Among the patients who visited the periodontal department of Chosun University Dental Hospital for implant surgery, we selected those who underwent sinus or alveolar ridge augmentation with Bontree® 1.0-granule with at least 1 year of functional loading. Patients were excluded if they (i) were not followed up, and (ii) did not have radiographs for evaluation. Three male patients were selected, and their ages ranged from 63 to 77 years (Table 1). This study was approved by the Institutional Review Board (IRB) of Chosun University Dental Hospital in Gwangju, Korea (CUDHIRB-2103-008). All treatment plans and procedures were explained to the patients, and they provided consent for surgery.

PMH, past medical history; ISQ (B/L), implant stability quotient (buccal/lingual); HTN, hypertension; CVD, cardiovascular disease; DM, diabetes mellitus.

#### **3. Results**

#### *3.1. Basic Research*

3.1.1. Material Characterization

Figure 1A shows the FE-SEM micrographs of two different granule sizes of Bontree®. At low magnification, both Bontree® granules had irregular shapes with clean and rough surface morphologies, and no cracks or defects were found on any Bontree® granule surface. High-magnification FE-SEM images clearly showed that micron-sized crystals uniformly covered the surface of the Bontree® granules. There were no significant differences between the surface qualities and morphologies of the two Bontree® granules.

**Figure 1.** (**A**) Representative low- and high-magnification FE-SEM images of the Bontree® granules: Bontree® 0.5-granule and Bontree® 1.0-granule refer to Bontree® granules with diameters of 0.5 mm and 1.0 mm, respectively. (**B**) The XRD spectra of the Bontree® 0.5-granule and Bontree® 1.0-granule. The OCP and HA peaks are indicated by \* and ೠ, respectively. FE-SEM, field-emission scanning electron microscopy; XRD, X-ray diffraction; OCP, octacalcium phosphate; HA, hydroxyapatite.

Figure 1B shows the XRD patterns of Bontree® 0.5-granule and Bontree® 1.0-granule. XRD was conducted to examine the presence of OCP and HA in Bontree® grafts. The presence of diffraction peaks at 2θ = 10.8◦, 26.3◦, 31.8◦, 32.2◦, 32.9◦, 39.8◦, 46.6◦, 49.3◦, and 53.2◦ indicated the formation of apatite (HA), while the strongest intensity of the OCP diffraction peak (100) was observed at 4.7◦ in both samples [15,16]. Phase analysis using XRD showed that OCP is the main constituent of Bontree®, regardless of size.

#### 3.1.2. In Vitro Study

The RANKL and ALP activities of the MG63 cells were assessed using ELISA to assess the role of Bontree® in bone formation and remodeling (Figure 2A,B). For comparison, a commercial biphasic calcium phosphate ceramic (MBCP+TM) was also assessed as a reference [17]. Prior to in vitro testing, each specimen was compressed into tablets. Bontree® refers to both Bontree® 0.5-granule and Bontree® 1.0-granule samples.

**Figure 2.** (**A**) RANKL activity and (**B**) ALP activity of MG63 cells in Bontree® and MBCP+TM cultures after 7 and 14 days of culture. Statistical significance was set at \* *p* < 0.05 and \*\* *p* < 0.01, respectively. RANKL, receptor activator of NF-κB ligand; ALP, alkaline phosphatase.

The level of RANKL, an osteoclast differentiation factor [18], was lower in cultures with Bontree® than in those with MBCP+TM during both culturing periods; it was 11.3 ng/mL and 19.0 ng/mL for Bontree® and 14.1 ng/mL and 28.6 ng/mL for MBCP+TM after 7 and 14 days, respectively. In contrast, the activity of ALP, an osteoblast differentiation marker [19], was significantly elevated in cultures with Bontree® (17.9 ng/mL and 22.4 ng/mL at 7 and 14 days, respectively), showing 1.9- and 1.6-fold higher values than in cultures with MBCP+TM (9.2 ng/mL and 13.9 ng/mL at 7 and 14 days, respectively).

#### 3.1.3. In Vivo Study

On the basis of the results of in vitro testing, Bontree® 0.5-granule and Bontree® 1.0-granule were selected for further evaluation of new bone formation in vivo. Figure 3A shows low- and high-magnification histological images of transverse cross-sections stained using hematoxylin and eosin; the cortical bone tissue stained light red or pink, and collagen fibers stained pale pink; the black areas represent the implanted Bontree® grafts [20]. To examine the size effect of bone grafts, two Bontree® grafts (Bontree® 0.5-granule and Bontree® 1.0-granule) were assessed in this in vivo study. At both 4 and 12 weeks after surgery, bone regeneration was observed in the histological images without any sign of local inflammation or rejection of the bone graft. In the control group (i.e., nongrafted defect), a small amount of new bone tissue was formed only at the edge of the tibial defect, and fibrous connective tissue and bone marrow nearly filled the entire tibial defect after 4 weeks of surgery.  Notably, in the case of the Bontree® 1.0-granules, thick and dense new bone tissues were consistently

generated between the graft granules, while relatively thin and less dense new bone tissues were formed with Bontree® 0.5-granules. Quantitative analysis of the new bone region (Figure 3B) showed 21.6% ± 5.3% and 26.0% ± 8.2% of new bone volumes after 4 weeks of grafting Bontree® 0.5- and 1.0-granules, respectively, which were significantly higher than that of the control (11.2% ± 2.9%) (*p* < 0.01).

**Figure 3.** (**A**) Representative histological images with low- and high-magnifications of the bone graft implantation regions in a rabbit tibia 4 and 12 weeks after surgery. Magnified regions are marked with yellow rectangles in the low magnification images. Black and white bars indicate 2 mm and 1 mm, respectively. Percentages of (**B**) new bone and (**C**) bone graft volumes compared with the tibial defect volume at 4 and 12 weeks after surgery. Statistical significance was set at \* *p* < 0.05 and \*\* *p* < 0.01, respectively.

The histological differences between the samples decreased with increasing implantation time. Twelve weeks after surgery (Figure 3A), all samples exhibited complete sealing

of the bone defects; the tibial defects were fully filled with newly formed compact bone, and there was no open gap at the side of the cortical bone defect regions. However, in the control and Bontree® 0.5-granule groups, regenerated bone tissues were not fully dense, and they showed macroscopic inner pores within the bone structures. In contrast, tibias grafted with Bontree® 1.0-granule showed a fully dense regenerated bone structure at the wound site, and they showed good histomorphological continuity with adjacent old bone tissues. From the quantitative results (Figure 3B), the Bontree® 1.0-granule group showed significantly higher new bone volumes (34.5% ± 2.5%) than the control (25.8% ± 1.5%) and Bontree® 0.5-granule (29.2% ± 4.8%) groups (*p* < 0.05).

Regarding bone graft degradation, Bontree® 0.5-granule and Bontree® 1.0-granule showed apparent degradation over 12 weeks (Figure 3C). Four weeks after surgery, large amounts of bone graft remained inside the tibial defects in both groups, and the retention rates were 5.8% ± 3.2% and 21.5% ± 8.0% for the Bontree® 0.5-granule and Bontree® 1.0-granule groups, respectively. In contrast, at the end of the in vivo test (week 12), the bone grafts were significantly degraded, showing only several coarse remnants in the marrow cavities in both groups. The retention rates were 2.7% ± 3.0% and 8.1% ± 6.4%, respectively, which were 2.2 and 2.7 times lower than those at week 4, respectively.

#### *3.2. Clinical Case Study*

#### 3.2.1. Description of Three Cases

All surgical procedures were performed by a single skilled periodontologist (W.-P.L.). Gargling was performed for 1 min using a 0.12% chlorhexidine solution before each surgery. After local anesthesia with 2% lidocaine containing epinephrine (1:100,000), bone was exposed by elevating a full-thickness flap after an alveolar ridge incision and a vertical incision with a #15 surgical blade. After curettage of the inflamed tissue, decortication was performed using a #330 carbide burr for bleeding. To facilitate the sticky bone substitute material condition, Bontree® and the entire blood harvested from the surgical site were mixed.

After each surgery, antibiotics (Augmentin 625 mg, Ilsung Pharm. Co., Seoul, Korea) and analgesics (aceclofenac 100 mg, Dona-A ST, Seoul, Korea) were administered orally for 7 days. In addition, the patients were advised to rinse the oral cavity with 0.12% chlorhexidine twice a day for 2 weeks. Complete stitch-out was performed 2 weeks after surgery. The implant stability quotient (ISQ) of each implant was measured at the second stage of implant surgery approximately 4 months after implantation.

The characteristic surgical procedure for each case is described below.

#### Case 1

A 77 year old man visited the clinic for implant placement in the 44–46 region. After local anesthesia, when implants (TS III SA®, Osstem, Seoul, Korea) were inserted, one implant thread was partially exposed buccally. Therefore, the guided bone regeneration (GBR) procedure was performed on the peri-implant dehiscence defect using a mixture of Bontree® and whole blood, followed by adaptation of a collagen barrier membrane (Ossix Plus®, Datum Dental Biotech, Lod, Israel). During the second-stage implant surgery performed 4 months after implantation, sufficient horizontal bone augmentation was observed around the implants. The final prosthesis was inserted approximately 6 months after implantation (Figure 4).

