Open Access is an initiative that aims to make scientific research freely available to all. To date our community has made over 100 million downloads. It’s based on principles of collaboration, unobstructed discovery, and, most importantly, scientific progression. As PhD students, we found it difficult to access the research we needed, so we decided to create a new Open Access publisher that levels the playing field for scientists across the world. How? By making research easy to access, and puts the academic needs of the researchers before the business interests of publishers.
We are a community of more than 103,000 authors and editors from 3,291 institutions spanning 160 countries, including Nobel Prize winners and some of the world’s most-cited researchers. Publishing on IntechOpen allows authors to earn citations and find new collaborators, meaning more people see your work not only from your own field of study, but from other related fields too.
To purchase hard copies of this book, please contact the representative in India:
CBS Publishers & Distributors Pvt. Ltd.
www.cbspd.com
|
customercare@cbspd.com
The potential use of calcium phosphate cements in endodontic therapy is an active area of research. Hydroxyapatite is one of the most commonly used calcium phosphate materials in medicine and dentistry. Biocompatibility of hydroxyapatite is closely related to its chemical composition, similar to dental and bony tissues. Recent studies have focused on new and modified formulations of calcium-phosphate-based biomaterials with improved mechanical and maintained favorable biological properties. Recently, two non-commercial new nanomaterials based on calcium silicates and hydroxyapatite have been synthesized. One is a calcium silicate system of tricalcium and dicalcium silicates (CS), and the other one is a mixture of the calcium silicate system and hydroxyapatite (HA-CS). Both CS and HA-CS are nanostructural materials. Particle size affects cement hydration and consequently setting time and final quality of the cement. Fast setting is a clear clinical advantage while cement composition and internal nanostructure are expected to provide biological behavior in vital tissues. The problem with furcation perforation repair is still not agreed upon as no currently available materials meet all the requirements of an ideal repair material as defined in the literature. Therefore, this study aimed to compare the tissue reaction of two new repair materials for furcation perforations.
Department for Restorative Dentistry and Endodontics, School of Dental Medicine, University of Belgrade, Belgrade, Serbia
Violeta Petrović
Department for Restorative Dentistry and Endodontics, School of Dental Medicine, University of Belgrade, Belgrade, Serbia
Vanja Opačić Galić
Department for Restorative Dentistry and Endodontics, School of Dental Medicine, University of Belgrade, Belgrade, Serbia
Smiljana Paraš
Faculty of Science and Mathematics, Department of Zoology, University of Banja Luka, Banja Luka, Republic of Srpska, Bosnia and Herzegovina
Vukoman Jokanović
“Vinča” Institute of Nuclear Sciences, University of Belgrade, Belgrade, Serbia
Slavoljub Živković
Department for Restorative Dentistry and Endodontics, School of Dental Medicine, University of Belgrade, Belgrade, Serbia
*Address all correspondence to: dr.marijanapopovic@gmail.com
1. Introduction
Root perforations represent communication between the root canal system of the tooth and the outer surface of the tooth root. Most often, they arise as a result of pathological processes, such as deep caries or resorption, and are revealed by a detailed clinical examination or X-ray analysis. However, most perforations occur mechanically, due to iatrogenic factors (during the opening of the access cavity, finding the entrance to the canals, mechanical processing of the root canal, or during the preparation phase for an intracanal post) [1, 2, 3, 4]. The resulting iatrogenic perforations affect the success of endodontic therapy and the long-term prognosis of the tooth. Root perforations leading to failure of endodontic treatment account for approximately 10% of all failures [5]. Errors during endodontic therapy can occur at any stage of the work. Predisposing factors for the occurrence of perforations are the presence of denticles, calcification, obliterated canals, internal resorptions, inability to identify canal orifices, extensive caries, inclined or rotated teeth, and presence of canal posts [1, 2, 3, 4]. Root perforation is a serious complication that requires rapid diagnosis, and timely and adequate therapy. Root perforations often lead to an inflammatory response and destruction of periodontal tissue and alveolar bone. The proliferation of granulomatous tissue, the proliferation of the epithelium, the formation of an endoperiodontal lesion and finally the loss of teeth can occur depending on the size of the perforation, its localization, and the intensity of the chronic inflammatory reaction [1, 6, 7].
Perforations of single-rooted teeth have a better prognosis than perforations of multi-rooted teeth. Perforations in the furcation area of multi-rooted teeth are a special problem. Iatrogenic perforations in the area of the furcation occur in approximately 2–12% of endodontically treated teeth and can have serious consequences on the outcome of endodontic therapy, but also on the preservation of the tooth itself [1, 8]. This perforation is an open door for bacteria to enter, either from the root canal or from the periodontal tissue, which causes an intense inflammatory response. The consequences of the inflammatory reaction can be bone resorption, growth of the gingival epithelium, and the appearance of a fistula, which further worsens the prognosis of the tooth [1].
Factors that determine the prognosis of a tooth with a perforation include the size and localization of the defect, the time elapsed from the diagnosis to the closure of the perforation, the duration of the contamination, the physical and chemical properties of the material used for its repair, or its sealing capabilities [9, 10]. The prognosis would be relatively good if the perforation was quickly detected and closed with a biocompatible material [11, 12]. There are now various biocompatible materials on the market that are used to close perforations in order to reduce the inflammatory reaction in the surrounding tissues [13, 14, 15]. The ideal material for closing these perforations should be non-toxic, non-resorbable, radiopaque, bactericidal, or bacteriostatic, and should provide a hermetic seal and prevent microleakage [16, 17].
In the literature, calcium silicate-based materials are most often used for such clinical situations. They are used in direct pulp capping procedures, closure of the root apex in apex surgery, therapy of root perforations, pulpotomy, and apexification [18, 19]. Biocompatibility, bioactivity, and sealing ability of calcium silicate cement have been proven in many studies due to their potential for dentin, cementum, bone, and periodontal ligament regeneration [18, 19]. The good biological properties of calcium-silicate cement are associated with calcium hydroxide, which is released during the bonding of the material and stimulates the proliferation and differentiation of various cells for tissue regeneration [20].
