching vibration band of PO2 groups in metaphosphate chains situated respectively at 1280 and 1150 cm�<sup>1</sup> . The asymmetric and symmetric stretching vibration bands of PO3 chain end groups situated respectively at 1100 and 1000 cm�<sup>1</sup> [12, 13, 15, 29, 32, 33].

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

Furthermore, the asymmetric and symmetric stretching vibrations of P▬O▬P bands are around 880, 780 and 720 cm<sup>1</sup> . The deformation mode of P▬O▬(PO4 <sup>3</sup>) groups at 535 and 480 cm<sup>1</sup> [12, 13, 15, 29, 32, 33].

FTIR spectra of (50-x/2)Na2O-xZnO-(50-x/2)P2O5 and (0.9-x)NaPO3-xSiO2- 0.1ZnO glasses are shown in **Figures 1a** and **b**.

As MO oxide is introduced, the asymmetric band of PO2 shifts from 1280 cm<sup>1</sup> shift to lower frequency as showed **Figures 1a** and **b** indicating the depolymerization of phosphate chains when x increases [12, 13, 15, 29, 32, 33].

For higher ZnO content, FTIR spectra revealed the displacement of the asymmetric stretching mode vibration of the P▬O▬P band from 880 to 920 cm<sup>1</sup> when x rises from 0 to 33 mol%. This result can be correlated to the increase in the covalence character of P▬O▬P bridges when monovalent cation Na<sup>+</sup> was replaced by divalent cation (such as Zn2+) [12, 13, 15, 29, 32, 33].

It may be also attributed to the shortening of phosphate chain length due to the higher field strength and the size of the metallic cation, when the ratio O/P increases for (50-x/2)Na2O-xZnO-(50-x/2)P2O5 [12, 13, 15, 29, 32, 33].

The FTIR spectra of NaPO3 glasses revealed also two bands in the frequency range 780–720 cm<sup>1</sup> which are attributed to the presence of two P▬O▬P bridges in metaphosphate chains based on (P2O6) <sup>2</sup> groups (**Figure 1a** and **b**) [12, 13, 15, 29, 32, 33]. However, it is interesting to note that for the series glasses containing 30 and 33 mol% ZnO FTIR spectra exhibit only a single band at 740 cm<sup>1</sup> assigned to the P▬O▬P linkage in pyrophosphate group (P2O7) <sup>4</sup> (**Figure 1a**) [12, 13, 15, 29, 32, 33]. These spectral changes depend essentially on the glass composition.

On the other hand, this result could be explained by disruption of the infinite metaphosphate chains when MO oxide is gradually incorporated suggesting the depolymerization of the skeleton of (P2O6 <sup>2</sup>)<sup>∞</sup> into short phosphate groups such as: P2O7 <sup>4</sup> and PO4 <sup>3</sup> [12, 13, 15, 29, 32, 33].

Similar FTIR spectra have been recorded for zinc phosphate-based silicate glasses, with a general formula: (0.9-x)NaPO3-xSiO2-0.1ZnO, (0 ≤ x ≤ 0.1 mol) as mentioned **Figure 1b**.

The FTIR spectra of zinc phosphate-based silicate glasses revealed the appearance of some bands assigned to phosphate-silicate glasses in the range 1000–1300 cm<sup>1</sup> as shown **Figure 1b**. The asymmetric stretching vibration bands of silicate and phosphate tetrahedron is located at 1100 cm<sup>1</sup> . It seems that a smaller band at 840 cm<sup>1</sup> is attributed to the symmetric stretching vibration of O▬Si▬O in metasilicate (Q2 ) when SiO4 tetrahedron shared two oxygen with their neighbor (Si▬O▬2NBO). The symmetric stretching vibration band Si▬O▬Si is about 1080 cm<sup>1</sup> . In addition, the bending vibration of Si▬O▬Si and O▬Si▬O bonds is around 460 cm<sup>1</sup> [12, 13, 15, 29, 32, 33].

When SiO2 is incorporated, FTIR spectra revealed the displacement of the asymmetric stretching mode vibration of PO2 band from 1280 to 1250 cm<sup>1</sup> when x increases from 0 to 10 mol%. This result can be probably due to the depolymerization of the infinite metaphosphate chains with the addition of SiO2 oxide.

Furthermore, for 0.8 NaPO3-0.1 SiO2-0.1 ZnO glass composition, FTIR spectrum revealed the appearance of a only a single band at 760 cm<sup>1</sup> assigned to P▬O▬P bands in phosphate dimers (P2O7 <sup>4</sup>) as mentioned **Figure 1b**. These spectral changes have been observed for magnesium and manganese phosphate glasses that could be attributed to the reduction of infinite phosphate groups (P2O6 <sup>2</sup>) into shorter phosphate groups such as: P2O7 <sup>4</sup> and PO4 <sup>3</sup> [12, 13, 15, 29, 32, 33].

#### *2.4.2 Raman spectroscopy*

Raman spectroscopy is an adequate technique for the analysis of glass matrix structure. Raman bands are generally characteristics of structures involving chains

The molar electronic polarizability α<sup>m</sup> was calculated using the relation of

*Refractive index, molar refractivity (Rm), molar electronic polarizability (αm) and band gap energy of (0.9-x)*

0.9NaPO3-0.1ZnO 1.44 4.06 10.23 1.50 1.05 0.88NaPO3-0.02SiO2-0.1ZnO 1.45 4.07 10.30 1.70 1.06 0.86NaPO3-0.04SiO2-0.1ZnO 1.46 4.13 10.40 2.00 1.08 0.84NaPO3-0.06SiO2-0.1ZnO 1.47 4.15 10.50 2.25 1.09 0.82NaPO3-0.08SiO2-0.1ZnO 1.48 4.18 10.55 2.35 1.10 0.8NaPO3-0.1SiO2-0.1ZnO 1.49 4.20 10.60 2.35 1.11

**) Rm (cm<sup>3</sup> mol**�**<sup>1</sup>**

Π*NRm* (3)

**) Eopt (ev) (αm/Vm)** � **<sup>10</sup>**�**<sup>25</sup>**

*αO*2� � �

. The asymmetric and symmetric

)

QN is known as the Lorentz function and N is the Avogadro

*Vm*) ratio. This variation can be probably due to the

*<sup>α</sup><sup>m</sup>* <sup>¼</sup> <sup>3</sup> 4

number. **Table 4** reports the values of Rm and αm. These parameters increase

*Vm*). This quantity shows that the refractive index (n) of the studied glasses

value (*αO*2� = 2.45 Å) compared to copper rich glasses [3, 10, 11].

powder with 300 mg of spectroscopic grade dried KBr.

chains situated respectively at 1280 and 1150 cm�<sup>1</sup>

and 1000 cm�<sup>1</sup> [12, 13, 15, 29, 32, 33].

**Table 4** shows that Rm increases from 10.23 to 10.60 m3 mol�<sup>1</sup> and α<sup>m</sup> are between 4.06 and 4.20 Å. These variations indicate that the refractive index as a function of both density and molar electric polarizability of glassy compounds [3, 10, 11].

In the present work, we found that the refractive index (n) depends on the ratio

For Na2O ionic-based glasses, the polarizability of oxygen ions has the smaller

Infrared spectra of the glass series have been recorded by Perkin-Elmer (FTIR 2000) spectrometer using KBr pellets in the frequency range 400–4000 cm�<sup>1</sup> at room temperature. The samples were prepared by grinding about 9 mg of glass

For undoped zinc phosphate glasses, NaPO3, FTIR spectrum revealed an asymmetric and symmetric stretching vibration band of PO2 groups in metaphosphate

stretching vibration bands of PO3 chain end groups situated respectively at 1100

Duffy et al. suggested that increasing the optical basicity (<sup>Λ</sup> <sup>¼</sup> <sup>1</sup>*:*67 1 � <sup>1</sup>

indicates an increase in the effective electronic density of the oxide ions and accordingly, increasing covalency in the oxygen-cation bonding [3, 10, 11]. The decrease in the molar volume for zinc-based phosphate-silicate glasses induces an increase in the rigidity and the compactness of the vitreous network, when SiO2 oxide is progressively introduced, because Si▬O▬P bonds are more

Clasius-Mosotti as follows [3, 10, 11]:

*NaPO3-xSiO2-0.1ZnO (0* ≤ *x* ≤ *0.1 Mol) glass series.*

**Glass composition n α<sup>m</sup> (Å3**

*Contemporary Topics about Phosphorus in Biology and Materials*

4

gradually with the incorporation of SiO2 oxide.

where the value of <sup>3</sup>

increases linearly versus (*<sup>α</sup><sup>m</sup>*

ionic that P▬O▬P.

*2.4.1 FTIR spectroscopy*

**2.4 Spectroscopic analysis**

electronic polarizability of oxide ions.

(*<sup>α</sup><sup>m</sup>*

**174**

**Table 4.**

of linked tetrahedral that may be found in crystalline, glassy phosphates and silicate because it can detect the local changes in the environment of Si▬O▬Si and P▬O▬P bonds [1, 2].

For undoped zinc phosphate glasses, Raman spectrum revealed a large band around 1274 cm<sup>1</sup> and three weaker peaks at 1164, 685, and 380 respectively as

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

The bands located at 1274 and 1164 cm<sup>1</sup> are assigned to the asymmetric and

32, 33]. The large band at about 685 cm<sup>1</sup> is attributed to the symmetric vibration

metaphosphate chains [12, 13, 15, 29, 32, 33]. The low frequency attributed to the faint band at 380 cm<sup>1</sup> is related to the bending motion of phosphate polyhedral

With increasing MO content, we observe some decrease of the overall back-

be correlated to the distortion of P▬O▬P band which induces the shortening of the infinite metaphosphate chains suggesting the formation of pyrophosphate groups

From **Figure 2a**, it seems that the intensity of bands located at 1164 and 685 cm<sup>1</sup> decrease when MO oxide is progressively introduced. However, the Raman spectra revealed the displacement of these bands to higher frequencies to 1180 (d, e) and 780 cm<sup>1</sup> (d, e, f) with 30 and 33 mol% of MO level. This result can be probably due to the higher π character of P▬NBO bands that induces the depolymerization of

Similar Raman spectroscopic analysis have been recorded for (0.9-x) NaPO3-

The 31P MAS-NMR spectra of (50-x/2)Na2O-xZnO-(50-x/2)P2O5 glasses are

*31P MAS-NMR spectra of (50-x/2)Na2O-xZnO-(50-x/2)P2O5 glasses: (a) 0 mol% ZnO, (b) 5 mol% ZnO, (c) 10 mol% ZnO, (d) 15 mol% ZnO, (e) 20 mol% ZnO, (f) 25 mol% ZnO, (g) 30 mol% ZnO, (h) 33 mol%*

The incorporation of SiO2 oxide to the phosphate network generates the appearance of the asymmetric band around 850 cm<sup>1</sup> attributed to Si▬O▬Si bending modes. The band located at 560 cm<sup>1</sup> is attributed to Si▬O▬Si intertetrahedral linkages obtained in calcium and magnesium rich silicate glasses in order to link the distorted metaphosphate groups when SiO2 oxide is added [12, 13, 15, 29, 32, 33].

) [12, 13, 15, 29,

. These spectral changes can

symmetric vibrations of PO2 groups in metaphosphate chains (Q<sup>2</sup>

of the bridging oxygen linking two PO4 tetrahedrons (P▬O▬P) in

) with the increase of the O/P ratio [12, 13, 15, 29, 32, 33].

infinite metaphosphate chains when MO oxide is progressively added.

xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) glass compositions as shown in **Figure 2b**.

ground located at 600–800 cm<sup>1</sup> and 1100–1300 cm<sup>1</sup>

shown in **Figure 2a**.

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

[12, 13, 15, 29, 32, 33].

*2.4.3 31P MAS-NMR spectroscopy*

shown in **Figure 3**.

**Figure 3.**

*ZnO.* **177**

(Q<sup>1</sup>

The Raman spectra were recorded on powder of glasses using a Labram HR800 micro Raman model operating in the 50–4000 cm<sup>1</sup> range at room temperature equipped with an internal He-Ne laser source (λ = 488 nm).

**Figure 2a** reported the Raman spectra of zinc phosphate glasses having a general formula (50-x/2)Na2O-xZnO-(50-x/2)P2O5 (0 ≤ x ≤ 33 mol%) with an O/P ratio varies from 3 to 3.49.

#### **Figure 2.**

*a. Raman spectra of (50-x/2)Na2O-xZnO-(50-x/2)P2O5 glasses: (a) 0 mol% ZnO, (b) 10 mol% ZnO, (c) 20 mol% ZnO, (d) 25 mol% ZnO, (e) 30 mol% ZnO, (f) 33 mol% ZnO. b. Raman spectra of (0.9-x) NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) glasses: (a) 0 mol SiO2, (b) 0.02 mol SiO2, (c) 0.04 mol SiO2, (d) 0.06 mol SiO2, (e) 0.08 mol SiO2, (f) 0.10 mol SiO2.*

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

For undoped zinc phosphate glasses, Raman spectrum revealed a large band around 1274 cm<sup>1</sup> and three weaker peaks at 1164, 685, and 380 respectively as shown in **Figure 2a**.

The bands located at 1274 and 1164 cm<sup>1</sup> are assigned to the asymmetric and symmetric vibrations of PO2 groups in metaphosphate chains (Q<sup>2</sup> ) [12, 13, 15, 29, 32, 33]. The large band at about 685 cm<sup>1</sup> is attributed to the symmetric vibration of the bridging oxygen linking two PO4 tetrahedrons (P▬O▬P) in metaphosphate chains [12, 13, 15, 29, 32, 33]. The low frequency attributed to the faint band at 380 cm<sup>1</sup> is related to the bending motion of phosphate polyhedral [12, 13, 15, 29, 32, 33].

With increasing MO content, we observe some decrease of the overall background located at 600–800 cm<sup>1</sup> and 1100–1300 cm<sup>1</sup> . These spectral changes can be correlated to the distortion of P▬O▬P band which induces the shortening of the infinite metaphosphate chains suggesting the formation of pyrophosphate groups (Q<sup>1</sup> ) with the increase of the O/P ratio [12, 13, 15, 29, 32, 33].

From **Figure 2a**, it seems that the intensity of bands located at 1164 and 685 cm<sup>1</sup> decrease when MO oxide is progressively introduced. However, the Raman spectra revealed the displacement of these bands to higher frequencies to 1180 (d, e) and 780 cm<sup>1</sup> (d, e, f) with 30 and 33 mol% of MO level. This result can be probably due to the higher π character of P▬NBO bands that induces the depolymerization of infinite metaphosphate chains when MO oxide is progressively added.

Similar Raman spectroscopic analysis have been recorded for (0.9-x) NaPO3 xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) glass compositions as shown in **Figure 2b**.

The incorporation of SiO2 oxide to the phosphate network generates the appearance of the asymmetric band around 850 cm<sup>1</sup> attributed to Si▬O▬Si bending modes. The band located at 560 cm<sup>1</sup> is attributed to Si▬O▬Si intertetrahedral linkages obtained in calcium and magnesium rich silicate glasses in order to link the distorted metaphosphate groups when SiO2 oxide is added [12, 13, 15, 29, 32, 33].

#### *2.4.3 31P MAS-NMR spectroscopy*

The 31P MAS-NMR spectra of (50-x/2)Na2O-xZnO-(50-x/2)P2O5 glasses are shown in **Figure 3**.

#### **Figure 3.**

*31P MAS-NMR spectra of (50-x/2)Na2O-xZnO-(50-x/2)P2O5 glasses: (a) 0 mol% ZnO, (b) 5 mol% ZnO, (c) 10 mol% ZnO, (d) 15 mol% ZnO, (e) 20 mol% ZnO, (f) 25 mol% ZnO, (g) 30 mol% ZnO, (h) 33 mol% ZnO.*

of linked tetrahedral that may be found in crystalline, glassy phosphates and silicate

The Raman spectra were recorded on powder of glasses using a Labram HR800 micro Raman model operating in the 50–4000 cm<sup>1</sup> range at room temperature

**Figure 2a** reported the Raman spectra of zinc phosphate glasses having a general formula (50-x/2)Na2O-xZnO-(50-x/2)P2O5 (0 ≤ x ≤ 33 mol%) with an O/P ratio

*a. Raman spectra of (50-x/2)Na2O-xZnO-(50-x/2)P2O5 glasses: (a) 0 mol% ZnO, (b) 10 mol% ZnO, (c) 20 mol% ZnO, (d) 25 mol% ZnO, (e) 30 mol% ZnO, (f) 33 mol% ZnO. b. Raman spectra of (0.9-x) NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) glasses: (a) 0 mol SiO2, (b) 0.02 mol SiO2, (c) 0.04 mol SiO2,*

*(d) 0.06 mol SiO2, (e) 0.08 mol SiO2, (f) 0.10 mol SiO2.*

because it can detect the local changes in the environment of Si▬O▬Si and

equipped with an internal He-Ne laser source (λ = 488 nm).