#### Case 2

A 69 year old man visited the clinic for bone augmentation and implant placement in the 24–27 region. As a severely atrophic alveolar ridge was expected, vertical ridge augmentation was performed first using Bontree® mixed with whole blood and a titanium mesh (Jeil Medical, Seoul, Korea) prefabricated on a 3D-printed model. The first stage of implant surgery for 24, 26, and 27 was performed 6 months after ridge augmentation. High primary stability was achieved at the time of implant (Luna®, Shinhung Co., Seoul, Korea) placement, and sufficient buccal and lingual marginal bone width was confirmed. Four months after implantation, vestibular loss and a lack of buccal attached mucosa were observed with vertical bone loss. Therefore, an apically positioned flap was performed simultaneously with the second-stage implant surgery, followed by the application of an absorbable periodontal dressing (Reso-Pac®, Hager & Werken GmbH & Co., KG, Germany). The final prosthesis was provided approximately 6 months after implantation (Figure 5).

**Figure 4.** Clinical case 1. (**A**) Clinical view before bone augmentation and implantation of 44, 45, and 46. (**B**) Peri-implant dehiscence defects at 44 and 45. (**C**) Bone augmentation at 44 and 45 using Bontree® and collagen barrier membrane. (**D**) Radiograph after bone augmentation and implantation. (**E**) Clinical view after flap elevation during the second-stage implant surgery 4 months after implantation. (**F**) Radiograph after the second-stage implant surgery. (**G**,**H**) Clinical and radiographic views of the final prosthesis 6 months after implantation. (**I**) Radiograph at 1 year after loading.

#### Case 3

A 63 year old man visited the clinic for sinus augmentation and implant placement in the 16–17 region. Because oroantral communication (OAC) was expected on radiographic evaluation, vertical ridge augmentation with simultaneous sinus floor elevation was planned. After a piezoelectric lateral bony window osteotomy and sinus membrane elevation, sinus grafting was performed with a mixture of Bontree® and whole blood, followed by bony window replacement. Next, vertical ridge augmentation was performed with Bontree® mixed with whole blood and a titanium-reinforced d-PTFE membrane (Cytoplast® Ti-250, Osteogenics Biomedical, Lubbock, TX, USA). A single-stage implant surgery for 16 and 17 was planned 6 months postoperatively, because the radiograph showed sufficient hard tissue volume in the 16–17 region, indicating the resolution of OAC. A core biopsy was conducted before drilling at site 16 for implant placement. The biopsy was harvested through the alveolar process at a depth of 10 mm using a trephine bur with an inner diameter of 2 mm. High primary stability was obtained when the implants (Superline®, Dentium Co. Ltd., Seoul, Korea) for 16 and 17 were placed, and sufficient buccal and lingual marginal bone width was also confirmed. The harvested specimens were fixed using paraformaldehyde in 4% buffered saline, followed by demineralization. The specimens were processed into paraffin blocks, and a microtome was used for microsectioning. Next, hematoxylin and eosin staining was performed. Four months after

implantation, modified periosteal fenestration [21–23], which we first suggested as a free gingival graft alternative, followed by the application of an absorbable periodontal dressing, was performed because of the loss of attached mucosa buccally. The final prosthesis was inserted 6 months after the single-stage implant surgery (Figure 6).

**Figure 5.** Clinical case 2. (**A**,**B**) Clinical and radiographic views before vertical ridge augmentation in the 24, 26, and 27 regions. (**C**) Three-dimensional printed model of the severely atrophic posterior maxilla. (**D**) A prefabricated titanium mesh on the 3D-printed model before clinical application. (**E**) Buccal view after flap elevation. (**F**) Bontree® mixed with whole blood. A sticky bone graft with good manipulability is observed. (**G**) Buccal view of vertical ridge augmentation using Bontree® and a titanium mesh. (**H**) Radiograph after vertical ridge augmentation. (**I**) Occlusal view of 24, 26, and 27 implants placed 6 months after vertical ridge augmentation. (**J**) Radiograph after the first-stage implant surgery. (**K**) Occlusal view of the second-stage implant surgery using apically positioned flap and punching technique 4 months after implantation. (**L**) Radiograph after the second-stage implant surgery. (**M**,**N**) Clinical and radiographic views of the final prosthesis 6 months after implantation. (**O**) Radiograph at 1 year after loading.

**Figure 6.** Clinical case 3. (**A**,**B**) Clinical and radiographic views before alveolar ridge and sinus augmentation in the 16 and 17 regions. (**C**) Buccal view after a piezoelectric bony window osteotomy and sinus membrane elevation. An oroantral communication is seen. (**D**) Sinus graft with Bontree®. (**E**) Vertical ridge augmentation with Bontree® and titanium-reinforced d-PTFE membrane after the replacement of the bony window. (**F**) Radiograph after sinus and ridge augmentation. (**G**) Buccal view before placement of implants 16 and 17, 6 months after sinus and ridge augmentation. A coronally positioned mucogingival junction line is observed. (**H**) Occlusal view of implants 16 and 17 immediately after placement. (**I**) The core biopsy in the 16 region using a trephine bur. (**J**) Radiograph after the single-stage implant surgery. (**K**) Histological examination. New bone formation (NB) is observed with no inflammatory tissue (H&E-stained; original magnification × 100). (**L**) Buccal view after performing a modified periosteal fenestration 4 months after the single-stage implant surgery. (**M**,**N**) Clinical and radiographic views of the final prosthesis 6 months after the single-stage implant surgery. (**O**) Radiograph at 1 year after loading.

#### 3.2.2. Clinical and Radiological Findings

A total of eight implants were placed in three patients who underwent sinus or alveolar ridge bone grafting with Bontree®. None of the patients had postoperative complications other than slight swelling at the surgical site. At 4 months after implantation, the ISQ values were >60 for all implants, indicating good implant stability (Table 1). All eight implants in all three patients were followed up for at least 12 months after functional loading, and the success rate, which was evaluated on the basis of the International

Congress of Oral Implantologists Pisa Consensus implant health scale [24], was 100% (Table 1). The radiograph at the 1 year follow-up showed integration of the implant with the regenerated bone and no bone loss or peri-implant radiolucency. No decrease in graft height was observed on any radiograph, and healthy peri-implant mucosa was established around all implants during the 1 year loading period.

#### 3.2.3. Histological Findings

Histological analysis at the site of 16 in case 3, where sinus and ridge grafting was performed with Bontree®, revealed the deposition of newly formed bone around the residual synthetic bone graft and satisfactory incorporation of the newly formed bone with the residual synthetic bone graft. No foreign body reactions or inflammatory signs were detected (Figure 6K).

#### **4. Discussion**

In this study, the safety and efficacy of Bontree® for bone regeneration were demonstrated using in vitro and in vivo experiments. In addition, in patients with bony defects in the alveolar ridge or maxillary sinus, oral rehabilitation using dental implants was possible after ridge or sinus augmentation using Bontree®.

Although OCP is a promising bone graft material with excellent biological properties, there are only a few commercially available OCP-based products, such as Ti-Oss® (Chiyewon, Guri, Korea) or BonarcTM (Toyobo Co. Ltd., Osaka, Japan) bone grafts. However, even these products are not fully composed of Ca- and P-containing apatites; rather, they consist of a combination of OCP with naturally occurring bovine bone materials or collagens. In contrast, the recently introduced OCP bone graft, Bontree®, is fully composed of apatite compounds with 80 wt.% OCP and 20 wt.% HA, and it exhibits a commercially acceptable quality, defect-free homogeneous surface morphology, and relatively uniform size distribution (Figure 1A). In addition, numerous micron-scale crystals covering the entire surface of Bontree® granules are beneficial for increasing the rate of dissolution and resorption of Ca and P ions under physiological conditions, thereby enabling active mineralization on their surfaces.

Short-term in vitro and long-term in vivo studies were performed to investigate the bone formation ability of Bontree®. In the in vitro experiments, RANKL and ALP were used as key markers for bone formation. RANKL is a transmembrane protein that controls the differentiation, maturation, and activation of osteoclasts, which results in bone resorption during remodeling, whereas ALP is an enzyme secreted by osteoblasts and is involved in the mineralization and calcification of newly formed bone tissues. Therefore, the suppressed RANKL activity of MG63 cells is closely associated with reduced osteoclastic differentiation and bone resorption, while the improved ALP activity indicates better osteoblastic differentiation and new bone formation on the bone grafts [19,25,26]. If the bone graft material has insufficient ability to regenerate bone tissues, pores between the bone grafts allow infiltration of soft tissues that hinder the formation of new bone [27,28]. Moreover, the soft tissue is unable to provide sufficient mechanical support to the host bone and prevents the escape of bone grafts from the implanted site [29]. As shown in Figure 2A,B, Bontree® exhibited significantly higher values of ALP activity than MBCP+TM at 12 days of MG63 cell culturing, and a significant amount of newly formed bone tissues appeared uniformly distributed over the entire region of the Bontree® 0.5-granule- and Bontree® 1.0-granule-grafted tibial defects at 4 weeks after surgery. In previous studies, OCP showed outstanding osteoconductive and osteoinductive properties owing to its unique physicochemical properties [30]. When degraded and converted to HA, it releases numerous inorganic PO4 3− and Ca2+ ions, which are known to promote osteoblastic differentiation and maturation better than HA, as evidenced by the increased ALP activity (Figure 2B) [31]. In addition, OCP is beneficial for the adsorption of bone-forming molecules and proteins, which could further increase the osteogenic differentiation of cells and initiate bone regeneration [32].