Mineral trioxide aggregate (MTA) was introduced in endodontics in 1990, as a new material with the property of closing the communication between the tooth and the outer surface of the root [18, 19]. In vitro and in vivo studies have shown that MTA has significant sealing ability and marginal adaptation [18, 19]. This material has a long initial setting time (3 h), as a result of its chemical composition and hygroscopic nature, and it also contains traces of heavy metals [21]. MTA contains bismuth oxide as an X-ray contrast agent, which interferes with cement hydration processes [22] and can react with dentin collagen, leading to tooth discoloration [23]. Newer generations of MTA cement have a slightly shorter initial setting time, as a result of changes in chemical composition, and contain non-toxic radiocontrast agents such as zirconium oxide [24]. The main disadvantages of MTA are the long setting time and difficult handling of this material, despite its optimal sealing ability and other advantages [25].
Despite its biocompatibility and bioconductivity, hydroxyapatite has not found its place in endodontic therapy, mainly due to inadequate mechanical properties. In order to improve the mechanical properties and bioactivity of hydroxyapatite, it was added to calcium silicate cements [26].
Technological progress and the influence of nanotechnology in the field of biomedical research have led to the emergence of new nanostructured materials. This new field of science deals with controlling matter, energy and/or information at the atomic and molecular level and enables the synthesis of new materials that are increasingly similar to natural biomolecular structures [27].
Nanomedicine is a branch of medicine that is based on the medical application of nanotechnologies, through the application of nanomaterials, nanoelectronic biosensors, and molecular nanotechnology [27, 28]. The very size of nanomaterial particles (< 100 nm), which is similar to the size of biological molecules and structures (proteins 5 nm, organelles 100–200 nm), points to the possible application of nanomaterials in in vivo and in vitro biomedical research [29]. Nanomedicines made from nanomaterials and nanoparticles (compared to drugs made from the same materials in a classical way) have up to ten times greater interactive surface, which unequivocally indicates a possible improvement in the pharmacokinetic and pharmacodynamic properties of the drug. Chen (2012) points out that the bioactivity, biocompatibility, stability, and mechanical properties of nanomaterials are determined by their composition, structure, morphology, and crystal size, as well as the method of synthesis [30].
Nanomaterials have revolutionized medicine and are increasingly being used in this field. These materials have the ability to mimic the surface properties of natural tissues, are highly cytocompatible and biocompatible, and therefore show excellent properties for use in tissue engineering and regenerative medicine [27]. The pronounced activity of nanoparticles improves the hydration of nanostructured calcium silicate cement, improving their hardening and bonding, as well as their physical and chemical properties [31]. However, there are concerns about the biological behavior of these materials, as nanoparticles usually deposit in mitochondria, causing structural damage to cells. Commercial nanostructured calcium silicate cement with hydroxyapatite (BioAggregate, Innovative bioceramics, Vancouver, BC, Canada) shows similar toxicity in cell cultures, but lower systemic toxicity compared to commercial MTA, which was related to differences in the content of heavy metals and differences in production [32].
Two new experimental nanostructured cements were recently developed at the Institute of Nuclear Sciences “Vinča” according to the recipe of V. Jokanović and associates. The goal was to synthesize materials with good biological properties, short bonding time, and without heavy metals and bismuth oxide. The first cement (CS) was based on dicalcium and tricalcium silicate, and the second (HA-CS) was a mixture of hydroxyapatite with CS, in a ratio of 2:1. CS was synthesized using a hydrothermal sol-gel methodology and self-propagating combustion waves [33]. Hydroxyapatite was synthesized by the hydrothermal method. Both materials contain barium sulfate as an X-ray contrast agent. The setting time of CS and HA-CS is 10 minutes and 15 minutes. In vitro testing of CS and HA-CS showed no genotoxic effects on human cells [34]. Their biocompatibility was confirmed in studies where these materials were applied in subcutaneous tissue in rats or as direct pulp capping material in rabbits [35, 36].
Examination of the cytotoxicity of these materials on cell culture of fibroblasts and MRC-5 cells showed better results in comparison with MTA [37].
Calcium silicate hydrate gel (CSH) and calcium hydroxide, the main soluble fraction of cement that dissolves into Ca2+ and OH− ions, are formed by the hydration of calcium silicate materials. The reaction of released calcium with phosphates from tissue fluids represents the physicochemical basis of the bioactivity of calcium silicate materials [38]. It was established that the formation of an apatite layer on the surface of the material is not only a consequence of the release of calcium, but also the formation of Si-OH groups on the surface of the cement, which acts as centers of apatite nucleation and precipitation. Consequently, the synergistic action of calcium silicate hydrate gel as an apatite nucleator and calcium from dissolved calcium hydroxide as a precipitation accelerator is responsible for the rapid precipitation of apatite. Additionally, the hydrophilic substrate facilitates the bond with the apatite layer due to the presence of OH- groups on the surface of the cement [39]. The formation of apatite on the surface of calcium silicate actually goes through several stages. Amorphous calcium phosphate first forms on the surface of the material and then transforms into apatite, which later matures into type B carbonate apatite, which represents the biological phase of hydroxyapatite of bone, cementum, and dentin.
Calcium silicate cements are superior to most of the endodontic materials known to date in terms of their biocompatibility, bioactivity, and sealing properties. The fact that they possess a strong inductive potential for the regeneration of damaged tissues has contributed to the fact that these cements are now considered the materials of choice for numerous clinical indications.
The improved physical properties of calcium silicate cements obtained by the application of nanotechnology in terms of lower degree of solubility, higher structural stability of the material, and shorter bonding time, as well as lower cytotoxicity and better biocompatibility of these materials [40, 41], are sufficient reasons for examining their application in vivo conditions. The chemical nature of the material and the method of synthesis should ensure satisfactory biological behavior of these materials in living tissues. Quality marginal sealing, that is, adequate marginal adaptation of the material, should prevent the flow of tissue fluids and consequent bacterial microleakage, which is a significant factor for the long-term success of endodontic treatment.