*Contemporary Topics about Phosphorus in Biology and Materials*

P▬O▬P bonds [1, 2].

varies from 3 to 3.49.

**Figure 2.**

**176**

The characteristic features of undoped zinc phosphate glasses are isotopic peaks at �21 and �6.88 ppm. The first one is attributed to the Q<sup>2</sup> tetrahedral sites in metaphosphate groups and the second is assigned to the Q<sup>1</sup> groups at the end of chain [12, 13, 15, 29, 32, 33].

*α ν*ð Þ¼ *A h<sup>ν</sup>* � *Eopt <sup>n</sup> hν*

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

• n: a constant which determines the type of the optical transition. For direct allowed transition *n* = 2 and in the case of indirect allowed transition n= <sup>1</sup>

For glassy materials, the indirect transitions are valid according to Tauc relations

The values of indirect optical band gap energy (Eopt) were determined by the

<sup>2</sup> = 0. This latter shows that the Eopt increases with the incorporation SiO2 from 1.5 to 2.35 eV. This quantity is not only influenced by the chemical composition also by the structural rearrangement in the glass matrix [4, 10, 11, 14, 16, 37, 41]. **Figure 5** shows clearly that the Eopt values dependent strongly on the composition of the glass also on the oxygen bonding in the vitreous network [4, 10, 11, 14, 16, 37, 41]. Any changes of oxygen bonding suggesting the formation of nonbridging oxygen (NBOs) causes a change of the absorption characteristics of the

The higher energy is required to excite an electron from bridging oxygen (BO)

*The (αhν)2 as a function of photon energy of hν of (0.9-x)NaPO3-xSiO2-0.1ZnO (0* ≤ *x* ≤ *0.1 mol) glasses: (a) 0 mol SiO2, (b) 0.02 mol SiO2, (c) 0.04 mol SiO2, (d) 0.06 mol SiO2, (e) 0.08 mol SiO2, (f) 0.10 mol SiO2. \*Obtaining the lines corresponding to the curves of (αhν)2 against photon energy (hν) is probably due to the*

than from non-bridging oxygen (NBO). As a result the increase in Eopt values

where

(αhν)

**Figure 4.**

**179**

*superposition effect for all the glass compositions.*

• A: an energy-independent constant

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

**Figure 4** represents the variation (αhν)

extrapolation of the linear region of (αhν)

NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) series glasses.

• Eopt: the optical band gap energy

[4, 10, 11, 14, 16, 37, 41].

[4, 10, 11, 14, 16, 37, 41].

glass [4, 10, 11, 14, 16, 37, 41].

[4, 10, 11, 14, 16, 37, 41].

(5)

2

<sup>2</sup> versus photon energy (hν) for (0.9-x)

<sup>2</sup> against photon energy (hν) plots at

Based on literature, the chemical shift at +1.4 ppm is attributed to NaPO3 chain end groups. From **Figure 3**, one can note the appearance of two isotopic peaks around 21–�18.80 ppm and �6.88–�3.90 ppm for the glass series.

When the MO oxide is introduced to the vitreous network, the intensity peak attributed to Q<sup>1</sup> tetrahedral sites increases and becomes the major spectral feature [12, 13, 15, 29, 32, 33]. These results are in good agreement for zinc phosphate glasses with the structural study of (100-x)NaPO3-xZnO glasses (0 ≤ x ≤ 33.3 mol%) performed by Montagne et al. [12, 13, 15, 29, 32, 33].

From **Figure 3,** it seem that 31P MAS-NMR spectra exhibit only single peak assigned to Q<sup>1</sup> tetrahedral sites attributed to pyrophosphate groups resulted from the distortion of metaphosphate chains when MO oxide is progressively introduced.

Furthermore, the phosphorus chemical shift depends essentially on the phosphorus-ligand bond (P-O) for phosphate compounds and the electronic density of the non-bridging oxygen (NBO). **Figure 3** mentioned that the Q<sup>2</sup> chemical shift becomes less shielded when MO is added. This decrease is probably due to the higher electronegativity of M2+ compared to Na<sup>+</sup> also the increase of π fraction of P-NBO resulting from the decondensation of phosphate chains when ZnO is incorporated. This suggests that Zn2+ ions are only bonded to pyrophosphate groups described by Q<sup>1</sup> tetrahedral sites. As a result the increase of shielding Q2 sites from �21 ppm in NaPO3 glass to �18.80 ppm in 33.5 Na2O-33 ZnO-33.5 P2O5 glass composition [12, 13, 15, 29, 32, 33].

#### *2.4.4 UV-VIS spectroscopy*

UV-VIS-NIR absorption spectra of the glassy compounds were carried out by means of Perkin-Elmer Lambda 950 spectrometer at room temperature under air. Optical measurements were recorded in the range of 200 and 1800 nm.

Optical absorption, particularly the absorption edge, is useful for the investigation of optically-induced transitions and for getting information about the band gap energy [4, 10, 11, 14, 16, 37, 41]. This parameter is very interesting for the applications of the materials to be studied. It is known that the optical transition occurs through the region between conduction and valence bands (optical band gap) directly or indirectly.

However, the optical transition involves an energy transfer caused by electron transitions between conduction and valence bands [4, 10, 11, 14, 16, 37, 41].

The optical absorption coefficient α(hν) of the prepared glasses was calculated at different wavelengths by using the relation [4, 10, 11, 14, 16, 37, 41]:

$$a = \frac{1}{d} \ln \left( \frac{I}{I\_0} \right) \tag{4}$$

where *d* represents the thickness of the glass composition and ln *<sup>I</sup> I*0 is the absorbance.

For the optical measurements, one can note the absence of the absorption sharp edge which characterizes the vitreous nature of the prepared glasses [4, 10, 11, 14, 16, 37, 41].

According to Davis and Mott, the expression of the absorption coefficient α (ν) as a function of photon energy (hν) for direct and indirect optical absorption, was given by the relation as follows:

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

$$a(\nu) = \frac{A\left(h\nu - E\_{\rm opt}\right)^n}{h\nu} \tag{5}$$

where

The characteristic features of undoped zinc phosphate glasses are isotopic peaks

Based on literature, the chemical shift at +1.4 ppm is attributed to NaPO3 chain end groups. From **Figure 3**, one can note the appearance of two isotopic peaks

When the MO oxide is introduced to the vitreous network, the intensity peak attributed to Q<sup>1</sup> tetrahedral sites increases and becomes the major spectral feature [12, 13, 15, 29, 32, 33]. These results are in good agreement for zinc phosphate glasses with the structural study of (100-x)NaPO3-xZnO glasses (0 ≤ x ≤ 33.3 mol%)

From **Figure 3,** it seem that 31P MAS-NMR spectra exhibit only single peak assigned to Q<sup>1</sup> tetrahedral sites attributed to pyrophosphate groups resulted from the distortion of metaphosphate chains when MO oxide is progressively introduced. Furthermore, the phosphorus chemical shift depends essentially on the phosphorus-ligand bond (P-O) for phosphate compounds and the electronic density of the non-bridging oxygen (NBO). **Figure 3** mentioned that the Q<sup>2</sup> chemical shift becomes less shielded when MO is added. This decrease is probably due to the higher electronegativity of M2+ compared to Na<sup>+</sup> also the increase of π fraction of P-NBO resulting from the decondensation of phosphate chains when ZnO is incorporated. This suggests that Zn2+ ions are only bonded to pyrophosphate groups described by Q<sup>1</sup> tetrahedral sites. As a result the increase of shielding Q2 sites from �21 ppm in NaPO3 glass to �18.80 ppm in 33.5 Na2O-33 ZnO-33.5 P2O5 glass

UV-VIS-NIR absorption spectra of the glassy compounds were carried out by means of Perkin-Elmer Lambda 950 spectrometer at room temperature under air.

Optical absorption, particularly the absorption edge, is useful for the investigation of optically-induced transitions and for getting information about the band gap energy [4, 10, 11, 14, 16, 37, 41]. This parameter is very interesting for the applications of the materials to be studied. It is known that the optical transition occurs through the region between conduction and valence bands (optical band gap)

However, the optical transition involves an energy transfer caused by electron

The optical absorption coefficient α(hν) of the prepared glasses was calculated at

*<sup>d</sup>* ln *<sup>I</sup> I*0 

For the optical measurements, one can note the absence of the absorption sharp edge which characterizes the vitreous nature of the prepared glasses [4, 10, 11, 14,

According to Davis and Mott, the expression of the absorption coefficient α (ν) as a function of photon energy (hν) for direct and indirect optical absorption, was

(4)

is the

*I*0 

transitions between conduction and valence bands [4, 10, 11, 14, 16, 37, 41].

*<sup>α</sup>* <sup>¼</sup> <sup>1</sup>

where *d* represents the thickness of the glass composition and ln *<sup>I</sup>*

different wavelengths by using the relation [4, 10, 11, 14, 16, 37, 41]:

Optical measurements were recorded in the range of 200 and 1800 nm.

at �21 and �6.88 ppm. The first one is attributed to the Q<sup>2</sup> tetrahedral sites in metaphosphate groups and the second is assigned to the Q<sup>1</sup> groups at the end of

around 21–�18.80 ppm and �6.88–�3.90 ppm for the glass series.

performed by Montagne et al. [12, 13, 15, 29, 32, 33].

*Contemporary Topics about Phosphorus in Biology and Materials*

chain [12, 13, 15, 29, 32, 33].

composition [12, 13, 15, 29, 32, 33].

*2.4.4 UV-VIS spectroscopy*

directly or indirectly.

absorbance.

16, 37, 41].

**178**

given by the relation as follows:


For glassy materials, the indirect transitions are valid according to Tauc relations [4, 10, 11, 14, 16, 37, 41].

**Figure 4** represents the variation (αhν) <sup>2</sup> versus photon energy (hν) for (0.9-x) NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol) series glasses.

The values of indirect optical band gap energy (Eopt) were determined by the extrapolation of the linear region of (αhν) <sup>2</sup> against photon energy (hν) plots at (αhν) <sup>2</sup> = 0. This latter shows that the Eopt increases with the incorporation SiO2 from 1.5 to 2.35 eV. This quantity is not only influenced by the chemical composition also by the structural rearrangement in the glass matrix [4, 10, 11, 14, 16, 37, 41]. **Figure 5** shows clearly that the Eopt values dependent strongly on the composition of the glass also on the oxygen bonding in the vitreous network [4, 10, 11, 14, 16, 37, 41]. Any changes of oxygen bonding suggesting the formation of nonbridging oxygen (NBOs) causes a change of the absorption characteristics of the glass [4, 10, 11, 14, 16, 37, 41].

The higher energy is required to excite an electron from bridging oxygen (BO) than from non-bridging oxygen (NBO). As a result the increase in Eopt values [4, 10, 11, 14, 16, 37, 41].

#### **Figure 4.**

*The (αhν)2 as a function of photon energy of hν of (0.9-x)NaPO3-xSiO2-0.1ZnO (0* ≤ *x* ≤ *0.1 mol) glasses: (a) 0 mol SiO2, (b) 0.02 mol SiO2, (c) 0.04 mol SiO2, (d) 0.06 mol SiO2, (e) 0.08 mol SiO2, (f) 0.10 mol SiO2. \*Obtaining the lines corresponding to the curves of (αhν)2 against photon energy (hν) is probably due to the superposition effect for all the glass compositions.*

the dissolution, mixing or dilution process. The integration of the raw signal deter-

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

Experiments were carried out by dissolving the same mass of solids (25 mg) in

For (50-x/2)Na2O-xZnO-50-x/2)P2O5 glass composition, it seems that the dissolution phenomenon is endothermic for lower ZnO content and becomes exothermic

Furthermore, the change in thermal signs is probably correlated to structural

These results were correlated to spectroscopic investigations which revealed the formation of pyrophosphate groups for zinc, manganese and magnesium phosphate glasses, resulting from the cleavage of the P-O-P bridges when the amount of MO

Calorimetric study of glasses has been carried out for several decades. However, the thermochemical investigations of glassy compounds have been considered using a thermodynamic approach based on the Miedema's model in order to evaluate the

In the case of glassy compounds, the knowledge of the formation enthalpy is an important chemical data which can be used to determine the Gibbs free energy of formation of the selected compounds and to have an idea about their stability. The determination of the formation enthalpy of (100-x)NaPO3-xZnO glass series involves the formation enthalpy of sodium trimetaphosphate, (NaPO3)3, crystal). Because of the very old value reported in literature [42], this quantity has

Sodium trimetaphosphate (NaPO3)3, was synthesized by thermal decomposition of sodium dihydrogen phosphate (NaH2PO4). This later was obtained by thermal

*Evolution of molar enthalpy of dissolution of (100-x)NaPO3-xZnO(0* ≤ *x* ≤ *33 mol%) glasses.*

)) are shown in **Figure 6**.

) suggesting the formation of pyro-

mined the heat dissolution of the studied compound.

modifications of metaphosphate groups (Q<sup>2</sup>

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

formation enthalpy of binary alloys [42].

been determined again by the same technique.

*3.2.1 Synthesis and characterization of samples*

The plots of heat dissolution of glasses (ΔsolH (kJ mol<sup>1</sup>

oxide is progressively increases [11–13, 15, 29, 32, 33, 42].

**3.2 Thermochemical study of zinc phosphate glasses**

when ZnO oxide is progressively incorporated in the vitreous network.

) when ZnO oxide is introduced.

4.5 ml of solvent.

phosphate units (Q<sup>1</sup>

**Figure 6.**

**181**

**Figure 5.** *Variation of optical band gap energy of (0.9-x)NaPO3-xSiO2-0.1ZnO (0* ≤ *x* ≤ *0.1 mol) glass series.*

When x increases from 0 to 10 mol% of SiO2, the optical band gap energy rises from 1.5 to 2.35 eV. This variation can be explained by the structural modifications which suggest the distortion of metaphosphate chains inducing the increase in the number of non-bridging oxygen (NBOs).

Because the NBOs bonds are predominantly ionic character and consequently have lower bond energies [34]. The higher value of the band gap energy revealed the increase of the cross-linking network due the introduction of SiO2 [4, 10, 11, 14, 16, 37, 41].

From **Figure 5**, it seems that the Eopt is in the order of 2.35 eV for 0.82 NaPO3-0.08 SiO2-0.1 ZnO and 0.8 NaPO3-0.1 SiO2-0.1 ZnO glass compositions. This result can be correlated to the structural changes due to the formation of P-O-Si ionic bands [11].

#### **3. Thermochemical study of phosphate glasses**

#### **3.1 Calorimetric dissolution of zinc, manganese and magnesium phosphate glasses**

The calorimetric study was performed by determining the energy resulting from the dissolution of the glasses in a suitable solvent [11–13, 15, 29, 32, 33, 42].

Phosphate glasses are soluble in mineral acids [11–13, 15, 29, 32, 33, 42]. Furthermore, the dissolution process has been carried out in order to find the suitable solvent which dissolves entirely the glassy compounds and should not give rise to any secondary phenomena.

For this purpose, our investigations were covered all the usual acids, bases and their mixtures such as: HNO3, HCl, NaOH, KOH, CH3COOH.

The calorimetric profile shows that the 4.5% weight of phosphoric acid solution is the best solvent for the thermochemical requirements of phosphate glasses.

The dissolution of phosphate glasses were recorded by means the C80 (SETARAM) at 25°C. This equipment possesses two identical cells: the reference and the measuring cell. The reference cell should contain only the solvent but the measuring cell was provided with the solid to be dissolved or the liquid to be mixture. The superior compartment contains the attack solution (solvent) which is tightly separated from the lower one by a movable cover.

The reference and the measuring cell are surrounded by thermoelectric piles with high performance. These latters permit to detect the heat flow resulted from *Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

the dissolution, mixing or dilution process. The integration of the raw signal determined the heat dissolution of the studied compound.

Experiments were carried out by dissolving the same mass of solids (25 mg) in 4.5 ml of solvent.