In terms of the size effect of bone grafts on bone regeneration, macroscopic bone grafts generally provide more free space than their microscopic counterparts for cells to move into and attach to, thereby promoting vascularization and ingrowth of newly formed bone tissue into the bone grafts [33]. As shown in Figure 3A,B, Bontree® 1.0-granule was associated with greater amounts of newly formed bone tissues than the Bontree® 0.5-granule at both 4 and 12 weeks. According to the results of in vitro and in vivo experiments, the Bontree® 1.0-granule has a high potential as a commercially applicable novel bone graft material, and it was selected for the evaluation of clinical significance.

On the basis of the results of this basic research, we attempted to perform alveolar ridge or sinus augmentation using Bontree® 1.0-granule during dental implant surgery. Eight implants in three patients showed a 100% success rate after 1 year of functional loading. During the surgical procedure, clinical and systemic complications, such as pain, severe inflammation, or significant infection, were not observed. In Case 1, GBR was performed on the peri-implant dehiscence bony defects using Bontree® and an absorbable collagen barrier membrane, and clinically successful horizontal hard tissue gain was observed after 4 months. In Cases 2 and 3, vertical ridge or sinus augmentation with Bontree® was performed 6 months before implant placement, sufficient horizontal/vertical bone width and height were obtained clinically, and sufficient implant primary stability ≥ 40 N was achieved. All ISQ values measured approximately 4 months after implant placement were >60, indicating successful secondary implant stability. In addition, no characteristic marginal bone loss around the implant was observed on radiographs for up to 1 year after loading. The clinically and radiographically successful bone regeneration with Bontree® can be attributed to the excellent new bone formation ability unique to OCP. In other words, OCP seems to provide a nucleus for promoting osteogenesis and creates numerous starting points for ossification, unlike HA and β-TCP [34]. In addition, Kawai et al. [14] reported that OCP-based bone grafts achieved comparable osseointegration to autologous bone grafts in a study of the mandible in a canine model. However, due to the limitations inherent to retrospective studies, the number of samples was too small, and we cannot confirm the clinical characteristics of Bontree® only on the basis of the three cases included in this study. In addition, the follow-up period of approximately 1 year may be too short to evaluate the success rate of the implants. Therefore, as a followup study, comparative studies with other bone substitute materials, long-term studies, and prospective studies using Bontree® in a much larger number of cases are needed.

#### **5. Conclusions**

According to our basic research and clinical applications, Bontree® showed significantly higher ALP activity than a commercial biphasic calcium phosphate ceramic (MBCP+TM). Although both Bontree® 0.5-granule and Bontree® 1.0-granule are mainly composed of OCP distributed uniformly regardless of size, the Bontree® 1.0-granule has a higher potential for new bone formation. The clinical cases had predictable and successful outcomes, which demonstrated the safety and efficacy of Bontree® in alveolar ridge or sinus augmentation.

**Author Contributions:** Conceptualization, W.-P.L.; methodology, T.-S.J. and W.-P.L.; validation, S.-Y.K.; formal analysis, T.-S.J. and W.-P.L.; investigation, J.-S.K.; resources, J.-S.K. and W.-P.L.; data curation, T.-S.J. and J.-S.K.; writing—original draft preparation, J.-S.K., T.-S.J. and W.-P.L.; writing—review and editing, S.-Y.K. and W.-P.L.; visualization, T.-S.J. and W.-P.L.; supervision, S.-Y.K.; project administration, W.-P.L. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Institutional Review Board Statement:** This study was conducted in accordance with the guidelines of the Declaration of Helsinki and approved by the Institutional Review Board of the Dental Hospital of Chosun University in Gwangju, Korea (CUDHIRB-2103-008).

**Informed Consent statement:** Informed consent was obtained from all patients involved in the study.

**Data Availability Statement:** The datasets generated or analyzed during the current study are available from the corresponding author upon reasonable request.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

## *Article* **Titanium Nitride Plating Reduces Nickel Ion Release from Orthodontic Wire**

**Arata Ito, Hideki Kitaura \*, Haruki Sugisawa, Takahiro Noguchi, Fumitoshi Ohori and Itaru Mizoguchi**

Division of Orthodontics and Dentofacial Orthopedics, Graduate School of Dentistry, Tohoku University, 4-1 Seiryo-machi, Aoba-ku, Sendai 980-8575, Japan<EMAIL_ADDRESS>(A.I<EMAIL_ADDRESS>(H.S<EMAIL_ADDRESS>(T.N<EMAIL_ADDRESS>(F.O<EMAIL_ADDRESS>(I.M.)

**\*** Correspondence<EMAIL_ADDRESS>Tel.: +81-22-717-8374

**Abstract:** The leaching of metal ions from orthodontic appliances is a problem for their use in patients with metal allergies. Despite the development of a number of non-metal orthodontic appliances, including brackets, non-metal wires are not yet available. Therefore, it is necessary to modify the surfaces of orthodontic wires to prevent the leaching of metal ions into the oral environment for use in such patients. This study was performed to examine whether plating of orthodontic wire with titanium nitride (TiN), which does not impair its mechanical properties, could prevent the leaching of metal ions from the wire on immersion in acid. To investigate the acid corrosion resistance of the wire, the amount of metal ions eluted from the wire immersed in acid was measured by using inductively coupled plasma mass spectrometry (ICP-MS) and the dimethylglyoxime (DMG) test, the properties of the wire surface were examined by stereomicroscopy and scanning electron microscopy, and the surface roughness was measured using a surface roughness tester. The results indicated that TiN plating of orthodontic wire significantly suppressed the elution of metal ions on immersion in acid.

**Keywords:** TiN ion plating; metal allergy; orthodontic wire

#### **1. Introduction**

The number of patients with metal allergies has increased in recent years [1,2]. The metals used in dentistry are considered to be among the factors responsible for metal allergies. Many orthodontic materials also contain such metals associated with allergic reactions, and therefore care is required in the orthodontic treatment of patients with metal allergies [3–5]. Some metals commonly used in dental practice, including nickel (Ni), cobalt (Co), and chromium (Cr), are known allergens for patients with metal-allergic disease [6,7]. These metals are also present in various orthodontic materials [8]. Several studies have shown that Ni is the most common allergen associated with metal-allergic disease [2,9,10]. Previously, we evaluated the surface elements of orthodontic materials by X-ray fluorescence spectroscopy, and found that Ni was present in almost all the orthodontic metal materials [8]. Metal ions can be eluted from orthodontic materials due to the electrochemical corrosion caused by saliva, electrolytes in food debris, and acids produced by bacteria [11]. Therefore, orthodontic patients can be exposed to Ni from orthodontic appliances. In fact, Fors et al. [12] reported significantly elevated Ni levels in the dental plaque of patients with orthodontic appliances than in that of nonorthodontic patients.

Despite the development of many non-metal orthodontic appliances, including brackets, non-metal wires are still not available [5]. Therefore, it is necessary to improve orthodontic wires to achieve high corrosion resistance. Plating can be used to provide increased surface hardness, wear resistance, corrosion resistance, low friction, and improved esthetics, as well as to improve interactions with biological and material substrates [13]. Physical

**Citation:** Ito, A.; Kitaura, H.; Sugisawa, H.; Noguchi, T.; Ohori, F.; Mizoguchi, I. Titanium Nitride Plating Reduces Nickel Ion Release from Orthodontic Wire. *Appl. Sci.* **2021**, *11*, 9745. https://doi.org/ 10.3390/app11209745

Academic Editor: Vittorio Checchi

Received: 6 September 2021 Accepted: 14 October 2021 Published: 19 October 2021

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2021 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

vapor deposition (PVD) ion plating is a particularly suitable plating method to produce thin films on substrates. TiN layers fabricated by ion plating have demonstrated favorable characteristics in terms of corrosion resistance [14], wear resistance [15], hardness [16], and low friction [17], and may therefore be useful for the modification of the surfaces of orthodontic wires.

We described the mechanical and frictional properties of TiN-coated wires previously [17], but the details of corrosion resistance have yet to be elucidated. We focused on two acids as potential sources of metal ion leaching from orthodontic wires in the oral cavity. First, as fixed orthodontic appliances induce the retention of bacterial plaque, leading to an increase in levels of caries-inducing bacteria in the oral cavity, such as *Lactobacillus* spp. and *Streptococcus mutans* [18], and bacteria corrode orthodontic appliances [19], we focused on lactic acid, which is the main metabolite that causes low pH [20]. We also focused on gastric acid, mainly hydrochloric acid (HCl), which enters the oral cavity due to gastroesophageal reflux disease (GERD), which has a high incidence, of 10–20%, in adults [21]. In addition to acid erosion of teeth and dissolution of dental restorations [22], gastric acid has been reported to be involved in the elution of metal ions from orthodontic wires [23].

In the present study, we evaluated the corrosion resistance of TiN-plated stainless steel (SS) and nickel titanium (NiTi) orthodontic arch wire immersed in acids by measuring the concentrations of released metal ions, by stereomicroscopic and scanning electron microscopic observation of the wire surface, and by surface roughness measurement.