The aim of this study was to evaluate the inflammatory reactions of the periradicular tissue and the formation of calcified tissue after implantation of CS and HA-CS in the furcation defects of the teeth of Vietnamese pigs and in the root canals of the teeth of rabbits.
2.1 Application of CS and HA-CS in furcation defects of teeth of Vietnamese pigs
Experimental research was carried out at the Institute of Surgery of the Faculty of Veterinary Medicine, University of Belgrade, and at the Institute for Biological Research “Siniša Stanković,” University of Belgrade.
Permission for experimental work with animals was obtained from the Ethics Committee of the Faculty of Veterinary Medicine and from the Ethics Committee of the Faculty of Dentistry (16/29), conducted according to international standards ISO10993-2 (Requirements for animal welfare) and ISO 7405.
Three Vietnamese pigs (Sus scrofa verus), both sexes, aged 24 months and with an average weight of 25 kg were the animal model in this experimental research. The protocol of the European Good Laboratory Practice (86/609/EEC), which implies the implementation of the main principles of asepsis and antisepsis, the realization of the experiment in the minimum necessary time without physical and mental pain of the animals (International Organization for Standardization, 1997), was respected during the work.
The animals were housed in the experimental animal facility at the Faculty of Veterinary Medicine, University of Belgrade, during the experiment. Each experimental animal was housed in an individual cage in a controlled environment with a controlled diet and daily professional care. The animals were provided with appropriate care, nutrition, hygienic conditions, and their health condition was checked 3 times a day.
The animals were deprived of food for 6 hours before the operation, and water for 3 hours before the operation, in order to rule out complications during the experiment. Premedication with atropine in a dose of 0.03–0.04 mg/kg by intramuscular injection was performed in all three animals, and after 15 minutes, the animals were put under general anesthesia by administering xylazine (2% Xylazin, Cp Pharma, Bergdorf, Germany) in a dose of 1.5–2 mg/kg and ketamine (Laboratorio Sanderso S.A., Santiago, Chile) at a dose of 20–25 mg/kg intramuscularly. The average duration of anesthesia was about 100 minutes.
The surgical procedure was performed under aseptic conditions and in a way that ensures minimal trauma. Each tooth was individually cleaned, dried, and disinfected (30% hydrogen and 70% ethanol). Access cavities were formed on the lower premolars with a round diamond bur. Coronary pulp tissue was removed with a sterile, carbide, round bur. Accidental perforation of the floor of the cavum dentis in the area of the furcation of the tooth was made with the same bur, 2 mm in diameter. The cavities were washed with 5 ml of physiological solution, and hemostasis was performed with sterile cotton balls, gently and without pressure. Freshly mixed experimental, nanostructured materials, CS, and HA-CS mixed with distilled water in a ratio of 2:1 were placed on the perforation [33]. All cavities were closed with glass ionomer cement (GC FUJI VIII, GC Corporation, Tokyo, Japan) as a definitive filling.
For the next 3 days, the animals were housed in the inpatient unit of the Institute of Surgery, under constant medical supervision with analgesic therapy (Butorphanol 0.1 mg/kg/tm/i.m. for 6 h). Appropriate care, nutrition, and hygienic conditions were provided to the animals during the observation period, and their health conditions were checked three times a day.
Experimental animals were sacrificed after 30 days, by intravenous injection of 10 ml of pentobarbiturate solution (100 mg/kg).
The mandible and maxilla were separated from the rest of the skull after the removal of the soft tissues, and the tissue was fixed in 10% formalin for 48 hours. The tissue for histological analysis was taken in the form of block sections, where each block consisted of an experimental tooth with the surrounding bone. The samples were decalcified in a decalcification solution: 8% HCl from 37% (v/v) concentrate and 10% formic acid (HCOOH) from 89% (v/v) concentrate (pH = 5) at 37°C. The success of complete decalcification was evaluated subjectively and experientially. The tissue was fixed in a circular tissue processor (Leica TP 1020, Germany) after decalcification, and then molded in paraffin blocks.
Serial tissue sections of 4 μm thickness (8 from each sample) were cut from paraffin molds in the mesiodistal direction. The sections were caught on a glass slide and placed at a temperature of 56–68°C (melting point of paraffin) for one hour, to fix the samples to the glass slide and dry them. The preparations were then stained with hematoxylin-eosin (HE) and Goldner trichrome staining. H&E was used for stereological measurements, and Goldner trichrome was used for bone mineralization detection.
Stereological measurements were performed to observe the difference in the representation of bone tissue and matrix cells in the CS and HA-CS groups. Microscopic preparations were analyzed by optical microscopy using an LM Leica microscope at magnifications ×10, ×40, ×100, ×200, and ×450. Pathohistological parameters were analyzed qualitatively, semiquantitatively, and quantitatively. Histomorphometric analysis was performed according to the cellularity and thickness of the newly formed tissue. Parameters were scored using a scoring system from 1 to 4, according to the modified criteria of Accorinte et al. [42].
2.1.1 Results
After the application of CS and HA-CS materials in the artificial furcation defects of the teeth of Vietnamese pigs, no inflammatory reaction was recorded in the periodontal and bone tissue of any sample (grade 1). No giant cells or microorganisms were found (score 1). A small number of material particles were detected in most of the CS and HA-CS group samples (grade 2).
The results of cytohistological and stereological analyzes of bone after implantation of CS and HA-CS materials in the furcation of porcine teeth are shown in Tables 1 and 2; and Figures 1–6.
Results of morphometric and stereological values of the periodontium in pig teeth with CS and HA-CS (mean value ± SD, p<0.05).
Statistically significant difference.
Figure 1.
Micrographs of histological sections of pig teeth: (A) CS, Goldner, and (B) HA-CS, H&E, applied material in the perforation area (yellow arrow), newly formed bone (black arrow), complete healing at the perforation site (green arrow), ×40.
Figure 2.
Micrographs of histological sections of pig teeth: (A) CS and (B) HA-CS, histochemical technique Picrosirius red, applied material in the area of the furcation (blue arrow), newly formed bone (green arrow), collagen depots (black arrow), ×40.
Figure 3.