The plots of heat dissolution of glasses (ΔsolH (kJ mol<sup>1</sup> )) are shown in **Figure 6**. For (50-x/2)Na2O-xZnO-50-x/2)P2O5 glass composition, it seems that the dissolution phenomenon is endothermic for lower ZnO content and becomes exothermic when ZnO oxide is progressively incorporated in the vitreous network.

Furthermore, the change in thermal signs is probably correlated to structural modifications of metaphosphate groups (Q<sup>2</sup> ) suggesting the formation of pyrophosphate units (Q<sup>1</sup> ) when ZnO oxide is introduced.

These results were correlated to spectroscopic investigations which revealed the formation of pyrophosphate groups for zinc, manganese and magnesium phosphate glasses, resulting from the cleavage of the P-O-P bridges when the amount of MO oxide is progressively increases [11–13, 15, 29, 32, 33, 42].

#### **3.2 Thermochemical study of zinc phosphate glasses**

Calorimetric study of glasses has been carried out for several decades. However, the thermochemical investigations of glassy compounds have been considered using a thermodynamic approach based on the Miedema's model in order to evaluate the formation enthalpy of binary alloys [42].

In the case of glassy compounds, the knowledge of the formation enthalpy is an important chemical data which can be used to determine the Gibbs free energy of formation of the selected compounds and to have an idea about their stability.

The determination of the formation enthalpy of (100-x)NaPO3-xZnO glass series involves the formation enthalpy of sodium trimetaphosphate, (NaPO3)3, crystal). Because of the very old value reported in literature [42], this quantity has been determined again by the same technique.

#### *3.2.1 Synthesis and characterization of samples*

Sodium trimetaphosphate (NaPO3)3, was synthesized by thermal decomposition of sodium dihydrogen phosphate (NaH2PO4). This later was obtained by thermal

**Figure 6.** *Evolution of molar enthalpy of dissolution of (100-x)NaPO3-xZnO(0* ≤ *x* ≤ *33 mol%) glasses.*

When x increases from 0 to 10 mol% of SiO2, the optical band gap energy rises from 1.5 to 2.35 eV. This variation can be explained by the structural modifications which suggest the distortion of metaphosphate chains inducing the increase in the

*Variation of optical band gap energy of (0.9-x)NaPO3-xSiO2-0.1ZnO (0* ≤ *x* ≤ *0.1 mol) glass series.*

Because the NBOs bonds are predominantly ionic character and consequently have lower bond energies [34]. The higher value of the band gap energy revealed the increase of the cross-linking network due the introduction of SiO2 [4, 10, 11, 14, 16, 37, 41]. From **Figure 5**, it seems that the Eopt is in the order of 2.35 eV for 0.82 NaPO3-0.08 SiO2-0.1 ZnO and 0.8 NaPO3-0.1 SiO2-0.1 ZnO glass compositions. This result can be correlated to the structural changes due to the formation of P-O-Si ionic bands [11].

**3.1 Calorimetric dissolution of zinc, manganese and magnesium phosphate**

the dissolution of the glasses in a suitable solvent [11–13, 15, 29, 32, 33, 42].

their mixtures such as: HNO3, HCl, NaOH, KOH, CH3COOH.

tightly separated from the lower one by a movable cover.

The calorimetric study was performed by determining the energy resulting from

Phosphate glasses are soluble in mineral acids [11–13, 15, 29, 32, 33, 42]. Furthermore, the dissolution process has been carried out in order to find the suitable solvent which dissolves entirely the glassy compounds and should not give rise to

For this purpose, our investigations were covered all the usual acids, bases and

The calorimetric profile shows that the 4.5% weight of phosphoric acid solution

The reference and the measuring cell are surrounded by thermoelectric piles with high performance. These latters permit to detect the heat flow resulted from

is the best solvent for the thermochemical requirements of phosphate glasses. The dissolution of phosphate glasses were recorded by means the C80 (SETARAM) at 25°C. This equipment possesses two identical cells: the reference and the measuring cell. The reference cell should contain only the solvent but the measuring cell was provided with the solid to be dissolved or the liquid to be mixture. The superior compartment contains the attack solution (solvent) which is

number of non-bridging oxygen (NBOs).

**glasses**

**180**

**Figure 5.**

any secondary phenomena.

**3. Thermochemical study of phosphate glasses**

*Contemporary Topics about Phosphorus in Biology and Materials*

dehydration of its commercial monohydrate from NaH2PO4.H2O (Fluka of purity higher than 99%) at 150°C during 24 hours.

So the whole results allow to derive the enthalpy of reaction (R1) at 298.15 K. Taking into account the enthalpies of formation of NaH2PO4 (s) and H2O (liq) [42], we can derive that for sodium trimetaphosphate ((NaPO3)3, (crystal)) at 298.15 K.

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

For zinc-based phosphate glasses, the formation enthalpy of the glass series can be determined by considering a hypothetical reaction based on ((NaPO3)3(crystal)) and ZnO (sd). Their formation enthalpies can be derived by considering the following cycle which involves dissolution and mixing processes. The designed states from E5 to E8 are the solutions resulted from the different chemical operations [42]:

• E5 state presents the phenomena of dissolution for (NaPO3)3 in Slv1 (4.5%)

• E6 state designed the dissolution process of ZnO in Slv1 (4.5%) (ΔsolH3).

• E8 state designed the dissolution process of glasses in Slv1 (4.5%) (ΔsolH4).

According to this cycle ΔrH2 can be expressed as: ΔrH2 = (100-x) ΔsolH2 + 3x

� ð Þ <sup>100</sup> � <sup>x</sup> <sup>Δ</sup>fH NaPO ð Þ<sup>3</sup> <sup>3</sup>*;* crystal � 3x <sup>Δ</sup>fH ZnO ð Þ *;*sd (6)

<sup>þ</sup><sup>x</sup> <sup>Δ</sup>solH ZnO ð Þþ *;*sd <sup>Δ</sup>fH° ð Þ ZnO*;*sd � <sup>Δ</sup>solH glass ð Þþ <sup>1</sup>*=*3ΔmixH3

ð Þ NaPO3 <sup>3</sup>*,* crÞÞ

(7)

ΔsolH3 + ΔmixH3 + ΔmixH4 � 3ΔsolH4 when E7 and E8 states are identical

So, the standard enthalpy of formation of the glass can be derived as:

ΔrH2 ¼ 3 ΔfH 100 ½ � ð Þ � x NaPO3 � xZnO*;* glass

ð Þ¼ glass <sup>1</sup>*=*3 100 ð Þ � <sup>x</sup> <sup>Δ</sup>solH NaPO ð Þ<sup>3</sup> <sup>3</sup>*;* cristal <sup>þ</sup> <sup>Δ</sup>fH°

(ΔmixH4 ≈ 0) [42]. This quantity also equals:

ΔfH°

**183**

• E7 state designed the mixing process of E5 + E6 states (ΔmixH3).

**3.4 Cycle for Na2O-ZnO-P2O5 series glass**

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

(ΔsolH2).

NaH2PO4 was placed in the furnace, then the temperature increases from 200 to 300°C during 24 hours in order to eliminate residual water and volatile gases. After any heat treatment, the powder was crushed.

Then, the temperature increases and maintains 500°C for one night. The final product was tested using an X-ray diffraction equipped by SEIFERT-XRD 3000 TT diffractometer which confirms that the final product is the sodium trimetaphosphate.

#### **3.3 Cycle for sodium trimetaphosphate ((NaPO3)3, crystal)**

Generally, the direct determination of the formation enthalpy of any compounds is impossible. For this reason, our investigation is based on considering a particular reaction which involves the studied compounds with other reactants and products.

The knowledge of the enthalpy of the hypothetical reaction and the formation enthalpy of the reactants and products allow determining the formation enthalpy of the compound to be studied.

For the sodium trimetaphosphate ((NaPO3)3, crystal); the following thermochemical cycle has been studied [42]. This later put into consideration the chemical reaction which involves dissolution, dilution and mixing processes.

The designed states from E1 to E4 are the solutions obtained from different chemical operations [42]:


ΔrH1 can be expressed as ΔrH1 = 3 ΔsolH1 + ΔmixH2 ΔmixH1 ΔsolH2 ΔdilH1 in which ΔsolH are molar quantities.

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

So the whole results allow to derive the enthalpy of reaction (R1) at 298.15 K. Taking into account the enthalpies of formation of NaH2PO4 (s) and H2O (liq) [42], we can derive that for sodium trimetaphosphate ((NaPO3)3, (crystal)) at 298.15 K.

#### **3.4 Cycle for Na2O-ZnO-P2O5 series glass**

dehydration of its commercial monohydrate from NaH2PO4.H2O (Fluka of purity

NaH2PO4 was placed in the furnace, then the temperature increases from 200 to 300°C during 24 hours in order to eliminate residual water and volatile gases. After

Then, the temperature increases and maintains 500°C for one night. The final product was tested using an X-ray diffraction equipped by SEIFERT-XRD 3000 TT

Generally, the direct determination of the formation enthalpy of any compounds is impossible. For this reason, our investigation is based on considering a particular reaction which involves the studied compounds with other reactants and products. The knowledge of the enthalpy of the hypothetical reaction and the formation enthalpy of the reactants and products allow determining the formation enthalpy of

For the sodium trimetaphosphate ((NaPO3)3, crystal); the following thermochemical cycle has been studied [42]. This later put into consideration the chemical

The designed states from E1 to E4 are the solutions obtained from different

• E1 state presents the phenomena of dissolution for NaH2PO4 in Slv2 (4.45%)

• E2 state designed the dissolution process of (NaPO3)3 in Slv1 (4.5%) (ΔsolH2).

ΔrH1 can be expressed as ΔrH1 = 3 ΔsolH1 + ΔmixH2 ΔmixH1 ΔsolH2 ΔdilH1

diffractometer which confirms that the final product is the sodium

reaction which involves dissolution, dilution and mixing processes.

• E3 state designed the dilution process of 3H2O(liq) (ΔdiH1).

• E4 state designed the mixing process of E2 + E3 states (ΔmixH1).

**3.3 Cycle for sodium trimetaphosphate ((NaPO3)3, crystal)**

higher than 99%) at 150°C during 24 hours.

*Contemporary Topics about Phosphorus in Biology and Materials*

any heat treatment, the powder was crushed.

trimetaphosphate.

the compound to be studied.

chemical operations [42]:

in which ΔsolH are molar quantities.

**182**

(ΔsolH1).

For zinc-based phosphate glasses, the formation enthalpy of the glass series can be determined by considering a hypothetical reaction based on ((NaPO3)3(crystal)) and ZnO (sd). Their formation enthalpies can be derived by considering the following cycle which involves dissolution and mixing processes. The designed states from E5 to E8 are the solutions resulted from the different chemical operations [42]:


According to this cycle ΔrH2 can be expressed as: ΔrH2 = (100-x) ΔsolH2 + 3x ΔsolH3 + ΔmixH3 + ΔmixH4 � 3ΔsolH4 when E7 and E8 states are identical (ΔmixH4 ≈ 0) [42]. This quantity also equals:

$$\begin{aligned} \Delta\_{\text{f}} \mathbf{H}\_{2} &= 3 \,\Delta\_{\text{f}} \mathbf{H} \left[ (\mathbf{100} - \mathbf{x}) \text{NaPO}\_{3} - \mathbf{x} \text{ZnO}, \text{glass} \right] \\ &- (\mathbf{100} - \mathbf{x}) \Delta\_{\text{f}} \mathbf{H} \left( (\text{NaPO}\_{3})\_{3}, \text{crystal} \right) - 3 \mathbf{x} \,\Delta\_{\text{f}} \mathbf{H} \left( \text{ZnO}, \text{sd} \right) \end{aligned} \tag{6}$$

So, the standard enthalpy of formation of the glass can be derived as:

$$\begin{array}{c} \Delta\_{\text{f}}\text{H}^{\circ}(\text{glass}) = \text{1/3}(\text{100}-\text{x}) \left( \Delta\_{\text{sol}}\text{H}(\text{NaPO}\_{3})\_{3}, \text{critical} \right) + \Delta\_{\text{f}}\text{H}^{\circ}(\text{NaPO}\_{3})\_{3}, \text{cr}() \\ \quad + \text{x} \left( \Delta\_{\text{sol}}\text{H} \left( \text{ZnO,sd} \right) + \Delta\_{\text{f}}\text{H}^{\circ}(\text{ZnO,sd}) \right) - \Delta\_{\text{sol}}\text{H} \left( \text{glass} \right) + \text{1/3} \Delta\_{\text{mix}}\text{H}\_{3} \end{array} \tag{7}$$

#### *3.4.1 Dissolution processes*

**Tables 5**–**7** show the dissolution heat (Qr) of increasing the moles number (n) of NaH2PO4, (NaPO3)3 and ZnO solids in their corresponding solvents [42].

**Table 5** presents the dissolution process of NaH2PO4 with the variation of the quantity (n (mmol)) to be dissolved in 4.5 ml of phosphoric acid solution (4.45% (w/w) H3PO4) (Slv2).

**Table 6** presents the dissolution process of (NaPO3)3 with the variation of the quantity (n (mmol)) to be dissolved in 4.5 ml of phosphoric acid solution (4.5% (w/w) H3PO4) (Slv1).

**Table 7** presents the dissolution process for ZnO with the variation of the quantity (n (mmol)) to be dissolved in 4.5 ml of phosphoric acid solution (4.5% (w/w) H3PO4) (Slv1).

The plots of the variation of the measuring heats as a function of the moles number of solid leads to straight lines whose expressed as: Qr = An+b.

The slope (A) presents the molar dissolution enthalpy (ΔsolH) and b is the intercept increment.

Referring to a mathematical treatment developed in literature, the increment b is statistically not significant which leads to derive the dissolution enthalpy as [42]:


where (wi) is the reciprocal of the variance on ΔHi (wi = 1/σ<sup>2</sup>

for (NaH2PO4) (sd) in 4.45% weight of H3PO4 solution.

for ((NaPO3)3, (cristal)) in 4.5% weight of H3PO4 solution.

**Table 8** gathers the values of molar dissolution enthalpies with the

mass of solids (25 mg) in 4.5 ml of 4.5% weight of H3PO4 solution [42].

For zinc phosphate glasses, experiments were carried out by dissolving the same

**Compound NaH2PO4 (sd) in Slv2 (NaPO3)3 (sd) in Slv1 ZnO (sd) in Slv1** ΔsolH (kJ/mol) ΔsolH1 = 4.01 � 0.47 ΔsolH2 = 4.66 � 0.44 ΔsolH3 = �95.5 � 2.7

*Molar enthalpy of solution of dissolved compounds at the temperature T = 298.15 K and pressure p = 0.1 MPa*

For ZnO (sd) in 4.5% weight of H3PO4 solution.

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

Equations of the lines are as follows [42]:

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

phosphoric acid solution.

**ZnO (sd) in Slv1 (ΔsolH3)**

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

*Q r: Heat of solution. Slv1: 4.5% (w/w) H3PO4. <sup>a</sup>*

*confidence = 0.68).<sup>a</sup>*

**Table 7.**

corresponding errors.

*Slv1: 4.5% (w/w) H3PO4. Slv2: 4.45% (w/w) H3PO4. <sup>a</sup>*

*(level of confidence = 0.68).<sup>a</sup>*

**Table 8.**

**185**

the energy resulting by dissolving ni (mol) of the corresponding product in the

*Enthalpy of solution of ZnO (sd) in 4.5% (w/w) H3PO4 at T = 298.15 K and p = 0.1 MPa (level of*

**n (mmol) Q <sup>r</sup> (J) u(Q r) (J)** 0.2736 �26.081 0.12 0.3625 �34.600 0.33 0.2461 �23.500 0.24 0.2211 �21.215 0.22 0.3060 �29.124 0.50 0.3350 �32.010 0.21

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

ΔHi), and ΔHi is

ΔsolH1 ¼ 4*:*01 n (8)

ΔsolH2 ¼ 4*:*66 n (9)

ΔsolH3 ¼ �95*:*50 n*:* (10)

*Q r: Heat of solution.*

*Slv2: 4.45% (w/w) H3PO4. <sup>a</sup> Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

#### **Table 5.**

*Enthalpy of solution of NaH2PO4 (sd) in 4.45% (w/w) H3PO4 at the temperature T = 298.15 K and pressure p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*


*Q r: Heat of solution.*

*Slv1: 4.5% (w/w) H3PO4. <sup>a</sup>*

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

#### **Table 6.**

*Heat of solution of (NaPO3)3(cr) in 4.5% (w/w) H3PO4 at the temperature T = 298.15 K and p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*


*Q r: Heat of solution.*

*Slv1: 4.5% (w/w) H3PO4. <sup>a</sup>*

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

#### **Table 7.**

*3.4.1 Dissolution processes*

(w/w) H3PO4) (Slv2).