#### **2. Materials and Methods**

#### *2.1. TiN Ion Plating*

SS wire and NiTi wire (0.016 × 0.022 inches) (American Orthodontics, Sheboygan, WI, USA) were plated with TiN using the hollow cathode discharge method, as described previously [17]. Briefly, a 180-mm length of wire was wiped clean with ethanol and coated using an ion plating system (HCD Ion Plating System X-27; Tigold Co., Ltd., Chiba, Japan) with 1.49 Pa argon (Ar) gas, 1.69 Pa nitrogen (N2) gas, and a bias voltage of −20 V. The temperature of the substrate was kept at approximately 220 ◦C for 7 min, and after the wire was completely covered with the TiN layer, except at the two ends, it was slowly cooled under vacuum for 60 min. The coating thickness was set to 0.3 μm. The thickness of the TiN coating was measured as described previously [17].

#### *2.2. Metal Ion Release Testing*

Orthodontic wire samples consisting of uncoated SS wire, TiN-coated SS wire, uncoated NiTi wire, and TiN-coated NiTi wire were separately immersed in plastic dishes (untreated 60 × 15 mm dishes; Falcon, Tewksbury, MA, USA) containing 10 mL of physiological saline, sterile water, 35% HCl (pH −1.1) (Wako Pure Chemical Industries, Ltd., Osaka, Japan), and 88% lactic acid (pH 1.2) (Wako Pure Chemical Industries, Ltd., Osaka, Japan). The dishes were placed in an incubator at 37 ◦C for 30 min, and the concentrations of Cr, manganese (Mn), iron (Fe), and Ni ions released from the SS wires and of the Ni and Ti ions released from the Ni-Ti wires into the solutions were measured by triple quadrupole inductively coupled plasma mass spectrometry (ICP-MS) (ICP-QQQ-Agilent 8800; Agilent Technologies, Santa Clara, CA, USA). The concentrations of metal ions are shown as the mean ± SD of four replicates for each sample.

#### *2.3. DMG Tests*

The dimethylglyoxime (DMG) test was performed to detect the leaching of Ni from the surface of orthodontic wires immersed in physiological saline, sterile water, 35% HCl, and 88% lactic acid at 37 ◦C for 30 min. As described previously [24], two drops of 1% DMG in absolute ethanol and two drops of 10% ammonium hydroxide were added to a cotton swab, which was then rubbed on the sample for 30 s. The strength of the color reaction on the cotton swab was then evaluated visually; the development of a red color was considered a positive reaction, indicating the release of Ni from the wire surface. For positive reactions, points were allocated according to the following criteria: dark red, 3 points; red, 2 points; pink, 1 point; and colorless, 0 points. All the tests were performed in triplicate by the same operator. The values for all the tests are shown as the mean ± SD of three replicates for each sample.

#### *2.4. Observation of the Wire Surface*

The wire surface was then observed and photographed under a stereomicroscope (M165FC; Leica, Wetzlar, Germany) and by scanning electron microscopy (SEM) (VE-7800; Keyence, Osaka, Japan) at a magnification of 200× and 1000× with an acceleration voltage of 15 kV.

#### *2.5. Measurement of Surface Roughness*

Surface roughness was assessed using a surface roughness tester (Surftest SJ-210; Mitutoyo Corporation, Tokyo, Japan), according to ISO 1997, with a diamond tip radius of 5 μm, a measurement force of 4.0 mN, a scanning speed of 0.5 mm/s, a cut-off length of 0.8 mm, and a Gaussian filter. The samples were each fixed under the stylus of the tester, and the mean surface roughness values (Ra, μm) were determined. The values for each sample are shown as the mean ± SD of three replicates for each sample.

#### *2.6. Statistics*

Statistical analyses were performed with JMP® Pro software (version 16.0.0, SAS Institute Inc., Cary, NC, USA). The data obtained from the DMG test and the surface roughness test were analyzed using the Tukey–Kramer HSD test. The Shapiro–Wilk test for normality indicated that the ICP-MS data had a non-parametric distribution, so the concentrations of the eluted metal ions were analyzed using non-parametric comparisons for each pair with Wilcoxon's method. In all the analyses, *p* < 0.05 was taken to indicate statistical significance.

#### **3. Results**

#### *3.1. TiN Coating of Orthodontic Wires Suppresses the Elution of Metal Ions from the Wire Surface*

The amount of metal ions released from the TiN-coated and uncoated wires after immersion in sterile water, physiological saline, HCl, or lactic acid for 30 min, are shown in Tables 1 and 2. All the data are shown as the mean ± SD of four replicates.

Following immersion in HCl for 30 min, the TiN-coated SS wire showed significantly reduced Cr, Mn, Fe, and Ni ion concentrations compared to the uncoated SS wire and the HCl-only immersion solution, which was used as a control (Figure 1a). On the other hand, following the immersion of the NiTi wire in HCl for 30 min, the concentrations of Ni and Ti ions were significantly reduced in the TiN-coated wire sample compared to the uncoated NiTi wire sample and the HCl-immersion-solution-only control (Figure 1b). Furthermore, the leaching of Ni ions from the uncoated SS and NiTi wire immersed in lactic acid for 30 min was significantly increased compared to the lactic-acid-only immersion solution control (Figure 1c,d). There were no significant differences in the amounts of metal ions released when the wires were immersed in sterile water or physiological saline for 30 min between the TiN-coated wire, the uncoated wire, and the immersion-solution-only control.

**Figure 1.** Metal ions released from orthodontic arch wire immersed in different solutions. After immersion in HCl or lactic acid solution for 30 min, the amount of metal ions eluted from the wire was determined by ICP-MS. The concentrations of the eluted metal ions from TiN-coated SS wire and uncoated SS wire were: (**a**) SS immersed in HCl; (**b**) NiTi immersed in HCl; (**c**) SS immersed in lactic acid; (**d**) NiTi immersed in lactic acid. The statistical significance of differences was determined by the Shapiro-Wilk (\* *p* < 0.05, \*\* *p* < 0.01).

#### *3.2. DMG Test*

The results of the DMG test were scored according to the degree of red color development as an indicator of the strength of the positive response, as shown in Figure 2a. The DMG test results for Ni release are shown in Tables 3 and 4. The TiN-coated and uncoated SS and NiTi wire immersed in HCl for 30 min showed positive reactions on the DMG test, while negative reactions were observed for the other conditions. Following immersion in HCl for 30 min, the scores for the TiN-coated and uncoated SS wire were 1.0 ± 0.0 and 2.67 ± 0.58, respectively (Figure 2b, Table 3), and those for the TiN-coated and uncoated NiTi wire were 1.0 ± 0.0 and 2.67 ± 0.58, respectively (Figure 2c, Table 4).

**Figure 2.** DMG test: (**a**) Positive reactions were scored; (**b**) SS wire immersed in HCl; (**c**) NiTi wire immersed in HCl. The statistical significance of differences was determined by the Tukey-Kramer HSD test. (\* *p* < 0.05, \*\* *p* < 0.01).

**Table 3.** DMG test results for SS wire.

**Table 4.** DMG test results for NiTi wire.

#### *3.3. Observation of the Wire Surface*

Figure 3 shows stereomicrographs of the surfaces of the wire samples. Corrosion of the surface was found on both the TiN-coated and uncoated SS and NiTi wire after immersion in HCl for 30 min. There were no differences between the non-immersed wire and wire immersed in lactic acid for 30 min.

**Figure 3.** Stereomicrographs of orthodontic wire immersed in HCl or lactic acid for 30 min: (**a**) TiN-coated SS wire; (**b**) Uncoated SS wire; (**c**) TiN-coated NiTi wire; (**d**) Uncoated NiTi wire.

SEM images of the wire surface are shown in Figure 4. At a magnification of 200×, corrosion was found on both the TiN-coated and uncoated SS and NiTi wires after immersion in HCl for 30 min. At a magnification of 1000×, the TiN-coated SS and NiTi wires demonstrated pitting corrosion, while the uncoated SS and NiTi wires showed uniform corrosion. There were no differences in surfaces between non-immersed wires or TiN-coated and uncoated wire immersed in lactic acid for 30 min.

**Figure 4.** Scanning electron micrographs of orthodontic arch wire immersed in HCl or lactic acid for 30 min. (×200): (**a**) TiN-coated SS wire; (**b**) Uncoated SS wire; (**c**) TiN-coated NiTi wire; (**d**) Uncoated NiTi wire. (×1000): (**e**) TiN-coated SS wire; (**f**) Uncoated SS wire; (**g**) TiN-coated NiTi wire; (**h**) Uncoated NiTi wire.

#### *3.4. Surface Roughness*

Both the uncoated and the TiN-coated SS and NiTi wires immersed in HCl for 30 min showed significantly increased surface roughness compared to the non-immersed wire (Figure 5). On the other hand, immersion in lactic acid for 30 min was associated with significantly increased surface roughness compared to non-immersed wire only for the uncoated SS and NiTi specimens.

**Figure 5.** Comparison of surface roughness of orthodontic arch wires after immersion in HCl or lactic acid. After immersion in HCl or lactic acid for 30 min, the surface roughness (Ra) of orthodontic arch wires was assessed using a surface roughness tester: (**a**) SS wire immersed in HCl; (**b**) NiTi wire immersed in HCl; (**c**) SS wire immersed in lactic acid; (**d**) NiTi immersed in lactic acid. The statistical of differences was determined by the Tukey-Kramer HSD test. (\* *p* < 0.05, \*\* *p* < 0.01).