Micrographs of histological sections of pig teeth: (A) CS and (B) HA-CS, histochemical technique Picrosirius red, periodontium (green arrow), newly formed bone (black arrow), blood vessels (white arrow), ×40.
Figure 4.
Micrographs of histological sections of the peridontia of pig teeth: (A) CS, and (B) HA-CS, different width of peridontium (black bracket), blood vessels (red arrow), and peridontium (black arrow), Goldner, ×450.
Figure 5.
Micrographs of histological sections of pig teeth: (A) CS, and (B) HA-CS, H&E, periodontium (white arrow), osteoblast activity surface (black arrow), and newly formed bone (yellow arrow), ×100.
Figure 6.
Micrographs of histological sections of pig teeth: (A) CS, and (B) HA-CS, Goldner, remnants of applied materials (red arrow), periodontium (yellow arrow), newly formed bone (white arrow), and newly formed tendon tissue (black arrow), ×100.
The results showed that the volume and numerical density of osteocytes, as well as the number of osteocytes, was higher after implantation of HA-CS material than after implantation of CS material in the area of the tooth furcation, which represented a statistically significant difference. The volume density of the mineralized extracellular matrix in the bone tissue in the HA-CS group was higher compared to the CS group. A higher percentage of newly formed bone was observed in the HA-CS group (25.66%) compared to the CS group. The mean bone marrow volume density was lower in the HA-CS group compared to the CS group, and the numerical bone marrow density was higher in the HA-CS group compared to the CS group. These data imply that the mitogenic effect led to an increase in the number of bone marrow cells and not to an increase in its volume in the bone tissue. This was the reason to make individual measurements of volume and numerical densities, as well as the number of mesenchymal cells and fibroblasts in the bone marrow.
The results of these measurements showed that the volume and numerical density, as well as the number of the mentioned cells in the bone marrow, were higher in the HA-CS group compared to the CS group. Also, less extracellular substance was observed in the same tissue in the bone marrow. Stereological parameters of blood capillaries showed a slight increase in values in the tissue of the HA-CS group compared to the CS group, which is a consequence of increased mitosis of all types of bone tissue cells.
Newly formed calcified tissue was observed in all samples after the application of both CS and HA-CS (Figure 1). In all samples of the HA-CS group, the implanted material was completely separated from the adjacent tissue by newly formed, properly structured, calcified continuous tissue (grade 1) (Figure 1B), whose thickness was 190–210 μm. Incomplete closure of the perforation by newly formed calcified tissue (green arrow) 125–160 μm thick is present in CS material in most samples. The newly formed calcified tissue in the CS group mostly had an irregular morphology (grade 2).
HA-CS performed significantly better than CS in terms of continuity and thickness of newly formed calcified tissue (p = 0.008).
Immunostaining for collagen (Picrosirius red) is a method by which, based on the detection of collagen, the quality of newly formed bone is determined. The results showed that the newly formed bone under the HA-CS material was of better quality and denser (Figure 2) and that there were more depots of extracellular collagen in the form of strips compared to the CS group (black arrow).
The newly formed bone was more regular and denser in contact with the periodontium in the HA-CS group than in the CS group (Figures 2 and 3). A thin layer of periodontium (green arrow) is observed in most samples under the applied material (Figure 3). The activity of osteoblasts leading to new bone formation (black arrow) and the presence of newly formed blood vessels (white arrows) are also observed. Blood vessels are more numerous and larger in the HA-CS than in the CS group.
The results of morphometric and stereological measurements of the periodontium tissue are shown in Table 2; and Figures 4–6.
The thickness of the newly formed periodontium in the area of the furcation in the HA-CS group was 26.88% higher than in the CS group. The surface of the periodontium in the area of the furcation was 16.34% higher in the HA-CS group compared to the CS group. The volume density of the periodontium was 27.75% higher in HA-CS compared to CS, which was a statistically significant difference.
Histological and morphometric and stereological data show that the newly formed periodontal tissue was better and more regular in HA-CS material (Figure 4, black bracket) (Table 2). The connective tissue of the periodontium under the applied HA-CS material was more regular and dense (Table 2, Figure 4, black arrow), and the blood vessels were narrower and more numerous (red arrow). The volume density of the periodontium was statistically significantly higher (p = 0.021673) in the HA-CS group compared to the periodontium of the CS group. The blood vessels of the periodontium in the CS group were wider, dilated, with remnants of clustered erythrocytes.
Histological analysis shows that the newly formed periodontium in the HA-CS group is thicker compared to the CS group. Mesenchymal cells with osteoblastic differentiation were observed at the periphery of the newly calcified tissue. These cells are more numerous and fill the entire space between the periodontium and newly formed bone in the HA-CS material, while they are found in smaller numbers in the CS group (Figure 5).
The periodontium is histologically more regular, thicker, and more reactive to staining in the HA-CS group compared to CS. Newly formed bone (white arrow) is observed below the periodontium and the beginnings of tendon tissue formation (black arrow) are observed below it (Figure 6). Newly formed bone is histologically more regular and thicker, as well as tendon tissue which is thicker in teeth with HA-CS compared to CS.
2.2 Implantation of CS and HA-CS in root canals of rabbit teeth
Experimental research was carried out at the Institute of Surgery of the Faculty of Veterinary Medicine, University of Belgrade. Permission for experimental work with animals was obtained from the Ethics Committee of the Faculty of Dentistry, University of Belgrade, Serbia (Protocol No. 36/21/2013), conducted according to international standards ISO10993-2 (Requirements for animal welfare) and ISO 7405 [41].
Four rabbits from the genus Orictolagus cuniculus, aged 12 months and with an average weight of 4 kg, were included in the research.
Animals were kept in standard, individual cages, given ad libitum access to water, and standard rabbit chow. Daily monitoring of the animals was performed while the experiment lasted. Before the surgery, the animals were put under general anesthesia with Xylazine (2% Xylazine, Czech Republic) 35 mg/kg body weight and Ketamidor (100% Ketamidor 100 mg/ml, Richter Pharma AG, Austria) 5 mg/kg body weight.