(4.5% (w/w) H3PO4) (Slv1).

(4.5% (w/w) H3PO4) (Slv1).

**NaH2PO4 (sd) in Slv2 (ΔsolH1)**

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

*p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*

**(NaPO3)3(cr) in Slv1 (ΔsolH2)**

intercept increment.

*Q r: Heat of solution. Slv2: 4.45% (w/w) H3PO4. <sup>a</sup>*

*Q r: Heat of solution. Slv1: 4.5% (w/w) H3PO4. <sup>a</sup>*

*(level of confidence = 0.68).<sup>a</sup>*

**Table 6.**

**184**

**Table 5.**

**Tables 5**–**7** show the dissolution heat (Qr) of increasing the moles number (n) of

**Table 5** presents the dissolution process of NaH2PO4 with the variation of the quantity (n (mmol)) to be dissolved in 4.5 ml of phosphoric acid solution (4.45%

**Table 6** presents the dissolution process of (NaPO3)3 with the variation of the quantity (n (mmol)) to be dissolved in 4.5 ml of phosphoric acid solution

**Table 7** presents the dissolution process for ZnO with the variation of the quantity (n (mmol)) to be dissolved in 4.5 ml of phosphoric acid solution

The plots of the variation of the measuring heats as a function of the moles

The slope (A) presents the molar dissolution enthalpy (ΔsolH) and b is the

Referring to a mathematical treatment developed in literature, the increment b is statistically not significant which leads to derive the dissolution enthalpy as [42]:

*Enthalpy of solution of NaH2PO4 (sd) in 4.45% (w/w) H3PO4 at the temperature T = 298.15 K and pressure*

*Heat of solution of (NaPO3)3(cr) in 4.5% (w/w) H3PO4 at the temperature T = 298.15 K and p = 0.1 MPa*

number of solid leads to straight lines whose expressed as: Qr = An+b.

**n (mmol) Q <sup>r</sup> (J) uQ <sup>r</sup> (J)** 0.2548 1.084 0.05 0.3700 1.328 0.08 0.3440 1.282 0.06 0.3000 1.167 0.11 0.1366 0.680 0.06

**n (mmol) Q <sup>r</sup> (J) uQ <sup>r</sup> (J)** 0.1995 0.950 0.08 0.2473 1.171 0.06 0.2990 1.370 0.07 0.3452 1.590 0.12 0.2626 1.240 0.10

NaH2PO4, (NaPO3)3 and ZnO solids in their corresponding solvents [42].

*Contemporary Topics about Phosphorus in Biology and Materials*

*Enthalpy of solution of ZnO (sd) in 4.5% (w/w) H3PO4 at T = 298.15 K and p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*

$$\Delta \text{sol}\,\text{H} = \begin{array}{c} \Sigma(\text{w}\_{i}\text{\*}\Delta\text{H}\text{in}\_{i})\\ \Sigma(\text{w}\_{i}\text{\*}\text{n}\_{i}^{2}) \end{array} = \text{A} \begin{array}{c} \text{(kJmol}^{-1}\text{)} \end{array}$$

where (wi) is the reciprocal of the variance on ΔHi (wi = 1/σ<sup>2</sup> ΔHi), and ΔHi is the energy resulting by dissolving ni (mol) of the corresponding product in the phosphoric acid solution.

Equations of the lines are as follows [42]:

$$
\Delta\_{\text{sol}}\text{H}\_1 = 4.01\,\text{n}\tag{8}
$$

for (NaH2PO4) (sd) in 4.45% weight of H3PO4 solution.

$$
\Delta\_{\text{sol}}\text{H}\_2 = 4.66\,\text{n}\tag{9}
$$

for ((NaPO3)3, (cristal)) in 4.5% weight of H3PO4 solution. For ZnO (sd) in 4.5% weight of H3PO4 solution.

$$
\Delta\_{\text{sol}}\text{H}\_{\text{3}} = -\text{95.50 n.}\tag{10}
$$

**Table 8** gathers the values of molar dissolution enthalpies with the corresponding errors.

For zinc phosphate glasses, experiments were carried out by dissolving the same mass of solids (25 mg) in 4.5 ml of 4.5% weight of H3PO4 solution [42].


**Table 8.**

*Molar enthalpy of solution of dissolved compounds at the temperature T = 298.15 K and pressure p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*

#### *Contemporary Topics about Phosphorus in Biology and Materials*


#### **Table 9.**

*Evolution of heat solution of (100-x)NaPO3-xZnO phosphate glasses in 4.5% (w/w) H3PO4 at the temperature T = 298.15 K and pressure p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*

**Table 9** reports the variation of the heat dissolution for the glass series published in a previous work [33]. This evolution shows that the dissolution heat decreases linearly with the incorporation of ZnO oxide. As a result the inversion in the thermal signs for the studies glasses. This variation can be explained by the structural changes of the phosphate network suggesting the distortion of metaphosphate chains revealed by 31P MAS-NMR analysis.

Plotting of dissolution heat versus ZnO proportion (ΔsolH (kJ mol<sup>1</sup> )) is reported in **Figure 6**. It seems that the calorimetric dissolution of the glass series is endothermic for lower ZnO proportion and becomes exothermic above 18 mol% of ZnO. This variation can be correlated to the cleavage of P▬O▬P bridges which suggests the appearance of pyrophosphate groups (Q<sup>1</sup> ), revealed by 31P MAS-NMR spectroscopic analysis, when ZnO oxide is progressively incorporated in the vitreous network [33].

#### *3.4.2 Mixing processes*

#### *3.4.2.1 For ((NaPO3)3, crystal) cycle*

E2 was provided with various amounts of (NaPO3)3 (6–9 mg) which have been dissolved in 4.5% (w/w) H3PO4 solution but E3 is 4.4% (w/w) H3PO4 solution. Mixing the same volumes of E2 and E3 (around 2 ml) leads to a solution having the mean value of acid composition (E4 with 4.45% (w/w) H3PO4 or [H3PO4.116.90H2O]. Consequently, the concentration of Slv2 was fixed as 4.45% (w/w) H3PO4 in order to get identical E1 and E4 states. **Table 10** reports E2 + E3 mixing enthalpy for different mole number (n) of (NaPO3)3 added in E2. This allowed to express ΔmixH1 as: ΔmixH1 = 0.07 n (R2 = 0.995) leading to a value of 0.07 kJ per (NaPO3)3 mole [42].

the various amounts of ZnO so the variation of energy (**Table 11**) is due to the presence of ZnO in the solution. The mixture of E7 and E8 states has no detectable

*Enthalpy of mixing: ΔmixH3 and ΔmixH4 at T = 298.15 K and p = 0.1 MPa (level of confidence = 0.68).a*

**n (mmol) Qr (J) u(Qr) (J)** 0.0187 0.0005 0 0.0257 0.0010 0 0.0157 0.0003 0 0.0300 0.0013 0

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

*Enthalpy of mixing: ΔmixH1 (E2 + E3) at the temperature T = 298.15 K and pressure p = 0.1 MPa (level of*

**u(ΔmixH3) (kJ/mol)**

**ΔmixH4 (E7 + E8) (kJ/mol)**

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa. E2: Solution with various amounts of (NaPO3)3 in 4.5% (w/w) H3PO4.*

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

**Composition ΔmixH3 (E5 + E6)**

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa. E5: solution with 27.3 mg of (NaPO3)3 in 4.5% (w/w) H3PO4. E6: solution with various amounts of ZnO in 4.5% (w/w) H3PO4. E8: solution with 25 mg of glass composition in 4.5% (w/w) H3PO4.*

**(kJ/ZnO mole)**

95 NaPO3-5 ZnO 10.15 0.10 ≈0 90 NaPO3-10 ZnO 4.00 0.13 ≈0 85 NaPO3-15 ZnO 5.10 0.21 ≈0 80 NaPO3-20 ZnO 3.23 0.24 ≈0 75 NaPO3-25 ZnO 2.74 0.25 ≈0 70 NaPO3-30 ZnO 3.00 0.15 ≈0 67 NaPO3-33 ZnO 2.12 0.20 ≈0

For the cycle corresponding to the sodium trimetaphosphate, addition of water to Slv1 [H3PO4.115.54 H2O] corresponds to a dilution process. The corresponding energy was calculated by linear interpolation of literature data considering the interval to which belongs each enthalpies of solution of the initial (4.5% (w/w)

The formation and dilution enthalpies were calculated from Ref. [42] and listed

The formation enthalpy of sodium trimetaphosphate ((NaPO3)3, crystal) has

ð½ � H3PO4*:*115*:*54H2O ޼�1288*:*255 Hð Þ 3PO4 4*:*5% (11)

ð½ � H3PO4*:*118*:*54H2O ޼�1288*:*272 Hð Þ 3PO4 4*:*4% (12)

. The obtained value differs from the

thermal effect [42].

*3.4.3 Dilution processes*

below:

**187**

*a*

*a*

**Table 11.**

**Table 10.**

*E3: 4.4% (w/w) H3PO4. Qr: Heat of mixing.*

*confidence = 0.68).<sup>a</sup>*

H3PO4) and final states (4.4% (w/w) H3PO4) [42].

Calculation gives <sup>Δ</sup>dilH1 <sup>=</sup> �0.017 kJ mol�<sup>1</sup> H3PO4.

ΔfH°

ΔfH°

been deduced as (�3762.5 � 175) kJ mol�<sup>1</sup>

(E1 + E4) mixing process, which noted as (ΔmixH2), was considered in order to check whether or not they correspond to the same final state. This operation led to an undetectable thermal effect [42].

#### *3.4.2.2 For Na2O-ZnO-P2O5 glasses cycle*

Mixing the same volumes (around 2 ml) of E5 and E6 which have the same concentration of phosphoric acid (4.5% (w/w) H3PO4), led to E7 solution. Were previously added to while The E5 sate was obtained by dissolving the average mass of ((NaPO3)3, (cristal)) (27.3 mg) whereas the E6 state was considered by dissolving *Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*


*a Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

*E2: Solution with various amounts of (NaPO3)3 in 4.5% (w/w) H3PO4.*

*E3: 4.4% (w/w) H3PO4.*

*Qr: Heat of mixing.*

#### **Table 10.**

**Table 9** reports the variation of the heat dissolution for the glass series published in a previous work [33]. This evolution shows that the dissolution heat decreases linearly with the incorporation of ZnO oxide. As a result the inversion in the thermal signs for the studies glasses. This variation can be explained by the structural changes of the phosphate network suggesting the distortion of metaphosphate

**Composition ΔsolH (kJ/mol) u (ΔsolH) (kJ/mol)** NaPO3 4.80 0.45 95 NaPO3-5 ZnO 3.70 0.20 90 NaPO3-10 ZnO 2.92 0.15 85 NaPO3-15 ZnO 1.70 0.10 80 NaPO3-20 ZnO 1.15 0.10 75 NaPO3-25 ZnO 3.21 0.20 70 NaPO3-30 ZnO 6.34 0.32 67 NaPO3-33 ZnO 13.05 1.00

)) is

), revealed by 31P MAS-NMR

Plotting of dissolution heat versus ZnO proportion (ΔsolH (kJ mol<sup>1</sup>

*Evolution of heat solution of (100-x)NaPO3-xZnO phosphate glasses in 4.5% (w/w) H3PO4 at the*

*temperature T = 298.15 K and pressure p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*

reported in **Figure 6**. It seems that the calorimetric dissolution of the glass series is endothermic for lower ZnO proportion and becomes exothermic above 18 mol% of ZnO. This variation can be correlated to the cleavage of P▬O▬P bridges which

spectroscopic analysis, when ZnO oxide is progressively incorporated in the vitre-

E2 was provided with various amounts of (NaPO3)3 (6–9 mg) which have been dissolved in 4.5% (w/w) H3PO4 solution but E3 is 4.4% (w/w) H3PO4 solution. Mixing the same volumes of E2 and E3 (around 2 ml) leads to a solution having the

[H3PO4.116.90H2O]. Consequently, the concentration of Slv2 was fixed as 4.45% (w/w) H3PO4 in order to get identical E1 and E4 states. **Table 10** reports E2 + E3 mixing enthalpy for different mole number (n) of (NaPO3)3 added in E2. This allowed to express ΔmixH1 as: ΔmixH1 = 0.07 n (R2 = 0.995) leading to a value of

(E1 + E4) mixing process, which noted as (ΔmixH2), was considered in order to check whether or not they correspond to the same final state. This operation led to

Mixing the same volumes (around 2 ml) of E5 and E6 which have the same concentration of phosphoric acid (4.5% (w/w) H3PO4), led to E7 solution. Were previously added to while The E5 sate was obtained by dissolving the average mass of ((NaPO3)3, (cristal)) (27.3 mg) whereas the E6 state was considered by dissolving

chains revealed by 31P MAS-NMR analysis.

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

*Contemporary Topics about Phosphorus in Biology and Materials*

ous network [33].

*a*

**Table 9.**

*3.4.2 Mixing processes*

*3.4.2.1 For ((NaPO3)3, crystal) cycle*

0.07 kJ per (NaPO3)3 mole [42].

**186**

an undetectable thermal effect [42].

*3.4.2.2 For Na2O-ZnO-P2O5 glasses cycle*

suggests the appearance of pyrophosphate groups (Q<sup>1</sup>

mean value of acid composition (E4 with 4.45% (w/w) H3PO4 or

*Enthalpy of mixing: ΔmixH1 (E2 + E3) at the temperature T = 298.15 K and pressure p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*


*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

*E5: solution with 27.3 mg of (NaPO3)3 in 4.5% (w/w) H3PO4.*

*E6: solution with various amounts of ZnO in 4.5% (w/w) H3PO4.*

*E8: solution with 25 mg of glass composition in 4.5% (w/w) H3PO4.*

#### **Table 11.**

*Enthalpy of mixing: ΔmixH3 and ΔmixH4 at T = 298.15 K and p = 0.1 MPa (level of confidence = 0.68).a*

the various amounts of ZnO so the variation of energy (**Table 11**) is due to the presence of ZnO in the solution. The mixture of E7 and E8 states has no detectable thermal effect [42].

#### *3.4.3 Dilution processes*

For the cycle corresponding to the sodium trimetaphosphate, addition of water to Slv1 [H3PO4.115.54 H2O] corresponds to a dilution process. The corresponding energy was calculated by linear interpolation of literature data considering the interval to which belongs each enthalpies of solution of the initial (4.5% (w/w) H3PO4) and final states (4.4% (w/w) H3PO4) [42].

The formation and dilution enthalpies were calculated from Ref. [42] and listed below:

> ΔfH° ð½ � H3PO4*:*115*:*54H2O ޼�1288*:*255 Hð Þ 3PO4 4*:*5% (11)

$$
\Delta\_{\rm f} \text{H}^{\circ} \text{(} [\text{H}\_{3}\text{PO}\_{4}.\text{118.54}\text{H}\_{2}\text{O}] \text{)} = -1288.272 \text{ (} \text{H}\_{3}\text{PO}\_{4}.\text{4.496}\text{)}\tag{12}
$$

Calculation gives <sup>Δ</sup>dilH1 <sup>=</sup> �0.017 kJ mol�<sup>1</sup> H3PO4.

The formation enthalpy of sodium trimetaphosphate ((NaPO3)3, crystal) has been deduced as (�3762.5 � 175) kJ mol�<sup>1</sup> . The obtained value differs from the

#### *Contemporary Topics about Phosphorus in Biology and Materials*


Amorphous state was investigated by means of FTIR, Raman, MAS-NMR and

)

) when

Spectroscopic analysis revealed the formation of pyrophosphate groups (Q<sup>1</sup>

Furthermore, the indirect optical band gap energy for zinc phosphate-based silicate glasses increases with the addition of SiO2 oxide. This suggests the increase in the NBOs resulting from the modification of P▬O▬P bridges which revealed the

On the other hand, the dissolution process is endothermic at lower MO content and become exothermic when MO oxide is progressively incorporated. The change in thermal sign could be correlated to the structural modification inducing the formation of P▬O▬M ionic bond which increases the rigidity and the compacity of

The glass formation enthalpy increases when ZnO oxide is progressively incor-

Furthermore, the variation of Tg values reflects an increase of the rigidity of the

Because of the large disorder that exists in the vitreous structure, the entropy factor (ΔrS°) should prevail and induce the decrease in ΔrG° value when ZnO

, Ismail Khattech<sup>2</sup>

When replacing of P by Zn induced a decrease in the binding energy which

glass network due to the formation of P▬O▬Zn ionic bonds. As a result, the increase in the stability of the phosphate network which is tightly related to the

UV-visible spectroscopy in order to study the structural role of MO oxide.