#### **4. Discussion**

As the release of metal ions from orthodontic materials in the oral environment is closely related to their corrosion, metal surfaces are coated to improve corrosion resistance [13]. A number of plating techniques and materials are available for coating such metal components, and TiN ion plating was used in this study. The Ti of TiN components forms a passivation film, which improves the surface properties of the substrate material [25,26]. TiN ion plating results in little degradation of the mechanical properties of the substrate material because it is performed at low temperatures. Previously, we reported

that TiN-coated orthodontic wire has a reduced friction coefficient, and the mechanical properties of the base material are maintained [17]. This study was performed to investigate whether the elution of metal ions can be suppressed by coating orthodontic wires with TiN film.

The ICP-MS data showed that the elution of metal ions was significantly suppressed in the TiN-coated wire compared to the uncoated wire after immersion in HCl for 30 min. Similarly, the DMG test confirmed that that the amount of Ni leached from both the SS and the NiTi wire after immersion in HCl were significantly lower for the TiN-coated than for the uncoated wire. The leaching of metal ions by HCl was suppressed by TiN coating. On the other hand, with immersion in lactic acid, Ni ions were leached from the uncoated but not the TiN-coated SS and NiTi wire, suggesting that the TiN-coated SS and NiTi wire was resistant to corrosion by lactic acid. Wire immersed in lactic acid showed no positive reaction on the DMG test, which may have been because the concentration of Ni was below the limit of detection of the test (10 μg/mL, 10 ppm) [27]. ICP-MS also showed that the amount of Ni ions eluted from the wire immersed in lactic acid was < 10 μg/mL, suggesting that it would be difficult to detect Ni elution associated with immersion in lactic acid on the DMG test. The amount of metal ions eluted from the TiN-coated SS and NiTi wire and the variations in the data were considered to be greater because of the high concentration of HCl used in this experiment. As the HCl derived from GERD has a lower concentration than that used in this study [23] and the immersion time in GERD-derived HCl is likely to be < 30 min in the oral cavity, the amount of elution is expected to be much smaller than in this experiment. However, as the threshold for Ni allergies is 1.5 μg [28], and there is a risk of onset of metal allergy when exposed to very small amounts of the metal, further experiments are required to examine the use of HCl at a concentration closer to that likely to be present in the oral cavity. Furthermore, the amount of Ni ion elution from uncoated SS wire immersed in lactic acid showed a high standard deviation. The immersion time of 30 min in lactic acid is considered to be the starting point of erosion for the SS wire used in this experiment. In some samples, the amount of Ni ion elution was small because dissolution had only just begun, and this may have resulted in the high degree of variation in the data.

The stereomicroscopy and SEM analyses showed the corrosion of all the wires following immersion in HCl, but the types of corrosion differed between the TiN-coated and uncoated wires. That is, pitting corrosion occurred in the TiN-coated SS and NiTi wire, while the uncoated SS and NiTi wire showed uniform corrosion. Such uniform corrosion is known to occur when metal is immersed in HCl [29], and this was confirmed in the present study. Pitting corrosion tends to occur in TiN with a coating thickness <6 μm [30]. In the present study, the TiN coating was 0.3 μm thick, which was intended to be used for orthodontic treatment, so the pitting corrosion may have occurred easily. Preliminary experiments showed that the peeling of TiN coating gradually became noticeable under the stereomicroscope when immersed in HCl for more than 30 min, so the immersion time was set to 30 min in this study.

The surface roughness of both the uncoated and the TiN-coated wire immersed in HCl was increased significantly compared to that of the non-immersed samples, suggesting that the wire surface was corroded by HCl, confirming the results of the microscopic analyses. Furthermore, the surface roughness of the TiN-coated wire was greater than that of the uncoated wire despite the reduced amount of metal ion elution. This may have been due to the difference in corrosion patterns between the TiN-coated and uncoated wires, as revealed by the stereomicroscopy and SEM analyses. The uniform corrosion of the uncoated wire was seen over its entire surface, but the pitting corrosion of TiN-coated wire induced by HCl in this study was randomly scattered, and the SEM images showed small fragments of coating material that seemed to have been peeled off around the pores, which may have caused an increase in the surface roughness. The surface roughness (Ra) measured in this study is a two-dimensional measurement, so a three-dimensional measurement of surface roughness (Sa) may be useful for further detailed analyses [31]. Microscopic examination indicated that the surface condition of the TiN-coated SS and NiTi wire was not altered by immersion in lactic acid and there was no change in the surface roughness, suggesting no corrosion by lactic acid. On the other hand, for the uncoated SS and NiTi wire, although microscopy revealed no changes in the surface condition, the surface roughness was increased by immersion in lactic acid, suggesting the possibility of corrosion by lactic acid.

Taken together, the results outlined above indicated that pitting corrosion occurred on the surface of TiN-coated SS and NiTi wire immersed in HCl, but that the corrosion did not spread over the entire surface of the wire, so the leaching of metal ions was significantly suppressed compared to the uncoated wire. The leaching of Ni ions from the TiN-coated SS and NiTi wire immersed in lactic acid was also shown to be significantly suppressed compared to the uncoated wire. These results indicated that TiN coating inhibits acid-induced corrosion and the elution of metal ions. GERD is associated with the regurgitation of gastric acid, mainly HCl, into the oral cavity. Thus, the results presented above suggest that the use of TiN-coated wires may be effective in preventing the leaching of metal ions from orthodontic appliances by HCl in gastric acid and reduce the rates of induction and exacerbation of metal allergy in patients with GERD. In addition, *Lactobacillus acidophilus* present in the oral cavity has been suggested to be related to the corrosion of metals mediated by lactic acid production [11]. The TiN-coated wire used in this study was resistant to corrosion by lactic acid, and is expected to be applicable to orthodontic treatment of patients with metal allergies.

We analyzed the acid corrosion resistance of TiN-coated wires in vitro. The following issues should be considered in future experiments. First, it is necessary to examine how long the TiN coating can be maintained in good condition in the oral cavity during orthodontic treatment. Second, the TiN coating is expensive, so cost reduction must be considered for clinical use. We plan to reduce the cost by coating as many wires as possible at once. Although TiN coating has excellent biocompatibility, as demonstrated previously in a device for atrial septal defect closure [14], it will be necessary to study its biocompatibility in the oral cavity, where it will be exposed to various chemical substances in food and drink, as well as saliva.

#### **5. Conclusions**

TiN plating of orthodontic wires can inhibit acid-mediated corrosion, and thus reduce the elution of Ni ions from the wire surface. These results suggest that TiN plating may be useful for the orthodontic treatment of patients with metal allergies. In addition, TiN plating has the potential for use in other dental applications, such as coating the metal parts of removable dentures and for wires to splint the teeth after orthodontic treatment.

**Author Contributions:** Conceptualization, A.I. and H.K.; methodology, A.I. and H.K.; validation, A.I. and H.K.; formal analysis, A.I.; investigation, A.I. and H.S.; resources, A.I. and H.K.; data curation, A.I.; writing—original draft preparation, A.I.; writing—review and editing, H.K.; visualization, A.I., T.N. and F.O.; supervision, H.K. and I.M.; project administration, H.K.; funding acquisition, A.I. All authors have read and agreed to the published version of the manuscript.

**Funding:** This work was supported in part by JSPS KAKENHI grants from the Japan Society for the Promotion of Science (No. K20K187740).

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** The data and analysis in this study are available on request from the corresponding author.

**Conflicts of Interest:** The authors declare there are no conflicts of interest.

#### **Abbreviations**

#### **References**

## *Article* **Influence of Selected Restorative Materials on the Environmental pH: In Vitro Comparative Study**

**Anna Lehmann 1,†, Kacper Nijakowski 1,\*,†, Michalina Nowakowska 2, Patryk Wo´s 2, Maria Misiaszek 2 and Anna Surdacka 1**

**Abstract:** In dental caries treatment, it is worth using such restorative materials that may limit plaque accumulation. The pH of the filling seems to be an important factor affecting the potential bacterial colonisation. Our study aimed to assess how selected restorative materials influence the environmental pH. A total of 150 specimens (30 of each: Ketac Molar, Riva LC, Riva SC, Filtek Bulk Fill, and Evetric) were placed in 100 sterile hermetic polyethene containers with saline and stored in 37 ◦C. The pH of each sample was measured using the electrode Halo HI13302 (Hanna Instruments, Poland) at specific points in time for 15 days. The initial pH levels were significantly lower for glass ionomer cements (3.9–4.7) compared to composites (5.9–6.0). With time, the pH increased for samples with glass ionomer cements (by nearly 1.5), whereas it decreased for samples with composites (maximally by 0.8). In the end, all materials were in the pH range between 5.3 and 6.0. The highest final pH was obtained with Ketac Molar at about 5.9. Double samples had lower pH values than single samples, irrespective of the type of material. In conclusion, immediately after application, restorative materials decreased the environmental pH, especially light-cured glass ionomer cements. For glass ionomers, within two weeks, the pH increased to levels comparable with composites.