The working field was disinfected with 5% tincture of iodine. Round diamond bur was used for preparing the class I cavities on the upper and lower central incisors. Access cavities were formed and coronary pulp tissue was removed with sterile, round, and carbide burs. Root canals were instrumented with K files #40 (VDV Gmbh, Germany) after extirpation of the radicular pulp and irrigated with 5 ml saline between each instrument. A new set of endodontic instruments was used for each animal. Then, the canals were dried with paper points and filled with freshly mixed material. Experimental and nanostructured cements, CS and HA-CS, were mixed with distilled water in a ratio of 2:1 [33]. The control material, mineral trioxide aggregate (White MTA, Angelus® Solu oes odontologicas, Londrina, Brazil) was mixed according to the manufacturer’s instructions, in a powder-to-liquid ratio of 3:1. MTA was applied in the right maxillary incisors of all animals. In both mandibular incisors and left maxillary incisors of two animals were applied CS, HA-CS was applied in the left maxillary incisors and both mandibular incisors of remaining two animals. The materials were introduced with a lentulo spiral and condensed with a hand compactor into the root canals. The cavities were then closed with voice ionomer cement (GC FUJI VIII, GC Corporation, Tokyo, Japan). Postoperatively, the animals received subcutaneously an analgesic (Butorphanol, 10 mg/ml, Richter Pharma AG Austria), 0.1 mg/kg body weight, every 8 h for the next three days and an antibiotic (Baitril®, 25 mg/ml, KVP Pharma und Veterinar Produkte GmbH), 10 mg/kg body weight, daily for the next five days. After 28 days, animals were sacrificed by intravenous injection of 10 ml of pentobarbital solution (pentobarbital sodium salt 100 mg ml-1, Sigma-Aldrich Chemie GmbH, Steinheim, Germany).
The treated teeth together with the bone tissue of the upper or lower jaw were cut with a diamond disc in the form of block sections and fixed in 10% formalin after removing the soft tissues and separating the upper and lower jaw. The samples were decalcified in a decalcification solution: 8% HCl from 37% (v/v) concentrate and 10% formic acid (HCOOH) from 89% (v/v) concentrate (pH = 5) at 37°C. The success of complete decalcification was evaluated subjectively and experientially. The tissue was fixed in a circular tissue processor (Leica TP 1020, Germany) after decalcification and then molded in paraffin blocks.
Serial tissue sections (eight per sample) of 5 μm thickness were cut from paraffin blocks and stained with hematoxylin eosin (HE) according to standard procedure. Glass histological slides were analyzed with an optical microscope (Olympus 5 microscope) using the morphometric software package “Cell-B” (Olympus), at magnifications of 40x, 100x, and 200x, by an experienced pathologist who did not know the types of examined materials. Histological parameters were analyzed qualitatively (presence of inflammation, general tissue condition, continuity, and morphology of calcified tissue), semi-quantitatively (presence of giant cells, particles of material, and microorganisms), and quantitatively (intensity of inflammatory reaction, thickness of calcified tissue). Histomorphometric analysis was performed according to the cellularity and thickness of the calcified tissue. Parameters were scored using a scoring system from 1 to 4 according to the modified criteria of Accorinte et al. [42].
2.2.1 Results
Half of the samples showed no inflammatory reaction after CS material implantation (grade 1) (Figure 7). In two samples, inflammatory reactions were moderate (grade 3). In one sample, inflammatory reaction was pronounced (grade 4), with deep tissue infiltration of inflammatory cells and abscess formation (grade 3). Particles of implanted material were detected in all samples, but in different numbers (grades 2–4). Giant cells were detected in half of the samples—(grade 1), and giant cells were detected in the other half in small numbers (grade 2). Microorganisms were not detected in any sample.
Figure 7.
CS. Photomicrograph of calcified tissue with proliferation of the connective tissue, moderate cellularity with slight macrophage infiltration. Discontinuous newly formed calcified tissue (black arrow) with a clear border between old and new osteoid (black line). Irregular structure of newly formed calcified tissue (HE, 200×).
The tissue was unchanged after implantation of the HA-CS material (grade 1) in most samples (Figure 8). Mild inflammatory reaction (grade 2) was detected in two samples (score 2). Material particles were detected in most samples (grade 2). Microorganisms and giant cells were not detected (grade 1).
Figure 8.
HA-CS. Continuous calcified tissue with lamellar structure (HE, 40×). Calcified tissue with regular mineralization. Viable osteocyte is presented in newly formed bone (black arrows).
Different inflammatory reaction was observed after MTA implantation (grades 2–3), with inflammatory cells near the implanted material (grade 2) (Figure 9). Material particles were detected in all samples (grade 2). Few giant cells were found in most samples (grade 2). Microorganisms were not detected (grade 1).
Figure 9.
MTA. Foci of fibrovascular proliferation (black arrow) in partially discontinuous newly formed calcified tissue (HE, 40×). Regular structure of calcified tissue with many osteocytes. Material particles (blue arrow) (HE, 200×).
There were no statistically significant differences in the intensity of inflammatory reactions between the tested materials. There were statistically significant differences between HA-CS and CS regarding the extent of inflammation (p = 0.004).
Newly formed calcified tissue was mostly irregular in morphology in samples with CS (score 2), deposited in a thickness greater than 250 μm (scores 1–2) in most samples, but discontinuous with foci of fibrovascular tissue (scores 2–3) (Figures 7A, B).
The implanted material was completely separated from the adjacent tissue by newly formed, regularly structured, and calcified continuous tissue in most samples with HA-CS (score 1). The thickness of the newly formed tissue varied between 150 and 250 μm (scores 1–2). Mesenchymal cells with osteoblastic differentiation were observed at the periphery of the newly calcified tissue (Figures 8A, B).
New calcified tissue was deposited in small amounts, up to 150 μm thick, in all samples with MTA (score 3). The newly formed calcified tissue was irregularly structured and discontinuous with foci of fibrovascular proliferation (scores 2–3) (Figures 9A, B).
HA-CS performed significantly better than MTA and CS in terms of continuity of newly formed calcified tissue (p = 0.03 and p = 0.010, respectively). There were significant differences in calcified tissue thickness between CS and MTA (p = 0.004) and between HA-CS and MTA (p = 0.012).