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

resulting from the depolymerization of infinite metaphosphate groups (Q2

Furthermore, the glass formation enthalpy of (100-x)NaPO3-xZnO (0 ≤ x ≤ 33 mol%) glass series were determine by considering a thermochemical cycle involving the formation enthalpy of sodium trimetaphosphate ((NaPO3)3,

suggest the increase of the formation enthalpy of the glass series.

\*, Mohamed Jemal<sup>2</sup>

\*Address all correspondence to: refkaoueslati@gmail.com

provided the original work is properly cited.

1 Useful Materials Valorization Laboratory, National Center For Research in Materials Sciences, Technological Park of Borj Cedria, Soliman, Tunisia

2 Université de Tunis El Manar, Faculty of Science, Chemistry Department, Laboratory of Materials Crystal and Applied Thermodynamics LR15SE01, Tunis,

© 2019 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium,

the modifying oxide is gradually incorporated.

shortening of the metaphosphate chains.

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

crystal). This later was checked in this work.

porated in the vitreous network.

Gibbs free energy of formation ΔrG°.

concentration increases.

**Author details**

Tunisia

**189**

Refka Oueslati Omrani<sup>1</sup>

and Ahmed Hichem Hamzaoui<sup>1</sup>

the vitreous network.

#### **Table 12.**

*Formation enthalpy of: (100-x)NaPO3-xZnO glasses at the temperature T = 298.15 K and pressure p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*

**Figure 7.**

*Evolution of the standard enthalpy of formation at standard temperature and pressure of (100-x)NaPO3 xZnO (0* ≤ *x* ≤ *33 mol%) glasses [45].*

older by only 2.4%. It seems that the calculated value of the formation enthalpy of sodium trimetaphosphate ((NaPO3)3, crystal) is in good agreement with this determined previously in 1968. This confirms that the synthesized product is probably the sodium trimetaphosphate and not a mixture.

The variation of the formation enthalpy of the glass series are mentioned in **Table 12**. This latter shows that this quantity increases with the addition of ZnO oxide as reported in **Figure 7** [42].

#### **4. Conclusions**

The influence of ZnO, MgO, MnO and SiO2 addition on the structure, physical and optical properties of phosphate glasses and phosphate-based silicate glasses having a general formula: (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mn, Mg) where 3 ≤ O/P ≤ 3.49; (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) with O/P = 3 (0 ≤ x ≤ 33 mol%) and (0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol).

#### *Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

Amorphous state was investigated by means of FTIR, Raman, MAS-NMR and UV-visible spectroscopy in order to study the structural role of MO oxide.

Spectroscopic analysis revealed the formation of pyrophosphate groups (Q<sup>1</sup> ) resulting from the depolymerization of infinite metaphosphate groups (Q2 ) when the modifying oxide is gradually incorporated.

Furthermore, the indirect optical band gap energy for zinc phosphate-based silicate glasses increases with the addition of SiO2 oxide. This suggests the increase in the NBOs resulting from the modification of P▬O▬P bridges which revealed the shortening of the metaphosphate chains.

On the other hand, the dissolution process is endothermic at lower MO content and become exothermic when MO oxide is progressively incorporated. The change in thermal sign could be correlated to the structural modification inducing the formation of P▬O▬M ionic bond which increases the rigidity and the compacity of the vitreous network.

Furthermore, the glass formation enthalpy of (100-x)NaPO3-xZnO (0 ≤ x ≤ 33 mol%) glass series were determine by considering a thermochemical cycle involving the formation enthalpy of sodium trimetaphosphate ((NaPO3)3, crystal). This later was checked in this work.

The glass formation enthalpy increases when ZnO oxide is progressively incorporated in the vitreous network.

When replacing of P by Zn induced a decrease in the binding energy which suggest the increase of the formation enthalpy of the glass series.

Furthermore, the variation of Tg values reflects an increase of the rigidity of the glass network due to the formation of P▬O▬Zn ionic bonds. As a result, the increase in the stability of the phosphate network which is tightly related to the Gibbs free energy of formation ΔrG°.

Because of the large disorder that exists in the vitreous structure, the entropy factor (ΔrS°) should prevail and induce the decrease in ΔrG° value when ZnO concentration increases.

#### **Author details**

older by only 2.4%. It seems that the calculated value of the formation enthalpy of sodium trimetaphosphate ((NaPO3)3, crystal) is in good agreement with this determined previously in 1968. This confirms that the synthesized product is probably

*Evolution of the standard enthalpy of formation at standard temperature and pressure of (100-x)NaPO3-*

The variation of the formation enthalpy of the glass series are mentioned in **Table 12**. This latter shows that this quantity increases with the addition of ZnO

The influence of ZnO, MgO, MnO and SiO2 addition on the structure, physical and optical properties of phosphate glasses and phosphate-based silicate glasses having a general formula: (50-x/2)Na2O-xMO-(50-x/2)P2O5 (M = Zn, Mn, Mg) where 3 ≤ O/P ≤ 3.49; (50-x)Na2O-xMO-50P2O5 (M = Zn, Mn) with O/P = 3 (0 ≤ x ≤ 33 mol%) and (0.9-x)NaPO3-xSiO2-0.1ZnO (0 ≤ x ≤ 0.1 mol).

the sodium trimetaphosphate and not a mixture.

**Composition ΔfH°**

*Contemporary Topics about Phosphorus in Biology and Materials*

*Standard uncertainties u are u(T) = 0.01 K, u(p) = 10 kPa.*

*p = 0.1 MPa (level of confidence = 0.68).<sup>a</sup>*

*a*

**Table 12.**

**Figure 7.**

NaPO3 1260 50 95 NaPO3-5 ZnO 1213 49 90 NaPO3-10 ZnO 1174 47 85 NaPO3-15 ZnO 1132 45 80 NaPO3-20 ZnO 1090 44 75 NaPO3-25 ZnO 1070 43 70 NaPO3-30 ZnO 1003 40 67 NaPO3-33 ZnO 973 39

*Formation enthalpy of: (100-x)NaPO3-xZnO glasses at the temperature T = 298.15 K and pressure*

**(kJ/mol) u(ΔfH°**

**) (kJ/mol)**

oxide as reported in **Figure 7** [42].

*xZnO (0* ≤ *x* ≤ *33 mol%) glasses [45].*

**4. Conclusions**

**188**

Refka Oueslati Omrani<sup>1</sup> \*, Mohamed Jemal<sup>2</sup> , Ismail Khattech<sup>2</sup> and Ahmed Hichem Hamzaoui<sup>1</sup>

1 Useful Materials Valorization Laboratory, National Center For Research in Materials Sciences, Technological Park of Borj Cedria, Soliman, Tunisia

2 Université de Tunis El Manar, Faculty of Science, Chemistry Department, Laboratory of Materials Crystal and Applied Thermodynamics LR15SE01, Tunis, Tunisia

\*Address all correspondence to: refkaoueslati@gmail.com

© 2019 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

### **References**

[1] Sreedhar VB, Basavapoornima CH, Jayasankar CK. Spectroscopic and fluorescence properties properties of Sm3+ doped zincfluorophosphate glasses. Journal of Rare Earths. 2014; **32**:918

[2] Aguiar H, Solla EL, Serra J, Conzalez P, Leon B, Malz F, et al. Jager, Raman and NMR study of bioactive Na2O-MgO-CaO-P2O5-SiO2 glasses. Journal of Non-Crystalline Solids. 2008; **354**:5004-5008

[3] Zid MHM, Matori KA, Hj S, Aziz A, Zakaria A. Effect of ZnO on the physical and optical band gap of soda lime silicate glass. International Journal of Molecular Sciences. 2012;**13**: 7550-7558

[4] Ibrahim AM, Bader AM, Elshaikh HA, Mostafa AG, Elbashar YH. Effect of CuO addition on the dielectric parameters of sodium zinc phosphate glasses. Silicon. 2017. pp. 1265-1274

[5] Vedeanu N, Stanescu R, Filip S, Ardelean I, Cozar O. IR and ESR investigations on V2O5-P2O5-BaO glass system with opto-electronic potential. Journal of Non-Crystalline Solids. 2012; **358**:1881-1885

[6] Ahmina W, El Moudane M, Zriouil M, Taibi M. Effect of the content of MnO on the electric-dielectric properties of potassium-phosphate glasses. Journal of Materials and Environmental Science. 2017;**11**: 4193-4198

[7] Waclawska I, Szumera M, Sulowska J. Structural characterization of zinc-modified glasses from the SiO2- K2O-CaO-MgO. Journal of Alloys and Compounds. 2016;**666**:352-358

[8] Pascuta P, Bosca M, Borodi G, Vida-Simiti I, Culea E. Thermal, structural and magnetic properties of some zinc

phosphate glasses doped with manganese ions. Journal of Alloys and Compounds. 2011;**509**:4314-4319

[16] Magdas DA, Stefan R, Toloman D, Vedeanu NS. Copper ions influence on lead-phosphate glass network. Journal of Molecular Structure. 2014;**1056-1057**:

*DOI: http://dx.doi.org/10.5772/intechopen.88539*

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate…*

[25] Pascuta P, Borodi G, Jumate N, Vida-Simiti I, Viorel D, Culea E. The structural role of manganese ions in some zinc phosphate glasses and glass ceramics. Journal of Alloys and Compounds. 2010;**504**:479-483

[26] Pop L, Bolundut L, Pascuta P, Culea E. Influence of Er3+ ions addition on thermal and optical properties of phosphate-germanate system. Journal of Thermal Analysis and Calorimetry. 2019

[27] Jlassi I, Elhouichet H, Ferid M. Electrical conductivity and dielectric properties of MgO doped lithium phosphate glasses. Journal of Physics E.

[28] Ciceo-Lucacel R, Todea M, Simon V. Effect of selenium addition on network connectivity in P2O5-CaO-MgO-Na2O glasses. Journal of Non-Crystalline

Montagne L, Reval B, Jemal M. Effect of SrO content on the structure and properties of sodium-strontium metaphosphate glasses. The Journal of Physical Chemistry A. 2017:62-68

[30] Sevastiajova I, Aseev V, Tuzova L, Fedorov Y, Nikonorov N. Spectral and luminescence properties of manganese ions in vitreous lead metaphosphate. Journal of Luminescence. 2019:495-499

[31] Ciceo-Lucacel R, Todea M, Simon V. Effect of selenium addition on network connectivity in P2O5-CaO-MgO-Na2O glasses. Journal of Non-Crystalline

2016;**81**:219-223

Solids. 2018;**480**:10-13

Solids. 2018;**488**:10-13

Solids. 2014;**389**:66-71

[32] Oueslati-Omrani R, Krimi S, Videau JJ, Khattech I, El Jazouli A, Jemal M. Structural investigations and calorimetric dissolution of manganese phosphate glasses. Journal of Crystalline

[33] Oueslati-Omrani R, Krimi S, Videau JJ, Khattech I, El Jazouli A,

[29] Cherbib MA, Khattech I,

[17] Vedeanu N, Magdas DA, Stefan R. Structural modifications induced by addition of copper oxide to leadphosphate glasses. Journal of Non-Crystalline Solids. 2012;**358**:

[18] Zhang L, Liu S. Structure and crystallization behavior of 50CuOxTiO2-(50-x)P2O5. Journal of Non-Crystalline Solids. 2017;**473**:108-113

[19] Brow RK. Review: The structure simple of phosphate glasses. Journal of Non-Crystalline Solids. 2000;**263-264**:

[20] Aguiar H, Solla JM, Serra J, Gonzalez P, Leon B, Almeida N, et al. Orthophosphate nanostructures in SiO2- P2O5-CaO-Na2O-MgO biactive glasses. Journal of Non-Crystalline Solids. 2008;

[21] Abd el Ghany HA. Physicl and optical characterization of manganese ions in sodium-zinc-phosphate glass

matrix. IARJSET. 2018. p. 5

[22] Khor SF, Talib ZA, Malek F, Cheng EM. Optical properties of

[23] Walter G, Vogel J, Hoppe U, Hartmann P. The structure of CaO-Na2O-MgO-P2O5 invert glass. Journal of Non-Crystalline Solids. 2001;**296**:

ultraphosphate glasses containing mixed divalent zinc and magnesium ions. Optical Materials. 2013;**35**:629-633

[24] Li HC, Wang DG, Hu JH, Chen CZ. Influence of fluoride additions on biological and mechanical properties of Na2O-CaO-SiO2-P2O5 glass-ceramics. Materials Letters. 2013;**106**:373-376

314-318

3170-3174

1-28

**354**:4075-4080

212-223

**191**

[9] Oui MA, Azooz MA, Elbatal HA. Optical and infrared spectral investigations of cadmium zinc phosphate glasses doped with WO3 or MoO3 before and after subjecting to gamma irradiation. Journal of Non-Crystalline Solids. 2018;**494**:31-39

[10] Khor SF, Talib A, Mat Yunus WM. Optical properties of ternary zinc magnesium phosphate glasses. Ceramics International. 2012;**38**:935-940

[11] Omrani RO, Hamzaoui AH, Chtourou R, M'nif A. Structural, thermal and optical properties of phosphate glasses doped with SiO2. Journal of Non-Crystalline Solids. 2018; **481**:10-16

[12] Cherbib MA, Khattech I, Montagne L, Reval B, Jemal M. Structure properties relationship in calcium sodium metaphosphate and polyphosphate glasses. Journal of Non-Crystalline Solids. 2018;**485**:1-13

[13] Oueslati-Omrani R, Kaoutar A, El Jazouli A, Krimi S, Khattech I, Jemal M, et al. Structural and thermochemical properties of sodium magnesium phosphate glasses. Journal of Alloys and Compounds. 2015;**632**:766-771

[14] El-Maaref AA, Badr S, ElOkr KS, Abdel Wahab EA, ElOkr MM. Optical properties and radiatives rates of Nd3+ doped zinc-sodium phosphate glasses. Journal of Rare Earths. 2019;**37**:253-259

[15] Cherbib MA, Krimi S, El Jazouli A, Khattech I, Montagne L, Reval B, et al. Structure and thermochemical study of strontium sodium phosphate glasses. Journal of Non-Crystalline Solids. 2016; **447**:59-65

*Structural and Calorimetric Studies of Zinc, Magnesium and Manganese Based Phosphate… DOI: http://dx.doi.org/10.5772/intechopen.88539*

[16] Magdas DA, Stefan R, Toloman D, Vedeanu NS. Copper ions influence on lead-phosphate glass network. Journal of Molecular Structure. 2014;**1056-1057**: 314-318

**References**

**32**:918

**354**:5004-5008

7550-7558

**358**:1881-1885

4193-4198

**190**

[1] Sreedhar VB, Basavapoornima CH, Jayasankar CK. Spectroscopic and fluorescence properties properties of Sm3+ doped zincfluorophosphate glasses. Journal of Rare Earths. 2014;

*Contemporary Topics about Phosphorus in Biology and Materials*

phosphate glasses doped with

manganese ions. Journal of Alloys and Compounds. 2011;**509**:4314-4319

[9] Oui MA, Azooz MA, Elbatal HA. Optical and infrared spectral investigations of cadmium zinc phosphate glasses doped with WO3 or MoO3 before and after subjecting to gamma irradiation. Journal of Non-Crystalline Solids. 2018;**494**:31-39

[10] Khor SF, Talib A, Mat Yunus WM. Optical properties of ternary zinc magnesium phosphate glasses. Ceramics

International. 2012;**38**:935-940

[11] Omrani RO, Hamzaoui AH, Chtourou R, M'nif A. Structural, thermal and optical properties of phosphate glasses doped with SiO2. Journal of Non-Crystalline Solids. 2018;

[12] Cherbib MA, Khattech I, Montagne L, Reval B, Jemal M. Structure properties relationship in calcium sodium metaphosphate and polyphosphate glasses. Journal of Non-Crystalline Solids. 2018;**485**:1-13