**Keywords:** dental material; restorative material; composite; glass ionomer; resin-modified glass ionomer; pH; acidity; oral environment; dental filling; dental restoration

#### **1. Introduction**

Acid demineralisation of enamel due to the metabolic activity of bacterial plaque is the first stage of the caries process, leading to the formation of a cavity [1–3]. The destructed dental tissues are restored with various materials; however, on the contact surface of the tooth and the filling, there is usually a niche that favours the plaque accumulation and may increase the risk of secondary caries. Therefore, it is worth using such restorative materials that may prevent bacterial adhesion or limit the development of bacterial plaque. The pH of the filling seems to be an essential factor influencing the potential bacterial colonisation [4]. Numerous in vitro studies show that at acidic pH, the surface of the composite material is more stable than the surface of glass ionomer cement [4–6]. Thus far, few researchers have analysed the potential effect of the filling materials on the environmental pH. However, conventional glass ionomer cements (GICs) can buffer lactic acid and release fluoride, which appears to be very beneficial in a clinical approach [7–9].

It is believed that there are no clear limits to the number of fillings a dentist can perform in one visit. If necessary, it is recommended to perform several fillings in the same area within the conduction anaesthesia. However, other factors can affect the amount

**Citation:** Lehmann, A.; Nijakowski, K.; Nowakowska, M.; Wo´s, P.; Misiaszek, M.; Surdacka, A. Influence of Selected Restorative Materials on the Environmental pH: In Vitro Comparative Study. *Appl. Sci.* **2021**, *11*, 11975. https://doi.org/10.3390/ app112411975

Academic Editor: Vittorio Checchi

Received: 24 November 2021 Accepted: 13 December 2021 Published: 16 December 2021

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2021 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

of time a patient spends in the dental chair. Most often, apart from the time limitations of the doctor and the patient, the general health condition and ailments related to the temporomandibular joint can pose a limitation. After two hours with an open mouth, it is recognised that there may be irreversible changes in the joint [10]. There are no previous articles on how the pH in the oral cavity changes after tooth restoration depending on the number of fillings.

The literature reports the hydrolysis process of fillings under low and high pH conditions in vitro. The carious bacteria fermentation and gastric acid attack are most often simulated by lactic acid and hydrochloric acid, respectively [11,12]. Both composite materials and glass ionomer cements undergo hydrolysis at acidic pH [4,13,14]. However, GICs seem to exhibit more favourable properties than resin-based materials, among others: chemical adhesion to enamel and dentin in the presence of moisture, resistance to microleakage, good marginal integrity, dimensional stability at high humidity, coefficient of thermal expansion like tooth structure, biocompatibility, fluoride release, rechargeability with fluoride, and less shrinkage than resins upon setting with no free monomer being released [15]. In the case of composite materials, the key is proper preparation of the cavity edge, gradual material application, and a meticulous polishing procedure [1,16]. However, even in vitro, a microleakage can be observed in the most carefully placed composite [17]. Certain limitations can also be seen in the bonding systems. All adhesive systems are somehow unstable and susceptible to hydrolytic degradation, which may be responsible for the partial failure of the material adhesion to the tooth, leading to marginal leakage [18,19]. The literature shows that resin restoration degradation is a complex process involving the hydrolysis of the resin and the dentin collagen fibril phases within the hybrid layer. Inhibition of the collagenolytic activity, as well as the use of cross-linking agents, are the two main strategies to increase the resistance of the hybrid layer to enzymatic degradation [20,21]. There is still no answer as to whether the effect of the restorations' pH on the oral environment could be a helpful factor in reducing secondary caries. It is known that despite the continuous modernising of the filling materials, the enamel still has a more significant buffer capacity than the composite or glass ionomer cement [22,23].

Our study aimed to analyse the potential pH effect of the most used dental materials, i.e., composites and glass ionomers, on the oral environment modelled in vitro. We tried to answer the following questions:

#### **2. Materials and Methods**

*2.1. Materials Used in the Study*

In this in vitro study, we used five dental restorative materials. Detailed characteristics of the materials are presented in Table 1.

**Table 1.** Detailed characteristics of the dental restorative materials used in the study.

**Table 1.** *Cont.*

#### *2.2. Specimen Preparation*

All specimen preparation was completed by one operator to reduce variability. A total of 150 specimens (30 from one material) were prepared using metal moulds with 6 mm diameter and 2 mm thickness. All materials were inserted into the mould and intentionally overfilled. Then, the mould was sandwiched between transparent Mylar strips to expel excess material. The uncured resin-modified glass ionomer and composites were light-cured for 40 s, according to the manufacturer's instructions (light intensity of 1200 mW/cm2, Translux Wave, Kulzer GmbH, Hanau, Germany). Conventional glass ionomer cement (CGIC) was prepared according to the powder-liquid mixing ratio indicated by the manufacturers. After 30 min, the CGIC specimens were removed from the moulds.

#### *2.3. Specimen Storage and pH Evaluation*

The samples were placed in 100 sterile hermetic polyethene containers (ApteoCare, Sanmed, Bydgoszcz, Poland; lot: 06/KS/2021/S) and incubated in 37 ◦C. Each type of material was divided into 20 containers. In the first 10, there was one sample of the filling material and in the next 10, two samples. To each container, 5 mL of saline (Polpharma, Stargard Gdanski, Poland; lot: 1280620) was added.

The pH of each sample was measured using the electrode Halo HI13302 (Hanna Instruments, Olsztyn, Poland) by immersing it into the central part of the solution. Between procedures, the electrode was cleaned and recalibrated. All measurements were made in duplicate and then averaged. The pH evaluation was performed at the following time points: after 1, 6, 12 and 24 h and after 3, 6, 9, 12 and 15 days from the specimen preparation.

#### *2.4. Statistical Analysis*

The determined pH values were analysed using a two-way repeated-measures analysis of variance (in the model with the interaction). The pH changes were compared separately at specific points in time depending on the kind of used dental restorative material and the number of the incubated samples. The significance level was defined as α = 0.05. The statistical analysis was performed using Statistica 13.3 (Statsoft, Cracow, Poland).

#### **3. Results**

For a better interpretation, all results are presented in the form of graphs. Figures 1–3 show the differences in pH changes over time depending on the dental materials used.

**Figure 1.** Repeated-measures analysis of variance for pH values in the individual time points depending on the dental restorative material (without subdividing by number of samples).

**Figure 2.** Repeated-measures analysis of variance for pH values in the individual time points depending on the dental restorative material for the one sample.

**Figure 3.** Repeated-measures analysis of variance for pH values in the individual time points depending on the dental restorative material for the two samples.

In general, all materials of a given group showed similar trends in pH change over time, regardless of the number of samples. Glass ionomer materials showed a significant increase in initial pH level from acidic to more neutral. The lowest initial pH level was found for the light-cured glass ionomer. However, for this material, the final pH level was higher by nearly 1.5. A similar increase was found for Ketac Molar. In contrast, for composite materials, a decrease in pH level by less than 1 was observed. At the endpoint, all materials had pH levels close to each other between 5 and 6, with Ketac Molar having the highest value. Moreover, after about a week, this material had a significantly higher environmental pH than all the others.

The changes in pH over time for individual materials depending on the number of the samples are presented in Figures 4–8.

For glass ionomer materials, already after three days, significant differences between pH levels can be seen depending on the number of samples. In the case of dual samples, lower values were obtained. In contrast, for Evetric composite material, significant differences were observed in pH level depending on the number of samples during the first three days. Additionally, the two samples showed significantly lower values, although higher variability was determined for the measurements of composite materials. On the other hand, Filtek Bulk Fill was the only material without pH variation depending on the number of samples, although there were changes in pH over time.

**Figure 4.** Repeated-measures analysis of variance for pH values in the individual time points depending on the number of the samples for Ketac Molar Easymix.

**Figure 5.** Repeated-measures analysis of variance for pH values in the individual time-points depending on the number of the samples for RIVA SC.

**Figure 6.** Repeated-measures analysis of variance for pH values in the individual time points depending on the number of the samples for RIVA LC.

**Figure 7.** Repeated-measures analysis of variance for pH values in the individual time points depending on the number of the samples for Filtek Bulk Fill.

**Figure 8.** Repeated-measures analysis of variance for pH values in the individual time points depending on the number of the samples for Evetric.

#### **4. Discussion**

Unlike many researchers, we have chosen saline as a research medium to create an oral environment [13,18,24]. It should be noted that the use of distilled water, even at human body temperature, does not reflect oral conditions [25]. Several authors have already used saline with interesting findings [26–28]. Researchers also use artificial saliva to determine the pH of dental materials [29,30]. However, its composition is very variable; the search for the best composition is still ongoing [31]. Apart from the eating situation, the oral cavity contains maximally about 1.1 mL of saliva as this amount forces the swallowing reflex [32]. For our analysis, we used 5 mL of saline so that the pH electrode was completely immersed in the solution.

In our study, we found significant differences in the effects of different materials on the environmental pH. After 15 days, all tested materials reached a pH between 5 and 6 on the pH scale. It seems that this fact may have some clinical implications.