The complex interaction between the material and the host tissue is best demonstrated by in vivo tests. These tests, in addition to biocompatibility, enable the evaluation of the biofunctionality of the material. In these two experimental studies, the effects of the materials were evaluated after their implantation in artificially created perforations in the furcation area of the teeth of Vietnamese pigs and application in the root canals of rabbit teeth.
In a rabbit study, the CS and HA-CS materials induced an inflammatory reaction of the periradicular tissue that was similar in intensity to the control material (MTA). Inflammatory reactions were rated as mild to moderate in most samples, indicating a good tolerance of the host tissue to the applied materials. These findings are consistent with the results of other authors who examined the biocompatibility of materials with similar chemical composition [42].
In another study on Vietnamese pigs, no inflammatory reaction was noted in any sample after implantation of CS and HA-CS in the furcation area of the tooth. The results of de Silva and the authors show that the inflammatory reaction is most intense in the first seven days after the application of the tested material and that the intensity of the inflammation decreases over time [43].
Inflammatory reactions after the application of calcium silicate cements are the result of the release of calcium hydroxide during the setting of the material. High pH causes local tissue necrosis with the development of local inflammatory reactions [44]. As a consequence of the alkaline pH, silicate cement induces the expression of proinflammatory cytokines (IL-6 and IL-8) [45]. It is known that tissue necrosis is the initiator of the mineralization process [46]. Some studies show that repair processes could start even without necrosis or acute inflammation [47]. As the material sets as a function of time, the amount of calcium hydroxide released from calcium silicate cements decreases [47]. By binding the material, favorable conditions are created for the beginning of the reparation process.
Although no statistically significant differences were found between CS, HA-CS, and MTA in the rabbit tissue regarding the inflammatory response, the tissue condition in the samples with HA-CS was rated as the best. This finding may be a consequence of the composition of such material. HA-CS consists mainly of hydroxyapatite with a slightly lower pH value than CS and MTA [35]. Lower pH values are thought to promote alkaline phosphatase activity but cause a smaller zone of surface necrosis compared to highly alkaline materials such as calcium silicate cements [48].
The low number of giant cells in samples with CS and MTA, or their absence in samples with HA-CS in both experiments, indicates a low activity of tissue histocytes and a good tissue tolerance to the implanted materials.
Newly formed calcified tissue was observed in all samples of the investigated materials in both experiments. It confirms that the examined nanostructured cements have an inductive potential. Previous studies also confirm the formation of mineralized tissue after the implantation of materials with the similar chemical composition [49].
CS and HA-CS belong to the group of bioactive materials that have the ability to release biologically active ions. The main soluble component of these cements is calcium hydroxide. It is released during the binding of the material. Considering that calcium silicate materials are characterized by a long bonding time, calcium hydroxide is released over several weeks [38]. Calcium hydroxide dissolves calcium and hydroxyl ions in contact with tissue and tissue fluids. The continuous release of calcium ions from the material is considered to be crucial for the induction of calcified tissue formation. In addition to its role in chemotaxis, calcium regulates cell proliferation, differentiation, and mineralization [39].
Calcium-releasing materials have been confirmed to induce the proliferation of periodontal fibroblasts, the growth and differentiation of pulp cells, osteoblasts, osteoblast-like cells, and cementoblasts [20].
The release of hydroxyl ions from implanted material is associated with tissue mineralization processes. An increase in alkaline phosphatase (ALP) occurs as a result of high pH, leading to the expression of growth factors and the formation of calcified tissue.
The tested materials also have Si ions in their composition, which influence the bioactivity of the material [50], and the proliferation and differentiation of cells similar to osteoblasts. High concentrations of Si ions (> 30 ppm) can inhibit osteoclast growth and resorption processes, but can also increase the level of ALP that participates in the mineralization of newly calcified tissue [51].
The result of the application of both nanostructured materials is a thicker layer of calcified tissue compared to MTA. Materials synthesized by the sol-gel method, such as CS and HA-CS in these studies, have improved bioactivity compared to the same materials obtained by other methods [52]. The topography of the surface of materials is related to their chemical composition and structure, and affects the activity of cells, especially their adhesion and vitality [47]. These results can be explained by the nanostructure of CS and HA-CS particles, which is similar to bone structure.
The newly formed calcified tissue, associated with HA-CS in both experiments, was continuous and without foci of vascularized fibroblast proliferation, which was not the case with MTA and CS. Unlike CS and MTA, HA-CS contains phosphate ions that could be related to this histological finding. Studies show that similar or better quality of calcified tissues are described after the application of calcium silicate cements containing phosphate ions, compared to pure calcium silicate cement [49]. The authors attributed these findings to the greater amount of phosphate ions available for hydroxyapatite formation. The present studies confirm that materials with hydroxyapatite have a greater potential for tissue mineralization than MTA [53].
There were no observed microorganisms in any sample of the implanted materials. The presence of microbes usually correlates with inadequate crown restorations and subsequent microleakage. GIC is material with good sealing properties and it may be the reason for the obtained result in this study. However, it must be emphasized that with this type of histochemical staining, microorganisms are difficult to detect and can be removed during tissue preparation for histological analysis.
The application of CS and HA-CS into the furcation defects of the teeth of Vietnamese pigs showed a complete absence of tissue inflammatory reaction, while this response was minimal after the application of these materials into the root canals of rabbit teeth, similar to the control MTA. CS and HA-CS were more effective than MTA in inducing the formation of calcified tissue after implantation in the root canals of rabbit teeth. The best organized newly formed calcified tissue was observed after the application of HA-CS in root canals of rabbits and root perforation of Vietnamese pigs. The present results serve as a solid basis for further biological studies of CS and HA-CS.
The authors have stated explicitly that there is no conflict of interest in connection with this chapter.