[13] Oueslati-Omrani R, Kaoutar A, El Jazouli A, Krimi S, Khattech I, Jemal M, et al. Structural and thermochemical properties of sodium magnesium phosphate glasses. Journal of Alloys and

[14] El-Maaref AA, Badr S, ElOkr KS, Abdel Wahab EA, ElOkr MM. Optical properties and radiatives rates of Nd3+ doped zinc-sodium phosphate glasses. Journal of Rare Earths. 2019;**37**:253-259

[15] Cherbib MA, Krimi S, El Jazouli A, Khattech I, Montagne L, Reval B, et al. Structure and thermochemical study of strontium sodium phosphate glasses. Journal of Non-Crystalline Solids. 2016;

Compounds. 2015;**632**:766-771

**481**:10-16

**447**:59-65

Conzalez P, Leon B, Malz F, et al. Jager, Raman and NMR study of bioactive Na2O-MgO-CaO-P2O5-SiO2 glasses. Journal of Non-Crystalline Solids. 2008;

Aziz A, Zakaria A. Effect of ZnO on the physical and optical band gap of soda lime silicate glass. International Journal

Elshaikh HA, Mostafa AG, Elbashar YH. Effect of CuO addition on the dielectric parameters of sodium zinc phosphate glasses. Silicon. 2017. pp. 1265-1274

[5] Vedeanu N, Stanescu R, Filip S, Ardelean I, Cozar O. IR and ESR investigations on V2O5-P2O5-BaO glass system with opto-electronic potential. Journal of Non-Crystalline Solids. 2012;

[6] Ahmina W, El Moudane M,

of MnO on the electric-dielectric properties of potassium-phosphate glasses. Journal of Materials and Environmental Science. 2017;**11**:

[7] Waclawska I, Szumera M,

Zriouil M, Taibi M. Effect of the content

Sulowska J. Structural characterization of zinc-modified glasses from the SiO2- K2O-CaO-MgO. Journal of Alloys and Compounds. 2016;**666**:352-358

[8] Pascuta P, Bosca M, Borodi G, Vida-Simiti I, Culea E. Thermal, structural and magnetic properties of some zinc

[2] Aguiar H, Solla EL, Serra J,

[3] Zid MHM, Matori KA, Hj S,

of Molecular Sciences. 2012;**13**:

[4] Ibrahim AM, Bader AM,

[17] Vedeanu N, Magdas DA, Stefan R. Structural modifications induced by addition of copper oxide to leadphosphate glasses. Journal of Non-Crystalline Solids. 2012;**358**: 3170-3174

[18] Zhang L, Liu S. Structure and crystallization behavior of 50CuOxTiO2-(50-x)P2O5. Journal of Non-Crystalline Solids. 2017;**473**:108-113

[19] Brow RK. Review: The structure simple of phosphate glasses. Journal of Non-Crystalline Solids. 2000;**263-264**: 1-28

[20] Aguiar H, Solla JM, Serra J, Gonzalez P, Leon B, Almeida N, et al. Orthophosphate nanostructures in SiO2- P2O5-CaO-Na2O-MgO biactive glasses. Journal of Non-Crystalline Solids. 2008; **354**:4075-4080

[21] Abd el Ghany HA. Physicl and optical characterization of manganese ions in sodium-zinc-phosphate glass matrix. IARJSET. 2018. p. 5

[22] Khor SF, Talib ZA, Malek F, Cheng EM. Optical properties of ultraphosphate glasses containing mixed divalent zinc and magnesium ions. Optical Materials. 2013;**35**:629-633

[23] Walter G, Vogel J, Hoppe U, Hartmann P. The structure of CaO-Na2O-MgO-P2O5 invert glass. Journal of Non-Crystalline Solids. 2001;**296**: 212-223

[24] Li HC, Wang DG, Hu JH, Chen CZ. Influence of fluoride additions on biological and mechanical properties of Na2O-CaO-SiO2-P2O5 glass-ceramics. Materials Letters. 2013;**106**:373-376

[25] Pascuta P, Borodi G, Jumate N, Vida-Simiti I, Viorel D, Culea E. The structural role of manganese ions in some zinc phosphate glasses and glass ceramics. Journal of Alloys and Compounds. 2010;**504**:479-483

[26] Pop L, Bolundut L, Pascuta P, Culea E. Influence of Er3+ ions addition on thermal and optical properties of phosphate-germanate system. Journal of Thermal Analysis and Calorimetry. 2019

[27] Jlassi I, Elhouichet H, Ferid M. Electrical conductivity and dielectric properties of MgO doped lithium phosphate glasses. Journal of Physics E. 2016;**81**:219-223

[28] Ciceo-Lucacel R, Todea M, Simon V. Effect of selenium addition on network connectivity in P2O5-CaO-MgO-Na2O glasses. Journal of Non-Crystalline Solids. 2018;**480**:10-13

[29] Cherbib MA, Khattech I, Montagne L, Reval B, Jemal M. Effect of SrO content on the structure and properties of sodium-strontium metaphosphate glasses. The Journal of Physical Chemistry A. 2017:62-68

[30] Sevastiajova I, Aseev V, Tuzova L, Fedorov Y, Nikonorov N. Spectral and luminescence properties of manganese ions in vitreous lead metaphosphate. Journal of Luminescence. 2019:495-499

[31] Ciceo-Lucacel R, Todea M, Simon V. Effect of selenium addition on network connectivity in P2O5-CaO-MgO-Na2O glasses. Journal of Non-Crystalline Solids. 2018;**488**:10-13

[32] Oueslati-Omrani R, Krimi S, Videau JJ, Khattech I, El Jazouli A, Jemal M. Structural investigations and calorimetric dissolution of manganese phosphate glasses. Journal of Crystalline Solids. 2014;**389**:66-71

[33] Oueslati-Omrani R, Krimi S, Videau JJ, Khattech I, El Jazouli A, Jemal M. Structural and thermochemical study of Na2O-ZnO-P2O5 glasses. Journal of Crystalline Solids. 2014;**390**: 5-12

[34] Hurt JC, Phillips JC. Structural role of zinc oxide in glasses in the system Na2O-ZnO-SiO2. Journal of the American Ceramic Society. 1967;**53**

[35] Wan MH, Wong PS, Hussin R, Lintang HO, Edud S. Structural and luminescence properties of Mn2+ ions doped calcium zinc borophosphate glasses. Journal of Alloys and Compounds. 2014;**595**:39-45

[36] Hassaan MY, Moustafa MG, Osouda K, Kubuki S, Nishiba T. 57Fe and 119Sn Mössbauer, XRD, FTIR and DC conductivity study of Li2O-Fe2O3- SnO2-P2O5 glass and glass ceramics. Journal of Alloys and Compounds. 2018; **765**:121-127

[37] Mohan S, Kaur S, Kaur P, Singh DP. Spectroscopic investigations of Sm3+ doped lead alumino-borate glasses containing zinc, lithium and barium oxides. Journal of Alloys and Compounds. 2018;**763**:486-495

[38] Maji BK, Jena H, Asuvathraman R. Electric conductivity and glass transition temperature (Tg) measurements on some selected glasses used for nuclear waste immobilization. Journal of Non-Crystalline Solids. 2016; **434**:102-107

[39] Berwal N, Dhankhar S, Sharma P, Kundu RS, Punia R, Kishore N. Physical, structural and optical characterization of silicate modified bismuth-boratetellurite glsses. Journal of Molecular Structure. 2017;**1127**:636-644

[40] Belova EV, Kolyagin YA, Uspenskaya IA. Structure and glass transition temperature of sodiumsilicate glasses doped with iron. Journal of Non-Crystalline Solids. 2015; **423-424**:50-57

[41] Hammad AH, Abdel-Hameed SAM, Margha FH. Effects of crystallization and microstructure on the dc electrical conductivity in the system xCuO-(70-x) MnO-30SiO2. Journal of Alloys and Compounds. 2015;**627**:423-429

[42] Oueslati-Omrani R, Khattech I, Jemal M. Standard formation enthalpy of Na2O-ZnO-P2O5 series glasses. Chemistry Africa. 2018;**1**:43-51

[43] Cimek J, Stepien R, Klimczak M, Zalewska I. Developpement of thermally stable glass from SiO2-Bi2O3- PbO-ZnO-BaO oxide system suitable for all solid photonic crystal fibers. Optical Materials. 2017;**73**:277-283

[44] Jlassi I, Sdiri N, Elhouichet H, Ferid M. Raman and impedance spectroscopy methods of P2O5-Li2O-Al2O3 system doped with MgO. Journal of Alloys and Compounds. 2016;**645**: 125-130

[45] Videau JJ, Flem L. Les verres phosphates de la spécificité de l'atome de phosphore à la formation, la structure, et la durabilité chimiques de phosphates vitreux. Institut de chimie de la matière condensée de Bordeaux; 27 October 2010. HAL Id: Cel-00530128

[46] Montagne L, Palavit G, Dalaval R. 31P NMR in (100-x)NaPO3-xZnO glasses. Journal of Non-Crystalline Solids. 1997;**215**:1-10

[47] Montagne L, Palavit G, Delaval R. Effect of ZnO on the properties of (100-x)NaPO3-xZnO glasses. Journal of Non-Crystalline Solids. 1998;**223**:43-47

[48] Zotov N, Schlenz H, Brendebach B, Modrow H, Hormes J, Reinauer F, et al. Effects of MnO doping on the structure of sodium metaphosphate glasses. 2003; **58a**:419-428

**193**

10<sup>−</sup><sup>4</sup>

**Chapter 10**

*Latefa Sail*

**Abstract**

7.5 10<sup>−</sup><sup>3</sup>

mol/l).

**1. Introduction**

reduction reaction [1].

to 10<sup>−</sup><sup>5</sup>

as sulfur S2 [3].

Temperature Influence on

Inhibitory Efficiency of Three

Phosphate Inhibitors by Mass Loss

The effect of temperature on steel samples immersed in concrete pore solutions contaminated by chlorides incorporating three inhibitors based on phosphate (Na3PO4, K2HPO4, and Na2PO3F) was studied by gravimetric measurements at several ranges: 298, 308, and 318 K. The results obtained for the use of these three products show that the inhibitory efficacy is lower at 318 K than that detected at 308 and 298 K of temperature. Also, we find that the best inhibitory efficiency at 298 K was detected for Na2PO3F (75.80% at 0.05 mol/l of concentra-

**Keywords:** temperature, concrete pores, corrosion inhibitors, phosphate,

Corrosion of reinforcement in concrete is one of the most dangerous pathologies that attack reinforced concrete structures; the means of protection against corrosion are varied and expensive. During this last decade, a new alternative has been adapted which is the application of corrosion inhibitors either as an adjunct to the mass of fresh concrete or by impregnation on the facing of hardened concrete. Several families of corrosion inhibitor products have been developed to prove their protective effect against steel reinforcement corrosion initiated by the penetration of chlorides through the pores of concrete. The best known are phosphates, borates, silicates and carbonates. One of the peculiarities of these ions is that their hydrolysis releases hydroxide ions which will have the effect of increasing the pH of the medium and thus passivating the steel. Moreover, in the presence of oxygen, the anions will form with the metal cation a very insoluble iron III phosphate which will clog the anodic surface and displace the cathodic

The required concentration of passivative inhibitor, often of the order of

pH, the presence of depassivating ions such as chlorides or reducing agents such

Temperature is one of the factors that can alter the behavior of a material in a corrosive environment. It can modify the metal-inhibitory interaction in a medium [4].

mol/l [2], it depends in fact on many factors such as temperature,

mol/l) and then Na3PO4 (61.48% at

tion) followed by K2HPO4 (65.05% at 2.5 10<sup>−</sup><sup>3</sup>

gravimetric measurement efficiency

#### **Chapter 10**

Jemal M. Structural and thermochemical study of Na2O-ZnO-P2O5 glasses. Journal of Crystalline Solids. 2014;**390**:

*Contemporary Topics about Phosphorus in Biology and Materials*

[41] Hammad AH, Abdel-Hameed SAM, Margha FH. Effects of crystallization and microstructure on the dc electrical conductivity in the system xCuO-(70-x) MnO-30SiO2. Journal of Alloys and Compounds. 2015;**627**:423-429

[42] Oueslati-Omrani R, Khattech I, Jemal M. Standard formation enthalpy of Na2O-ZnO-P2O5 series glasses. Chemistry Africa. 2018;**1**:43-51

[43] Cimek J, Stepien R, Klimczak M, Zalewska I. Developpement of

thermally stable glass from SiO2-Bi2O3- PbO-ZnO-BaO oxide system suitable for all solid photonic crystal fibers. Optical

Materials. 2017;**73**:277-283

125-130

[44] Jlassi I, Sdiri N, Elhouichet H, Ferid M. Raman and impedance spectroscopy methods of P2O5-Li2O-Al2O3 system doped with MgO. Journal of Alloys and Compounds. 2016;**645**:

[45] Videau JJ, Flem L. Les verres phosphates de la spécificité de l'atome de phosphore à la formation, la

structure, et la durabilité chimiques de phosphates vitreux. Institut de chimie de la matière condensée de Bordeaux; 27 October 2010. HAL Id: Cel-00530128

[46] Montagne L, Palavit G, Dalaval R. 31P NMR in (100-x)NaPO3-xZnO glasses. Journal of Non-Crystalline

[47] Montagne L, Palavit G, Delaval R. Effect of ZnO on the properties of (100-x)NaPO3-xZnO glasses. Journal of Non-Crystalline Solids. 1998;**223**:43-47

[48] Zotov N, Schlenz H, Brendebach B, Modrow H, Hormes J, Reinauer F, et al. Effects of MnO doping on the structure of sodium metaphosphate glasses. 2003;

Solids. 1997;**215**:1-10

**58a**:419-428

[34] Hurt JC, Phillips JC. Structural role of zinc oxide in glasses in the system Na2O-ZnO-SiO2. Journal of the American Ceramic Society. 1967;**53**

[35] Wan MH, Wong PS, Hussin R, Lintang HO, Edud S. Structural and luminescence properties of Mn2+ ions doped calcium zinc borophosphate glasses. Journal of Alloys and Compounds. 2014;**595**:39-45

[36] Hassaan MY, Moustafa MG, Osouda K, Kubuki S, Nishiba T. 57Fe and 119Sn Mössbauer, XRD, FTIR and DC conductivity study of Li2O-Fe2O3- SnO2-P2O5 glass and glass ceramics. Journal of Alloys and Compounds. 2018;

[37] Mohan S, Kaur S, Kaur P, Singh DP. Spectroscopic investigations of Sm3+ doped lead alumino-borate glasses containing zinc, lithium and barium

[38] Maji BK, Jena H, Asuvathraman R.

measurements on some selected glasses used for nuclear waste immobilization. Journal of Non-Crystalline Solids. 2016;

[39] Berwal N, Dhankhar S, Sharma P, Kundu RS, Punia R, Kishore N. Physical, structural and optical characterization of silicate modified bismuth-boratetellurite glsses. Journal of Molecular Structure. 2017;**1127**:636-644

[40] Belova EV, Kolyagin YA, Uspenskaya IA. Structure and glass transition temperature of sodiumsilicate glasses doped with iron. Journal

of Non-Crystalline Solids. 2015;

oxides. Journal of Alloys and Compounds. 2018;**763**:486-495

Electric conductivity and glass transition temperature (Tg)

**765**:121-127

**434**:102-107

**423-424**:50-57

**192**

5-12

## Temperature Influence on Inhibitory Efficiency of Three Phosphate Inhibitors by Mass Loss

*Latefa Sail*

#### **Abstract**

The effect of temperature on steel samples immersed in concrete pore solutions contaminated by chlorides incorporating three inhibitors based on phosphate (Na3PO4, K2HPO4, and Na2PO3F) was studied by gravimetric measurements at several ranges: 298, 308, and 318 K. The results obtained for the use of these three products show that the inhibitory efficacy is lower at 318 K than that detected at 308 and 298 K of temperature. Also, we find that the best inhibitory efficiency at 298 K was detected for Na2PO3F (75.80% at 0.05 mol/l of concentration) followed by K2HPO4 (65.05% at 2.5 10<sup>−</sup><sup>3</sup> mol/l) and then Na3PO4 (61.48% at 7.5 10<sup>−</sup><sup>3</sup> mol/l).