It is well known that glass ionomer cements in the initial setting phase are characterised by a low pH that can irritate the pulp [13,15]. Therefore, it was surprising that they finally reached a pH higher than composites. Interestingly, conventional glass ionomers acidified the environment to a lesser extent than those modified with resin during the entire 15-day observation period. It is believed that this is a consequence of the lack of any buffering capacity of composite materials [4,33]. Many researchers indicate this fact as the main cause of secondary caries in materials based on synthetic resin [34,35]. We also observed that a single sample of the GIC lowered the environmental pH less than dual samples. The highest final pH about 5.9 was achieved by Ketac Molar. According to other researchers, it may even reach a pH close to 7 [13]. The first three days of material maturation seem to be crucial, when the pH increase was faster. The pH variation of Riva SC was similar but with a lower amplitude of the pH value at about 0.8 (compared to 1.4). This behaviour can be explained by releasing unreacted acrylic (or another organic acid) and its calcium salt, a weak acid, and conjugate base, which constitute a typical chemical buffer [35]. Some researchers state that it may result from the acid groups being bound to polymer molecules with limited diffusivity [2]. The most acidifying material was Riva LC which had the lowest initial pH of below 4.0 (below the critical point for enamel). In the

case of two samples, it was also the most acidic material after 15 days with a pH of about 5.2. It is commonly believed that RMGICs are more composites than glass ionomers [36,37], which was confirmed by the findings of our experiment.

The research methodology on composite materials focuses largely on subjecting them to cyclical changes in pH [38,39]. However, there is no clear answer as to whether these materials are chemically stable at a neutral pH. Despite the daily and routine placement of composite restorations, dentists' knowledge about the composite properties and the main factors of the polymerisation process is not adequate. It can lead to incorrect curing protocol, resulting in the incomplete polymerisation as well as relying on residual monomers and pH changes [40,41]. Moreover, chemical degradation of composite resin results in the release of final products, such as methacrylic acid which could acidify the environment [42,43].

Contrary to other studies [35,44], composite materials used in our research (Filtek Bulk Fill and Evetric) acidified the environment. Despite the initially relatively high pH of about 6, we observed a significant decrease to about 5.3 during the next 3 days of measurements. Around day 6, the pH started to rise slightly, and on day 15, it reached about 5.5. The Evetric composite, containing the Bis-EMA resin, turned out to be less acidifying in our study. Some researchers suggest that Bis-EMA, due to its hydrophobicity and high conversion character, is characterised by lower water sorption and lower solubility [45]. Another theory claims that UDMA composites are less hydrophilic than Bis-GMA-based composites. Also, it is suggested that the main difference between UDMA and Bis-GMA is their flexibility. Therefore, systems containing carboxylic or phosphate groups as functional monomers are more hydrophilic than the resins containing the Bis-GMA/TEGDMA monomer system [41,46]. It is believed that differences in the water absorption of the polymer network can be observed depending on the type of monomer. TEGDMA, Bis-GMA and UDMA seem to be the most hydrophilic [41].

Our study showed that medium-sized fillings are capable of inducing pH changes in the aquatic environment. Interestingly, the acidifying effect was visible in the case of resin-based materials, and two fillings lowered the environmental pH more than one. The literature shows that the lower the cross-linking of the composite, the higher the water absorption. It follows that the polymerisation quality of the composite may affect the pH stability [46]. In clinical practice, the irradiation of a portion of the material is not always optimal—a lower degree of conversion may adversely affect the quality of the filling and its subsequent durability [47]. We are planning further analyses to assess the possibility of changes in acidification by modifying the exposure time of the samples, the distance from the polymerisation lamp, and the diameter of the lamp optical fibre. The soaking of the filling materials in saline would also affect the dimensional, mechanical, and chemical properties. Therefore, it would also be advisable to evaluate the surface roughness of the samples or the release of chemicals such as residual monomers and metal ions.

#### **5. Conclusions**

Restorative materials lowered the environmental pH immediately after application, especially light-cured glass ionomer cements. Within two weeks, the pH level for glass ionomers rose to values comparable to composite materials. Moreover, dual samples reached lower pH levels than single samples. Self-cured materials showed larger amplitudes of pH changes than light-cured ones.

**Author Contributions:** Conceptualisation, A.L.; methodology, A.L. and K.N.; formal analysis, K.N. and A.L.; investigation and resources, A.L., K.N., M.N., P.W. and M.M.; writing—original draft preparation, A.L., K.N., and M.N.; writing—review and editing, K.N., A.L. and A.S.; visualisation, K.N.; supervision, A.S. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** Data are available on request from the corresponding author.

**Acknowledgments:** The project was implemented with the use of funds for science, awarded by the Poznan University of Medical Sciences.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

## *Article* **SEM and FT-MIR Analysis of Human Demineralized Dentin Matrix: An In Vitro Study**

**Lucia Memè 1,\*, Enrico M. Strappa 2, Riccardo Monterubbianesi 1, Fabrizio Bambini 1 and Stefano Mummolo 3**

**Abstract:** Recently, the demineralized dentin matrix has been suggested as an alternative material to autologous bone grafts and xenografts for clinical purposes. The aim of this study was to investigate the effect of different times of demineralization on the chemical composition and the surface morphology of dentinal particles. Extracted teeth were ground and divided into 5 groups based on demineralization time (T0 = 0 min, T2 = 2 min, T5 = 5 min, T10 = 10 min, and T60 = 60 min) with 12% EDTA. The analysis was performed using Fourier-Transform Mid-Infrared spectroscopy (FT-MIR) and Scanning Electron Microscopy (SEM) (*p* < 0.05). The FT-MIR analysis showed a progressive reduction of the concentration of both PO4 3− and CO3 2− in the specimens (T0 > T2 > T5 > T10 > T60). On the contrary, the organic (protein) component did not undergo any change. The SEM examination showed that increasing the times of demineralization resulted in a smoother surface of the dentin particles and a higher number of dentinal tubules.

**Keywords:** demineralized dentin matrix; human demineralized dentin matrix; human; bone graft; FT-MIR; SEM

#### **1. Introduction**

After tooth extraction, hard and soft tissue withstand remodeling processes. Alveolar ridge undergoes resorption, mostly in the horizontal but also in the vertical dimension [1]. Several studies analyzed the changes after tooth extraction [2]. The majority of horizontal and vertical changes take place during the first 3–6 months after tooth extraction and continue through the first year [3]. In fact, the horizontal bone loss is higher on the buccal side of the alveolar ridge than on the lingual/palatal side. On the other hand, the vertical resorption is minor and mainly on the buccal aspect of the alveolar ridge. This volumetric contraction of the alveolar ridge may jeopardize an appropriate implantsupported prosthetic rehabilitation.

Several surgical techniques have been developed to reduce or, at least, minimize the changes of soft and hard tissue following tooth loss. The main alveolar ridge preservation techniques concern on soft-tissue preservation, hard-tissue preservation (guided bone regeneration) and the combination of soft tissue and hard tissue preservation (socket seal technique) [4]. Moreover, different regenerative techniques have been tested: the use of bone grafts alone; barrier membranes alone, either resorbable or not; the combination of barrier membranes and bone graft. Metanalyses have shown that alveolar ridge preservation techniques are effective in significantly reducing vertical and horizontal alveolar ridge contraction [5,6]. Also Troiano et al. [7] reported positive results in the use of bone grafting and resorbable membrane compared with spontaneous healing. Autologous bone represents the ideal graft material due to its osteoinductive and osteoconductive properties [8]. Nevertheless, its limits are the small amount of bone graft available, the morbidity

**Citation:** Memè, L.; Strappa, E.M.; Monterubbianesi, R.; Bambini, F.; Mummolo, S. SEM and FT-MIR Analysis of Human Demineralized Dentin Matrix: An In Vitro Study. *Appl. Sci.* **2022**, *12*, 1480. https:// doi.org/10.3390/app12031480

Academic Editor: Vittorio Checchi

Received: 17 December 2021 Accepted: 27 January 2022 Published: 29 January 2022

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

of the donor area, and the risk of resorption of the bone graft itself. The research has led to developing alternatives, such as allogenic grafts, alloplastic grafts, and xenogenic grafts [9].

In the past decades, the use of demineralized dentin matrix (DDM) as a potential bone substitute has been proposed. Dentin consists of (i) 70% inorganic component (hydroxyapatite, tricalcium phosphate, octacalcium phosphate, and amorphous calcium phosphate); (ii) 20% organic component (collagen I 90%, collagen III and V in small quantities, and non-collagenic proteins); (iii) 10% water [10]. Its composition is, therefore, similar to that of bone tissue. Dentin also contains various growth factors, such as Fibroblast Growth Factors -2 (FGF-2), transforming growth factors-β1 (TGF-β1), insulin growth factor-1 (IGF-1) and, above all, bone morphogenetic proteins (BMPs) involved in the osteogenesis process [10].

Through demineralization, dentin can release the growth factors [11] and consequently express its osteoinductive and osteoconductive properties [12]. Thereby acting as a scaffold, dentin promotes the formation of bone tissue (osteoconduction), but at the same time, it releases the growth factors that promote the formation of bone (osteoinduction). Moreover, different in vivo studies have reported encouraging results [13–15] that led to this material being considered as a possible alternative to autologous bone grafts [16].

It has been shown that the preparation procedure, the shape, and the size of the dentin particles can influence dentin's osteoconductive and osteoinductive properties [17]. Excessive demineralization can damage the structure of the dentin and negatively affect the composition and function of growth factors; on the other hand, a reduced demineralization produces a scaffold with osteoconductive and osteoinductive properties [18].