References
1.Tsesis I, Fuss Z. Diagnosis and treatment of accidental root perforations. Endodontic Topics. 2006;13(1):95-107. DOI: 10.1111/j.1601-1546.2006.00213.x
2.Estrela C, Biffi JC, Moura MS, Lopes HP. Treatment of endodontic failure. In: Estrela C, editor. Endodontic Science. 2nd ed. São Paulo: Artes Médicas; 2009. pp. 917-952
3.Roda RS, Gettleman BH. Nonsurgical retreatment. In: Hargreaves KM, Berman LH, editors. Cohen’s Pathways of the Pulp. 11th ed. St. Louis: Elservier; 2016. pp. 324-386
4.Estrela C, Pécora JD, Estrela CRA, Guedes OA, Silva BS, Soares CJ, et al. Common operative procedural errors and clinical factors associated with root canal treatment. Brazilian Dental Journal. 2017;28(2):179-190. DOI: 10.1590/0103-6440201702451
5.Ingle JI. A standardized endodontic technique utilizing newly designed instruments and filling materials. Oral Surgery, Oral Medicine, and Oral Pathology. 1961;14:83-91. DOI: 10.1016/0030-4220(61)90477-7
6.Touré B, Faye B, Kane AW, Lo CM, Niang B, Boucher Y. Analysis of reasons for extraction of endodontically treated teeth: A prospective study. Journal of Endodontia. 2011;37(11):1512-1515. DOI: 10.1016/j.joen.2011.07.002
7.Holland R, Otobani Filho JA, Souza V, Nery MJ, Bernabé PF, Dezan JE. Mineral trioxide aggregate repair of lateral root perforations. Journal of Endodontia. 2001;27(4):281-284. DOI: 10.1097/00004770-200104000-00011
8.Farzaneh M, Abitbol S, Friedman S. Treatment outcome in endodontics: The Toronto study. Phases I and II: Orthograde retreatment. Journal of Endodontia. 2004;30:627-633
9.Eldeeb M, Tabibi A, Jensen JR. An evaluation of the use of amalgam, cavit and calcium hydroxide in the repair of furcation perforations. Journal of Endodontia. 1982;8:459-466
10.Oswald R. Procedural accidents and their repair. Dental Clinics of North America. 1979;23:593-616
11.Haghgoo R, Arfa S, Asgary S. Microleakage of CEM cement and ProRoot MTA as furcal perforation repair materials in primary teeth. Iranian Endodontic Journal. 2013;8:187-190
12.Gutmann JL, Harrison JW. Posterior endodontic surgery: Anatomical considerations and clinical techniques. International Endodontic Journal. 1985;18:8-34
13.Kakani AK, Veeramachaneni C, Majeti C, Tummala M, Khiyani L. A review on perforation repair materials. Journal of Clinical and Diagnostic Research. 2015;9(9):ZE09-ZE13. DOI: 10.7860/JCDR/2015/13854.6501
14.Estrela C, Decurcio D, Rossi-Fedele G, Silva JA, Guedes OA, Borges AH, et al. Root perforations: A review of diagnosis, prognosis and materials. Brazilian Oral Research. 2018;32:133-146. DOI: 10.1590/1807-3107bor-2018.vol32.0073. Epub 18 Oct 2018
15.Kakani AK, Veeramachaneni C. Sealing ability of three different root repair materials for furcation perforation repair: An in vitro study. Journal of Conservative Dentistry. 2020;23(1):62-65. DOI: 10.4103/JCD.JCD_371_19
16.Silveira CM, Sánchez-Ayala A, Lagravère MO, Pilatti GL, Gomes OM. Repair of furcal perforation with mineral trioxide aggregate: Long-term follow-up of 2 cases. Journal of the Canadian Dental Association. 2008;74:729-733
17.Felippe WT, Felippe MC, Rocha MJ. The effect of mineral trioxide aggregate on the apexification and periapical healing of teeth with incomplete root formation. International Endodontic Journal. 2006;39:2-9. DOI: 10.1111/j.1365-2591. 2005.01037.x
18.Torabinejad M, Parirokh M, Dummer PMH. Mineral trioxide aggregate and other bioactive endodontic cements: An updated overview—Part I: Vital pulp therapy. International Endodontic Journal. 2018;51:177-205
19.Torabinejad M, Parirokh M, Dummer PMH. Mineral trioxide aggregate and other bioactive endodontic cements: An updated overview—Part II: Other clinical applications and complications. International Endodontic Journal. 2018;51:284-317
20.Silva LAB, Pieroni KAMG, Nelson-Filho P, et al. Furcation perforation. Periradicular tissue response to biodentine as a repair material by histopathologic and indirect immunofluorescence analyses. Journal of Endodontia. 2017;43:1137-1142
21.Kum KY, Kim EC, Yoo YJ. Et al: Trace metal contents of three tricalcium silicate materials: MTA angelus, micro mega MTA and bioaggregate. International Endodontic Journal. 2014;47:704-710
22.Camilleri J. Hydratation characteristics of calcium silicate cements with alternative radiopacifiers used as root-end filling materials. Journal of Endodontia. 2010;36:502-508
23.Ramos JC, Palma PJ, Nascimento R, et al. 1-year in vitro evaluation of tooth discoloration induced by 2 calcium silicate-based cements. Journal of Endodontia. 2016;42:1403-1407
24.Camilleri J, Sorrentino F, Damidot D. Investigation of the hydration and bioactivity of radiopacified tricalcium silicate cement, Biodentine and MTA angelus. Dental Materials. 2013;29:580-593
25.Parirokh M, Torabinejad M. Mineral trioxide aggregate: A comprehensive literature review—Part III: Clinical applications, drawbacks and mechanism of action. Journal of Endodontia. 2010;36:400-413
26.Zhao Q , Qian J, Zhou H, Yuan Y, Mao Y, Liu C. In vitro osteoblast-like and endothelial cells response to calcium silicate/calcium phosphate cement. Biomedical Materials. 2010;5:1-8
27.Jokanović V. Nanomedicina najveći izazov 21. Vekas: Data Status; 2012. p. 180
28.Gui C, Dai X, Cui D. Advances of nanotechnology applied to biosensors. Nano Biomedicine & Engineering. 2011;3(4):260-273