**Keywords:** temperature, concrete pores, corrosion inhibitors, phosphate, gravimetric measurement efficiency

#### **1. Introduction**

Corrosion of reinforcement in concrete is one of the most dangerous pathologies that attack reinforced concrete structures; the means of protection against corrosion are varied and expensive. During this last decade, a new alternative has been adapted which is the application of corrosion inhibitors either as an adjunct to the mass of fresh concrete or by impregnation on the facing of hardened concrete. Several families of corrosion inhibitor products have been developed to prove their protective effect against steel reinforcement corrosion initiated by the penetration of chlorides through the pores of concrete. The best known are phosphates, borates, silicates and carbonates. One of the peculiarities of these ions is that their hydrolysis releases hydroxide ions which will have the effect of increasing the pH of the medium and thus passivating the steel. Moreover, in the presence of oxygen, the anions will form with the metal cation a very insoluble iron III phosphate which will clog the anodic surface and displace the cathodic reduction reaction [1].

The required concentration of passivative inhibitor, often of the order of 10<sup>−</sup><sup>4</sup> to 10<sup>−</sup><sup>5</sup> mol/l [2], it depends in fact on many factors such as temperature, pH, the presence of depassivating ions such as chlorides or reducing agents such as sulfur S2 [3].

Temperature is one of the factors that can alter the behavior of a material in a corrosive environment. It can modify the metal-inhibitory interaction in a medium [4].

The variation of temperature affects the rate of corrosion. According to Liu and Weyer [5], an increase in temperature increases the rate of corrosion. This result was confirmed in carbonated concrete and also that subject to aggressive environments like chloride ions penetration.

The objective of this research is based on the analysis of the evolution of the inhibitory efficiencies of three phosphate inhibitors (Na3PO4, K2HPO4 and Na2PO3F) as a function of the temperature variation: 298, 303 and 313 K.

#### **2. Methods and measurements**

In this section, gravimetric tests were performed to characterize the influence of temperature on inhibition efficiency for the three phosphate inhibitors used in this study.

#### **2.1 Gravimetric measurements**

These measurements consist in determining the weight loss of a steel sample subjected to specified conditions of temperature and relative humidity; they are calculated on the basis of three tests to determine the average. The steel sample is polished with abrasive paper ranging from 120 up to 1000 grades using a polisher at a speed of 500 rpm, then rinsed in distilled water, dried with an electric dryer then we weigh the mass M1.

The steel samples are introduced into beakers containing 50 ml of electrolytic solution in an inclined position as shown in **Figure 1**, hermetically closed, then they are placed in a thermostatic bath while adjusting the desired temperature, after 24 h, the samples are removed from beakers then, rinsed in distilled water, degreasing is carried out with acetone and then dried with the electric dryer, after that we weigh the mass M2.

#### **2.2 Study medium**

The medium of this study is a concrete synthetic medium which simulates concrete pores contaminated by 3% of chlorides given in **Table 1**.

**195**

**Table 2.**

*Medium concentrations.*

**Figure 2.**

*Molecular structure of the three tested inhibitors.*

*Temperature Influence on Inhibitory Efficiency of Three Phosphate Inhibitors by Mass Loss*

The steel used is circular shaped with a diameter of 27 ± 1 and 2 ± 2 mm of thickness, and the procedure of gravimetric tests is detailed in [9]. The corrosion rate is

**1 l distilled water Ca(OH)2 NaOH KOH CaSO4 2H2O NaCl** Wt (g/l) 2 0.4 0.56 0.27 30

Cr = ΔM/S.t (mg/h. cm<sup>2</sup>

Hence, ΔM represents the difference between the initial mass M1 and the final mass M2 after a time "t" equal to immersion time by hours. "S" is the surface of the

This value of the corrosion rate is the average of three tests carried out under the same conditions for an optimal concentration at a definite time. The value of the

> Cr0 − Cr Cr0

This study describes the corrosion behavior of steel immersed in synthetic

concrete pore solutions contaminated by chlorides for three phosphateinhibitors (Na3PO4, K2HPO4 and Na2PO3F), their molecular structure is given

**Inhibitor Concentration (mol/l)** Na3PO4 7.5 × 10<sup>−</sup><sup>3</sup> K2HPO4 2.5 × 10<sup>−</sup><sup>3</sup> Na2PO3F 5 × 10<sup>−</sup><sup>2</sup>

) (1)

. 100 (3)

ΔM = M1 − M2 (2)

*DOI: http://dx.doi.org/10.5772/intechopen.88130*

determined by the following formula:

metal exposed to the electrolytic solution.

inhibitory efficiency is given by the following formula: IE (%) = \_

**2.3 Steel preparation**

*Synthetic medium of concrete [6–8].*

**Table 1.**

**3. Tested inhibitors**

**Figure 1.** *Position of steel sample.*

*Temperature Influence on Inhibitory Efficiency of Three Phosphate Inhibitors by Mass Loss DOI: http://dx.doi.org/10.5772/intechopen.88130*


**Table 1.**

*Contemporary Topics about Phosphorus in Biology and Materials*

ments like chloride ions penetration.

**2. Methods and measurements**

**2.1 Gravimetric measurements**

we weigh the mass M1.

weigh the mass M2.

**2.2 Study medium**

this study.

The variation of temperature affects the rate of corrosion. According to Liu and Weyer [5], an increase in temperature increases the rate of corrosion. This result was confirmed in carbonated concrete and also that subject to aggressive environ-

In this section, gravimetric tests were performed to characterize the influence of temperature on inhibition efficiency for the three phosphate inhibitors used in

These measurements consist in determining the weight loss of a steel sample subjected to specified conditions of temperature and relative humidity; they are calculated on the basis of three tests to determine the average. The steel sample is polished with abrasive paper ranging from 120 up to 1000 grades using a polisher at a speed of 500 rpm, then rinsed in distilled water, dried with an electric dryer then

The steel samples are introduced into beakers containing 50 ml of electrolytic solution in an inclined position as shown in **Figure 1**, hermetically closed, then they are placed in a thermostatic bath while adjusting the desired temperature, after 24 h, the samples are removed from beakers then, rinsed in distilled water, degreasing is carried out with acetone and then dried with the electric dryer, after that we

The medium of this study is a concrete synthetic medium which simulates

concrete pores contaminated by 3% of chlorides given in **Table 1**.

The objective of this research is based on the analysis of the evolution of the inhibitory efficiencies of three phosphate inhibitors (Na3PO4, K2HPO4 and Na2PO3F) as a function of the temperature variation: 298, 303 and 313 K.

**194**

**Figure 1.**

*Position of steel sample.*

*Synthetic medium of concrete [6–8].*

#### **2.3 Steel preparation**

The steel used is circular shaped with a diameter of 27 ± 1 and 2 ± 2 mm of thickness, and the procedure of gravimetric tests is detailed in [9]. The corrosion rate is determined by the following formula:

$$\text{Cr} = \Delta \text{M/St} \left( \text{mg/hcm}^2 \right) \tag{1}$$

$$
\Delta \mathbf{M} = \mathbf{M1} - \mathbf{M2} \tag{2}
$$

Hence, ΔM represents the difference between the initial mass M1 and the final mass M2 after a time "t" equal to immersion time by hours. "S" is the surface of the metal exposed to the electrolytic solution.

This value of the corrosion rate is the average of three tests carried out under the same conditions for an optimal concentration at a definite time. The value of the inhibitory efficiency is given by the following formula: IE (%) = \_

$$\text{IE } \{\text{\(\%\)}\} = \frac{\text{CrO - Cr}}{\text{CrO}} \text{100} \tag{3}$$

#### **3. Tested inhibitors**

This study describes the corrosion behavior of steel immersed in synthetic concrete pore solutions contaminated by chlorides for three phosphateinhibitors (Na3PO4, K2HPO4 and Na2PO3F), their molecular structure is given

**Figure 2.**

*Molecular structure of the three tested inhibitors.*


**Table 2.** *Medium concentrations.* in **Figure 2**. The optimal concentration which provides maximum efficiencies for the three products cited was extracted from a previous study [10] (see **Table 2**).

#### **4. Results and discussions**

**Table 3** records the mass loss results, relating to the evolution of corrosion rates as well as the inhibitory efficiencies as a function of the temperature variation: 298, 303 and 318 K for the three inhibitors.

It can be seen from the results shown in **Table 3** that the corrosion rates decrease in the presence of the corrosion inhibitor, it reached the maximum at the optimal concentration, for the first inhibitor sodium phosphate Na3PO4 the maximum efficiency 69.28% was detected at a concentration of 7.5 × 10<sup>−</sup><sup>3</sup> mol/l at 298 K, we can see clearly that the inhibitory efficiency slightly decrease as a function of temperature increase. Likewise for K2HPO4, the best efficiency 67.44% was detected at 298 K for a concentration of 2.5 × 10<sup>−</sup><sup>3</sup> mol/l, also, the increase of temperature affects the inhibitory efficiency which decrease following temperature increasing, the same remark was recorded for Na2PO3F the maximal efficiency 75.8% was detected at 298 K. This phenomenon can be explained by the fact that the anodic processes (oxidation components of steel) and cathodic (proton reduction in acidic medium) are thermally activated.

This results in a current of exchange which increases the corrosion rate. Hunkeler [11] has shown in his studies that the influence of temperature on the rate of corrosion is greater than that the resistivity of the concrete.


**Figure 3** shows the evolution of inhibitory efficiencies as a function of temperature variation for different concentrations of tested inhibitors.

#### **Table 3.**

*Evolution of corrosion rates and inhibitory efficiencies as a function of temperature variation.*

**197**

**Figure 3.**

*Na3PO4, K2HPO4 and Na2PO3F.*

*Temperature Influence on Inhibitory Efficiency of Three Phosphate Inhibitors by Mass Loss*

**Figure 3** illustrates the influence of temperature variation on the inhibitory efficacy of the three phosphate-based inhibitors. Certainly, temperature is one of the factors that can alter the behavior of a material in a corrosive environment. It

*Evolution of inhibitory efficiencies as a function of temperature variation for different concentrations of* 

The increase in temperature causes the instability of inhibitory molecules and also reduces the inhibitory efficacies which was detected in previous researches [5]. It can be seen from **Figure 4** that inhibitory efficiencies are highest in the optimum concentration for all the studied temperature ranges, although they decrease

can modify the metal-inhibitory interaction in a given environment [4].

*DOI: http://dx.doi.org/10.5772/intechopen.88130*

*Temperature Influence on Inhibitory Efficiency of Three Phosphate Inhibitors by Mass Loss DOI: http://dx.doi.org/10.5772/intechopen.88130*

**Figure 3.**

*Contemporary Topics about Phosphorus in Biology and Materials*

efficiency 69.28% was detected at a concentration of 7.5 × 10<sup>−</sup><sup>3</sup>

of corrosion is greater than that the resistivity of the concrete.

ture variation for different concentrations of tested inhibitors.

**) IE % Cr 10<sup>−</sup><sup>3</sup>**

**Table 2**).

**4. Results and discussions**

303 and 318 K for the three inhibitors.

at 298 K for a concentration of 2.5 × 10<sup>−</sup><sup>3</sup>

 **(mg/h.cm2**

medium) are thermally activated.

**C (mol/l) Cr 10<sup>−</sup><sup>3</sup>**

Na3PO4

K2HPO4

Na2PO3F

in **Figure 2**. The optimal concentration which provides maximum efficiencies for the three products cited was extracted from a previous study [10] (see

**Table 3** records the mass loss results, relating to the evolution of corrosion rates as well as the inhibitory efficiencies as a function of the temperature variation: 298,

It can be seen from the results shown in **Table 3** that the corrosion rates decrease in the presence of the corrosion inhibitor, it reached the maximum at the optimal concentration, for the first inhibitor sodium phosphate Na3PO4 the maximum

we can see clearly that the inhibitory efficiency slightly decrease as a function of temperature increase. Likewise for K2HPO4, the best efficiency 67.44% was detected

affects the inhibitory efficiency which decrease following temperature increasing, the same remark was recorded for Na2PO3F the maximal efficiency 75.8% was detected at 298 K. This phenomenon can be explained by the fact that the anodic processes (oxidation components of steel) and cathodic (proton reduction in acidic

This results in a current of exchange which increases the corrosion rate. Hunkeler [11] has shown in his studies that the influence of temperature on the rate

**Figure 3** shows the evolution of inhibitory efficiencies as a function of tempera-

0 1.4 — 2.5 — 3.1 — 5 × 10<sup>−</sup><sup>3</sup> 0.444 68.28 0.964 61.44 1.231 60.29 7.5 × 10<sup>−</sup><sup>3</sup> 0.43 69.28 0.798 68.08 1.0875 64.92 10<sup>−</sup><sup>2</sup> 0.768 45.14 1.49 40.4 1.8764 39.47

0 1.4 — 2.5 — 3.1 — 10<sup>−</sup><sup>3</sup> 0.4629 66.93 0.8889 64.44 1.2478 59.77 2.5 × 10<sup>−</sup><sup>3</sup> 0.4558 67.44 0.8737 65.04 1.1737 62.14 5 × 10<sup>−</sup><sup>3</sup> 0.4689 66.5 0.89126 64.34 1.251 59.45

2.5 × 10<sup>−</sup><sup>2</sup> 0.3411 75.63 0.8675 65.29 1.3976 54.91 5 × 10<sup>−</sup><sup>2</sup> 0.3386 75.8 0.6749 73 0.8986 71.01 7.5 × 10<sup>−</sup><sup>2</sup> 0.3414 75.61 0.8344 66.62 1.1842 61.8

*Evolution of corrosion rates and inhibitory efficiencies as a function of temperature variation.*

1.4 — 2.5 — 3.1 —

**298 K 303 K 313 K**

 **(mg/h.cm2**

mol/l at 298 K,

 **(mg/h.cm2**

**) IE %**

mol/l, also, the increase of temperature

**) IE % Cr 10<sup>−</sup><sup>3</sup>**

**196**

**Table 3.**

*Evolution of inhibitory efficiencies as a function of temperature variation for different concentrations of Na3PO4, K2HPO4 and Na2PO3F.*

**Figure 3** illustrates the influence of temperature variation on the inhibitory efficacy of the three phosphate-based inhibitors. Certainly, temperature is one of the factors that can alter the behavior of a material in a corrosive environment. It can modify the metal-inhibitory interaction in a given environment [4].

The increase in temperature causes the instability of inhibitory molecules and also reduces the inhibitory efficacies which was detected in previous researches [5].

It can be seen from **Figure 4** that inhibitory efficiencies are highest in the optimum concentration for all the studied temperature ranges, although they decrease

**Figure 4.**

*Evolution of inhibitory efficiencies as a function of temperature variation.*

slightly as a function of temperature increase. As a result, the maximum inhibitory efficacy at T 298, 303 and 318 K deduced using gravimetric measurements was confirmed by sodium monofluorophosphate (Na2PO3F), followed by potassium monohydrogenphosphate (K2HPO4) and thirdly sodium phosphate (Na3PO4).

These results are in good agreement with previous research that used the same inhibitory products [10].

Indeed, sodium monofluorophosphate has been the subject of several studies [12–14], and it has proven remarkable inhibitory properties especially in the case of its use in zinc phosphate baths [15–17].

The variation of the temperature influences the rate of corrosion and consequently the mechanism of the inhibition [18]. According to Liu and Weyer [5], an increase in temperature increases the rate of corrosion.

#### **5. Conclusions**

Direct measurements of both corrosion rates and inhibitory efficiencies as a function of inhibitor concentrations, have confirmed that sodium monofluorophosphate (Na2PO3F) offers the best corrosion protection under study conditions (temperature 298, 303 and 313 K); its inhibitory efficiency has exceeded 70% for these temperatures.

This inhibitor has been the subject of several previous studies [12, 19–22], its effectiveness against corrosion has been confirmed especially when used in a carbonated concrete [14, 23–26] and also than for concrete solutions contaminated by chlorides.

We can also conclude that increase of temperature affects inhibitory efficiencies, which is in good concordance with literature. For inhibitors based of phosphate, the increase of temperature has a slight influence on the inhibitory efficiency for the study temperatures, moreover, at higher temperatures, the molecular activation will be greater, which leads to an increase in corrosion rates.

#### **Acknowledgements**

My sincere gratitude and thanks go to the members of the granular materials team: corrosion prevention at the EOLE laboratory Department of Civil Engineering University of Tlemcen-Algeria.

**199**

**Author details**

Latefa Sail

*Temperature Influence on Inhibitory Efficiency of Three Phosphate Inhibitors by Mass Loss*

My 'research areas' focus on corrosion prevention and repair by inhibitors.

Faculty of Technology, Aboubekr Belkaid University, Tlemcen, Algeria

© 2019 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium,

\*Address all correspondence to: saillatefa@yahoo.fr

provided the original work is properly cited.

*DOI: http://dx.doi.org/10.5772/intechopen.88130*

**Conflict of interest**

*Temperature Influence on Inhibitory Efficiency of Three Phosphate Inhibitors by Mass Loss DOI: http://dx.doi.org/10.5772/intechopen.88130*

### **Conflict of interest**

*Contemporary Topics about Phosphorus in Biology and Materials*

slightly as a function of temperature increase. As a result, the maximum inhibitory efficacy at T 298, 303 and 318 K deduced using gravimetric measurements was confirmed by sodium monofluorophosphate (Na2PO3F), followed by potassium monohydrogenphosphate (K2HPO4) and thirdly sodium phosphate (Na3PO4).

These results are in good agreement with previous research that used the same

Indeed, sodium monofluorophosphate has been the subject of several studies [12–14], and it has proven remarkable inhibitory properties especially in the case of

The variation of the temperature influences the rate of corrosion and consequently the mechanism of the inhibition [18]. According to Liu and Weyer [5], an

Direct measurements of both corrosion rates and inhibitory efficiencies as a function of inhibitor concentrations, have confirmed that sodium monofluorophosphate (Na2PO3F) offers the best corrosion protection under study conditions (temperature 298, 303 and 313 K); its inhibitory efficiency has exceeded 70% for

This inhibitor has been the subject of several previous studies [12, 19–22], its effectiveness against corrosion has been confirmed especially when used in a carbonated concrete [14, 23–26] and also than for concrete solutions contaminated by chlorides. We can also conclude that increase of temperature affects inhibitory efficiencies, which is in good concordance with literature. For inhibitors based of phosphate, the increase of temperature has a slight influence on the inhibitory efficiency for the study temperatures, moreover, at higher temperatures, the molecular activation will

My sincere gratitude and thanks go to the members of the granular materials team: corrosion prevention at the EOLE laboratory Department of Civil

**198**

inhibitory products [10].

**Figure 4.**

**5. Conclusions**

these temperatures.

**Acknowledgements**

its use in zinc phosphate baths [15–17].

increase in temperature increases the rate of corrosion.

*Evolution of inhibitory efficiencies as a function of temperature variation.*

be greater, which leads to an increase in corrosion rates.

Engineering University of Tlemcen-Algeria.

My 'research areas' focus on corrosion prevention and repair by inhibitors.

#### **Author details**

Latefa Sail Faculty of Technology, Aboubekr Belkaid University, Tlemcen, Algeria

\*Address all correspondence to: saillatefa@yahoo.fr

© 2019 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/ by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

### **References**

[1] Oly M. Contribution à l'évaluation des capacités des glycérophosphates pour la maintenance dans le béton armé, [thesis]. France: University of Toulouse; 2011

[2] Buchler M, Corrosion inhibitors for reinforced concrete, in Corrosion Reinforced Concrete Structures. In: H. Böhni. Switzerland: Formerly Swiss Federal Institute of Technology; 2005. pp. 190-214

[3] Helie M. Matériaux métalliques— Phénomènes de corrosion, Edition Ellypses. Val of Essonne: University of Evry; 2015. ISBN: 9782340004023

[4] Khenadeki A. Etude théorique et expérimentale de l'effet d'inhibition de la corrosion d'un acier au carbone par les dérivées de base de Schiff en milieu acide chlorhydrique [thesis]. Algeria: University of Tlemcen; 2013

[5] Liu T, Weyer RW. Modeling the dynamic corrosion process in chloride contaminated concrete structures. Cement and Concrete Research. 1998;**28**(3):365-379. DOI: 10.1016/ S0008-8846(98)00259-2

[6] Ghods P, Isgor OB, Mcrae B, Millar T. The effect of concrete pore solution composition on the quality of passive oxide films on black steel reinforcement. Cement and Concrete Composites. 2009. DOI: 10.1016/j. cemconcomp.2008.10.0032-11

[7] Moragues A, Macias A, Andrade C. Equilibria of the chemical composition of the concrete pore solution. Part I: Comparative study of synthetic and extracted solutions. Cement and Concrete Research. 1987;**17**(2):173-182. DOI: 10.1016/0008-8846(87)90100-1

[8] Page CL, Vennesland O. Pore solution compositions and chloride binding capacity of silica fume cement paste.

Materials and Structures. 1983;**16**(1): 19-25. DOI: 10.1007/BF02474863.

[9] Sail L, Ghomari F, Bezzar A, Benali O. Mass loss for assessment of the inhibitory efficiency of products to basis of phosphate. Canadian Journal on Environmental, Construction and Civil Engineering. 2011;**2**(5):111-117. Availabe from: http://www.ampublisher.com/ June%202011/ECCE-1106-018-Massloss-for-assessment-of-the-inhibitory. pdf

[10] Sail L. Etude de la performance d'inhibiteurs de corrosion à base de phosphate pour les constructions en béton armé [thesis]. Algeria: Tlemcen; 2013

[11] Hunkeler F. Grundlagen der korrosion und der potential messing baustahlbetonbauten, ASTRA Brücken unter halts for schung. Zürich: Verein Schweizer Strassen fachleute (VSS); Report No. 510; 1994

[12] Douche-Portanguen A, Prince W, Lutz T, Arliguie G. Detection or quantitative analysis of a corrosion inhibitor, the sodium monofluorophosphate, in concrete. Cement and Concrete Composites. 2005;**27**(6):679-687. DOI: 10.1016/j. cemconcomp.2004.11.002

[13] Pujol Lesueur VN. Etude du mécanisme d'action du monofluorophosphate de sodium comme inhibiteur de la corrosion des armatures métalliques dans le béton [thesis]. Paris, France: University of Pierre et Marie Curie; 2004

[14] Duprat M, Bonnel A, Dabisi F. Les monofluorophosphates de zinc et de potassium en tant qu'inhibiteurs de la corrosion d'un acier au carbone en solution de NaCl à 3%. Journal of Applied Electrochemistry. 1983;**13**:317-323

**201**

*Temperature Influence on Inhibitory Efficiency of Three Phosphate Inhibitors by Mass Loss*

toothpaste by capillary electrophoresis. Journal of Chromatography. 1999. DOI: 10.1016/S0021-9673(96)00926-0

[23] Alonso C, Andrade C, Argiz C, Malric B. Na2PO3F as inhibitor of corroding reinforcement in carbonated

concrete. Cement and Concrete Research. 1996;**26**(3):405-415. DOI: 10.1016/S0008-8846(96)85028-9

[24] Vézina D. Performance des inhibiteurs de corrosion, direction des laboratoires et chaussées. Technical

[26] Dhouibi L, Triki E, Salta M, Rodrigues P, Raharinaivo A. Studies on corrosion inhibition of steel reinforcement by phosphate and nitrite. Materials and Structures. 2003;**36**(8):530-540. DOI: 10.1007/

BF02480830

[25] Benzina Mechmeche L, Dhouibi L, Ben Ouezdou M, Triki E, Zucchi F. Investigation of the early effectiveness of an amino-alcohol based corrosion inhibitor using simulated pore solutions and mortar specimens. Cement and Concrete Composites. 2008;**30**:167-173. DOI: 10.1016/j. cemconcomp.2007.05.007

Newsletter. 1997;**2**:3

*DOI: http://dx.doi.org/10.5772/intechopen.88130*

[15] Zimmermann D, Mun˜oz AG, Schule JW. Formation of Zn-Ni alloys in the phosphating of Zn layers. Surface & Coatings Technology. 2005;**197**(2): 260-269. DOI: 10.1016/j.surfcoat.

[16] Kashyap A. Effects of water chemistry, temperature, gaseous cavitation et phosphate inhibitors on concrete corrosion [Master]. USA: University of Virginia; 2008

[17] Simescu F. Elaboration des revtements de phosphates de zinc sur armature à béton. Etude de leur comportement à la corrosion en milieu neutre et alcalin [thesis]. France:

[18] Khouikhi F. Etude de l'efficacité de deux inhibiteurs de corrosion dans les milieux multiphasiques (Eau, huile et gaz) [thesis]. Boumerdes, Algeria: University of M'Hamed Bougara; 2007

[19] Soylev TA, Richardson MG. Corrosion inhibitors for steel in concrete. State-of-the-art report. Construction and Building Materials. 2008;**22**(4):609-622. DOI: 10.1016/j.

[20] Farcas F, Chaussadent T, Fiaud C, Mabille I. Determination of the sodium monofluorophosphates in a hardened cement paste by ion chromatography. Analytica Chimica Acta. 2002;**472**(1-2):37-43. DOI: 10.1016/S0003-2670(02)00978-9

conbuildmat.2006.10.013

[21] Talamge JM, Biemer TA.

[22] Wang P, Li SF, Lee HK. Simultaneous determination of monofluorophosphate and fluoride in

Determination of potassium nitrate and sodium monofluorophosphate in the presence of phosphate and sulphate by high resolution ion chromaltography. Journal of Chromatography. 1987;**410**:494-499. DOI: 10.1016/S0021-9673(00)90084-0

University of Lyon; 2008

2004.07.129

*Temperature Influence on Inhibitory Efficiency of Three Phosphate Inhibitors by Mass Loss DOI: http://dx.doi.org/10.5772/intechopen.88130*

[15] Zimmermann D, Mun˜oz AG, Schule JW. Formation of Zn-Ni alloys in the phosphating of Zn layers. Surface & Coatings Technology. 2005;**197**(2): 260-269. DOI: 10.1016/j.surfcoat. 2004.07.129

[16] Kashyap A. Effects of water chemistry, temperature, gaseous cavitation et phosphate inhibitors on concrete corrosion [Master]. USA: University of Virginia; 2008

[17] Simescu F. Elaboration des revtements de phosphates de zinc sur armature à béton. Etude de leur comportement à la corrosion en milieu neutre et alcalin [thesis]. France: University of Lyon; 2008

[18] Khouikhi F. Etude de l'efficacité de deux inhibiteurs de corrosion dans les milieux multiphasiques (Eau, huile et gaz) [thesis]. Boumerdes, Algeria: University of M'Hamed Bougara; 2007

[19] Soylev TA, Richardson MG. Corrosion inhibitors for steel in concrete. State-of-the-art report. Construction and Building Materials. 2008;**22**(4):609-622. DOI: 10.1016/j. conbuildmat.2006.10.013

[20] Farcas F, Chaussadent T, Fiaud C, Mabille I. Determination of the sodium monofluorophosphates in a hardened cement paste by ion chromatography. Analytica Chimica Acta. 2002;**472**(1-2):37-43. DOI: 10.1016/S0003-2670(02)00978-9

[21] Talamge JM, Biemer TA. Determination of potassium nitrate and sodium monofluorophosphate in the presence of phosphate and sulphate by high resolution ion chromaltography. Journal of Chromatography. 1987;**410**:494-499. DOI: 10.1016/S0021-9673(00)90084-0

[22] Wang P, Li SF, Lee HK. Simultaneous determination of monofluorophosphate and fluoride in toothpaste by capillary electrophoresis. Journal of Chromatography. 1999. DOI: 10.1016/S0021-9673(96)00926-0

[23] Alonso C, Andrade C, Argiz C, Malric B. Na2PO3F as inhibitor of corroding reinforcement in carbonated concrete. Cement and Concrete Research. 1996;**26**(3):405-415. DOI: 10.1016/S0008-8846(96)85028-9

[24] Vézina D. Performance des inhibiteurs de corrosion, direction des laboratoires et chaussées. Technical Newsletter. 1997;**2**:3

[25] Benzina Mechmeche L, Dhouibi L, Ben Ouezdou M, Triki E, Zucchi F. Investigation of the early effectiveness of an amino-alcohol based corrosion inhibitor using simulated pore solutions and mortar specimens. Cement and Concrete Composites. 2008;**30**:167-173. DOI: 10.1016/j. cemconcomp.2007.05.007

[26] Dhouibi L, Triki E, Salta M, Rodrigues P, Raharinaivo A. Studies on corrosion inhibition of steel reinforcement by phosphate and nitrite. Materials and Structures. 2003;**36**(8):530-540. DOI: 10.1007/ BF02480830

**200**

*Contemporary Topics about Phosphorus in Biology and Materials*

Materials and Structures. 1983;**16**(1): 19-25. DOI: 10.1007/BF02474863.

[9] Sail L, Ghomari F, Bezzar A, Benali O. Mass loss for assessment of the inhibitory efficiency of products to basis of phosphate. Canadian Journal on Environmental, Construction and Civil Engineering. 2011;**2**(5):111-117. Availabe from: http://www.ampublisher.com/ June%202011/ECCE-1106-018-Massloss-for-assessment-of-the-inhibitory.

[10] Sail L. Etude de la performance d'inhibiteurs de corrosion à base de phosphate pour les constructions en béton armé [thesis]. Algeria: Tlemcen;

[11] Hunkeler F. Grundlagen der korrosion und der potential messing baustahlbetonbauten, ASTRA Brücken unter halts for schung. Zürich: Verein Schweizer Strassen fachleute (VSS);

[12] Douche-Portanguen A, Prince W,

Lutz T, Arliguie G. Detection or quantitative analysis of a corrosion inhibitor, the sodium monofluorophosphate, in concrete. Cement and Concrete Composites. 2005;**27**(6):679-687. DOI: 10.1016/j.

cemconcomp.2004.11.002

[13] Pujol Lesueur VN. Etude du mécanisme d'action du monofluorophosphate de sodium comme inhibiteur de la corrosion des armatures métalliques dans le béton [thesis]. Paris, France: University of

Pierre et Marie Curie; 2004

1983;**13**:317-323

[14] Duprat M, Bonnel A, Dabisi F. Les monofluorophosphates de zinc et de potassium en tant qu'inhibiteurs de la corrosion d'un acier au carbone en solution de NaCl à 3%. Journal of Applied Electrochemistry.

Report No. 510; 1994

pdf

2013

[1] Oly M. Contribution à l'évaluation des capacités des glycérophosphates pour la maintenance dans le béton armé, [thesis]. France: University of Toulouse;

[2] Buchler M, Corrosion inhibitors for reinforced concrete, in Corrosion Reinforced Concrete Structures. In: H. Böhni. Switzerland: Formerly Swiss Federal Institute of Technology; 2005.

[3] Helie M. Matériaux métalliques— Phénomènes de corrosion, Edition Ellypses. Val of Essonne: University of Evry; 2015. ISBN: 9782340004023

[4] Khenadeki A. Etude théorique et expérimentale de l'effet d'inhibition de la corrosion d'un acier au carbone par les dérivées de base de Schiff en milieu acide chlorhydrique [thesis]. Algeria:

University of Tlemcen; 2013

S0008-8846(98)00259-2

[6] Ghods P, Isgor OB, Mcrae B, Millar T. The effect of concrete pore solution composition on the quality of passive oxide films on black steel reinforcement. Cement and Concrete Composites. 2009. DOI: 10.1016/j. cemconcomp.2008.10.0032-11

[7] Moragues A, Macias A, Andrade C. Equilibria of the chemical composition of the concrete pore solution. Part I: Comparative study of synthetic and extracted solutions. Cement and Concrete Research. 1987;**17**(2):173-182. DOI: 10.1016/0008-8846(87)90100-1

[8] Page CL, Vennesland O. Pore solution compositions and chloride binding capacity of silica fume cement paste.

[5] Liu T, Weyer RW. Modeling the dynamic corrosion process in chloride contaminated concrete structures. Cement and Concrete Research. 1998;**28**(3):365-379. DOI: 10.1016/

2011

**References**

pp. 190-214

*Edited by David G. Churchill, Maja Dutour Sikirić, Božana Čolović and Helga Füredi Milhofer*

This book addresses a diverse set of topics regarding phosphorus chemistry, namely phosphates and closely related chemical systems. Divided into two sections, chapters cover such topics as phosphate dynamics and phosphates in biomaterials. This volume is a useful reference for scholars and researchers and will inspire readers to make future discoveries in the field.

Published in London, UK © 2020 IntechOpen © werbeantrieb / iStock

Contemporary Topics about Phosphorus in Biology and Materials

Contemporary Topics

about Phosphorus in Biology

and Materials

*Edited by David G. Churchill, Maja Dutour Sikirić, Božana Čolović* 

*and Helga Füredi Milhofer*