29.Fan JP, Kalia P, Di Silvio L, Huang J. In vitro response of human osteoblasts to multi-step sol-gel derived bioactive glass nanoparticles for bone tissue engineering. Materials Science and Engineering: C. 2014;36:206-214
30.Chenab F, Zhub Y, Wub J, et al. Nanostructured calcium phosphates: Preparation and their application in biomedicine. Nano Biomedicine & Engineering. 2012;4(1):41-49
31.Saghiri MA, Godoy FG, Gutmann JL, Lotfi M, Asatourian A, Sheibani N, et al. The effect of pH on solubility of nano-modified endodontic cements. Journal of Conservative Dentistry. 2014;17:13-17
32.De Deus G, Canabarro A, Alves G, Linhares A, Senne MI, Granjeiro JM. Optimal cytocompatibility of a bioceramic nanoparticulate cement in primary human mesenchymal cells. Journal of Endodontia. 2009;35:1387-1390
33.Jokanović V, Čolović B, Jokanović B, Živković S. Superplastic, quick-bonding endodontic mixtures and their hydration. International Journal of Applied Ceramic Technology. 2015;12(S2):E83-E91
34.Opačić-Galić V, Petrović V, Živković S, et al. New nanostructural biomaterials based on active calcium silicate systems and hydroxyapatite:Characterization and genotoxicity in human peripheral blood lymphocytes. International Endodontic Journal. 2013;46:506-516
35.Petrović V, Opačić Galić V, Jokanović V, Jovanović M, Basta Jovanović G, Živković S. Biocompatibility of a new nanomaterial based on calcium silicate implanted in subcutaneous connective tissue of rats. Acta Veterinaria-Beograd. 2012;62:697-708
36.Opačić-Galić V, Petrović V, Jokanović V, Živković S. Histological evaluation of tissue reactions to newly synthetized calcium silicate-and-hydroxyapatite-based bioactive materials-In vivo study. Srpski Arhiv za Celokupno Lekarstvo. 2017;145:370-377
37.Petrović V, Opačić-Galić V, Živković S, Nikolić B, Danilović V, Miletić V, et al. Biocompatibility of new nanostructural materials based on active silicate systems and hydroxyapatite: in vitro and in vivo study. International Endodontic Journal. 2015;48(10):966-975. DOI: 10.1111/iej.12391. [Epub 2014 Nov 5]
38.Sarkar NK, Caicedo R, Ritwik P, Moiseyeva R, Kawashima I. Physicochemical basis of the biologic properties of mineral trioxide aggregate. Journal of Endodontia. 2005;31:97-100
39.Gandolfi MG, Ciapetti G, Perut F, et al. Biomimetic calcium-silicate cements aged in simulated body solutions. Osteoblast response and analyses of apatite coating. Journal of Applied Biomaterials & Biomechanics. 2009;7:160-170
40.Jokanović V, Cicović B, Prokić BB, Tomanović N, Popović Bajić M, Živković S. Subchronic systemic toxicity of new endodontic material based on calcium hydroxyapatite and calcium silicates. Advances Mater Science and Engineering. 2018;2018:6. Article ID: 8493439. DOI: 10.1155/2018/8493439
41.Opačić Galić V, Petrović V, Popović Bajić M, Jokanovič V, Živkovič S, Nikolić B, et al. Physical properties and biocompatibility of Nanostructural biomaterials based on active calcium silicate systems and hydroxyapatite. In: Chaughule RS, editor. Dental Applications of Nanotechnology. Vol. Chapter 13. Switzerland AG: Springer Nature; 2018. pp. 247-271. DOI: 10.1007/978-3-319-97634-1_13
42.Accorinte Mde L, Holland R, Reis A, Bortoluzzi MC, Murata SS, Dezan E, et al. Evaluation of mineral trioxide aggregate and calcium hydroxide cement as pulp-capping agents in human teeth. Journal of Endodontia. 2008;34:1-6
43.da Silva GF, Guerreiro-Tanomaru JM, Sasso-Cerri E, Tanomaru-Filho M, Cerri PS. Histological and histomorphometrical evaluation of furcation perforations filled with MTA, CPM and ZOE. International Endodontic Journal. 2011;44(2):100-110. DOI: 10.1111/j.1365-2591.2010.01803.x. Epub 2010 Oct 12
44.Tran XV, Gorin C, Willig C, et al. Effect of a calcium-silicate-based restorative cement on pulp repair. Journal of Dental Research. 2012;91:1166-1171
45.Chen CC, Shie MY, Ding SJ. Human dental pulp cell response to new calcium silicate-based endodontic materials. International Endodontic Journal. 2011;44:836-842
46.Sangwan P, Sangwan A, Duha J, Rohilla A. Tertiary dentinogenesis with calcium hydroxide: A review of proposed mechanisms. International Endodontic Journal. 2013;46:3-19
47.Anselme K. Osteoblast adhesion on biomaterials. Biomaterials. 2000;21:667-681
48.Silva EJNL, Rosa TP, Herrera DR, Jacinto RC, Gomes BPFA, Zaia AA. Evaluation of cytotoxicity and physicochemical properties of calcium silicate-based endodontic sealer MTA Fillapex. Journal of Endodontia. 2013;39:274-277
49.Zarrabi MH, Javidi M, Jafarian AH, Joushan B. Histologic assessment of human pulp response to capping with mineral trioxide aggregate and a novel endodontic cement. Journal of Endodontia. 2010;36:1778-1781
50.Huan Z, Chang J. Calcium-phosphate-silicate composite bone cement: Self-setting properties and in vitro bioactivity. Journal of Materials Science: Materials in Medicine. 2009;20:833-841
51.Pietak AM, Reid JW, Stott MJ, Sayer M. Silicon substitution in the calcium phosphate bioceramics. Biomaterials. 2007;28:4023-4032
52.Li P, de Groot K. Better bioactive ceramics through sol-gel process. Journal of Sol-Gel Science and Technology. 1994;2:797-801
53.Zhang S, Yang X, Fan M. Bioaggregate and iRoot BP plus optimize the proliferation and mineralization ability of human dental pulp cells. International Endodontic Journal. 2013;46:923-929
Open access peer-reviewed chapter