The aims of this study were: (i) to examine the changes in the chemical composition of dentin particles using Fourier-transform mid-infrared spectroscopy (FT-MIR) analysis after different exposure times to demineralizing agent; (ii) to evaluate the surface morphology of the dentin particles by Scanning Electron Microscopy (SEM) analysis after different exposure times to demineralizing agent.

#### **2. Materials and Methods**

#### *2.1. Experimental Design and Specimen Preparation*

For this study, we used extracted teeth with unfavorable prognoses with the consent of patients. The criteria for exclusion were: (i) teeth with carious lesions; (ii) teeth with root canal treatment. After extraction of the specimens, the enamel and root cementum were removed with a dental drill (Figure 1a). Next, each tooth was washed with physiological solution, rinsed with an air-spray tool, and stored at −20 ◦C. The teeth were placed in a sterile container (Figure 1b) and ground using a Smart Dentin Grinder (KometaBio Inc., Cresskill, NJ, USA) for 3 s (Figure 1c). Then, the dentinal powder was sieved to distinguish two specimens with different grain sizes: the first with granules of smaller sizes (<300 μm) (A); and the second with granules of sizes between 300 μm and 1200 μm (B). The specimens were collected in sterile containers (Figure 1d,e). In this study, we used only the second specimen (B), while the first was excluded (specimen A).

Following the instructions of the machine producer, the dentinal powder was immersed in basic alcohol for 10 min in a sterile glass container. The basic alcohol was composed of 0.5 M NaOH and 20% ethanol. Next, the dentinal powder was picked up and washed two times with sterile phosphate-buffered saline solution (3 min for each wash) (Figure 1f). The dentinal powder was washed with a physiological solution and rinsed as much as possible.

The dentinal powder was randomly divided into five groups (0.075 g of dentin each) based on the time of demineralization (Table 1): T2, 2 min; T5, 5 min; T10, 10 min; T60, 60 min. T0 was not-demineralized and was considered as the control group.

**Figure 1.** Specimen preparation: (**a**) Example of the final specimens after enamel and cementum removal with a dental drill; (**b**) specimens in the sterile trituration chamber; (**c**) detail of Smart Dentin Grinder machine (KometaBio Inc., Cresskill, NJ, USA); (**d**) ground dentin with particle size less than 300 μm; (**e**) ground dentin with particle size between 300 μm and 1200 μm; (**f**) dentin particulate immersed in a sterile container with sterile saline solution.

**Table 1.** Group nomenclature based on demineralization process.

12 %EDTA was used to demineralize the dentin. Specifically, 1.7 g EDTA disodium salt was dissolved in 10mL distilled water at 25 ◦C. A saturated solution was obtained, from which only the supernatant was taken. Through this process, we obtained EDTA at 12%. The choice of this concentration was made according to previous articles published in the literature [12,17]. An amount of 1.1 mL EDTA was applied for the established time using a specific tool (Vortex). Once the process was completed, EDTA was removed, and the specimens were stored at −20 ◦C. Each specimen was washed two times with a physiologic solution (400 μL) to stop the demineralizing process. Next, the physiologic solution was aspirated.

The water had to be removed from the specimens to allow the spectrophotometry analysis. This process was completed by putting the specimens in a centrifuge (SpeedVac Concentrator, Savant SPD111V) (Figure 2a) for 30 min. The specimens were also weighed at different times: (i) ground and centrifuged dentin, before the demineralization process (D1); (ii) ground and demineralized dentin (D2); (iii) ground and centrifuged dentin after demineralization process (D3).

**Figure 2.** Instruments used in this study. (**a**) SpeedVac Concentrator, Savant SPD111V was used for the elimination of residual water; (**b**) FT-IR Perkin Elmer Spectrum One. Acquisition range 4000–450 cm−1; (**c**) scanning electron microscope (Supra 40, Zeiss).

#### *2.2. FT-MIR Analysis*

FT-MIR analysis was carried out using the FT-IR Perkin Elmer Spectrum One (Figure 2b) (Department of Life and Environmental Sciences, Polytechnic University of Marche, Ancona, Italy). The analysis of all specimens was performed in reflection UATR mode (Attenuate Total Reflection) with a range between 4000–600 cm−1 (spectral resolution 4 cm−1, 16 scans). The bands used for this study were: band 1021 cm−1, which indicated the percentage of PO4 3−; band 1649 cm−1, which corresponded to the protein component and was performed to determine the percentage of mineralization of each specimen; and band 872 cm−1, which reflected the CO3 2− percentage.

#### *2.3. SEM Analysis*

The specimens were gold-coated and analyzed through SEM (Department of Materials, Environmental Science and Urban Planning, Polytechnic University of Marche, Ancona, Italy). The superficial morphology of each specimen was evaluated through detector SE2 (Figure 2c) at different magnifications: (i) 400×; (ii) 2000×; (iii) 8000×.

#### *2.4. Statistical Analysis*

Statistical analysis of data included analysis of variance (ANOVA) and Tukey's test (*p* < 0.05).

#### **3. Results**

*3.1. Weight Analysis of the Specimens*

Table 2 shows the average weight of D1, D2 and D3.

**Table 2.** The weights of the specimens (expressed in g) in D1 (triturated and centrifuged dentin), D2 (triturated and demineralized dentin), and D3 (triturated, demineralized, and centrifuged dentin).

#### *3.2. FT-MIR*

MIR spectra were acquired for each specimen in the 4000–600 cm−1 spectral range, and NIR in the 10000–4000 cm−1 range. The main MIR absorption bands considered in this study are shown in Figure 3, which shows the spectrum of the dentin not subjected to demineralization (T0). Based on previous studies [19–21], the main absorption bands analyzed in this study were: (i) ~1649 cm−1 band corresponding to the protein component of dentin (band A); (ii) ~1021 cm−1 band corresponding to the phosphate ion PO4 3− (band B); (iii) ~872 cm−1 band corresponding to the stretching of the carbonate CO3 2− (band C).

**Figure 3.** Spectrum acquired through FT-MIR analysis of dentin particles not subjected to demineralization with 12% EDTA.

The variations in the MIR spectral profiles of the analyzed specimens (T0, T2, T5, T10, T60) were acquired in the spectral range 4000–600 cm−1. Then, the spectra were normalized compared to the 1649 cm−1 band, which corresponded to the protein component of dentin, and which remained constant even after the demineralization treatment. Through this procedure, it was possible to observe the trends of the PO4 3− and CO3 2− groups (Figure 4).

**Figure 4.** Spectra of the analyzed specimens normalized at the level of the 1649 cm-1 band, which corresponded to the protein component of dentin. Reductions were observed in the band at 1021 cm-1, which resembled the phosphate group, and at 872 cm-1, which corresponded to the carbonate group.

The concentrations of the inorganic components of dentin (PO4 3− and CO3 2− groups) decreased with increasing exposure time to the demineralizing agent. The amount of the mineralized component of the dentin was greater in T0, followed by T2, T5, T10. The T60 specimens were subject to the longest time of demineralization and showed the least amount of mineralized component.

The ratio between band B (1021 cm−1), which described the percentage of PO4 3−, and band A (1649 cm−1), which corresponded to the protein component, was measured to determine the percentage of mineralization of each specimen (Figure 5a). Band C (872 cm−1), which reflected the CO3 2− percentage, was also compared with band A (1649 cm−1), which corresponded to the protein component (Figure 5b).

**Figure 5.** FT-MIR Analysis. (**a**) The ratio between band B (1021 cm−1), which corresponded to the PO4 3− group, and band A (1649 cm−1), which reflected the protein component; (**b**) the ratio between band C (872 cm−1), which represented the CO3 2− group, and band A (1649 cm−1), which corresponded to the protein component. Different letters represent statistically significant differences (*p* < 0.05).

At T0, dentin was not treated with 12% EDTA, so it was considered as the control group, assuming 100% mineralization. Next, we calculated the quantity of PO4 3− and CO3 2− as a percentage of each specimen (Figure 6).

**Figure 6.** The graph represents the percentage of mineralization of the analyzed specimens after different exposure times. The blue expresses the quantity (%) of the PO4 3− group; the orange describes the quantity (%) of the CO3 2− group.

#### *3.3. SEM*

Finally, in this study, an SEM morphological evaluation of the specimens was performed (Figure 7). For each specimen, the obtained microanalysis was derived from the average of the results obtained on the most representative area. The voltage used was 25 kW, while the focal length was 15 mm.

**Figure 7.** The acquired SEM images of the specimens T0, T2, T5, T10, T60 at different magnifications: 400×, 2000×, and 8000×.

The images were acquired at different magnifications (400×, 2000×, and 8000×). This analysis allowed us to investigate the form and number of dentinal tubules exposed on the surface of the specimen. At 2000× magnification, it was seen that the number of the dentinal tubules in the same surface area increased T0 < T2 < T5 < T10 < T60.

Comparing the image at 8000× magnification of T0 with the one of T60, we observed a difference at the surface of the specimens. In the T0 specimens, the surface of the dentin was rough and non-homogeneous, while the dentin surface of the T60 specimen was smoother and more homogeneous. According to these results, it was assumed that the surface roughness tended to decrease with longer exposure to the demineralizing agent.

#### **4. Discussion**

An ideal bone graft should be biocompatible, biomechanically stable, capable of degrading over a certain time, and exhibit osteoconductive, osteogenic, and osteoconductive properties [22]. Although the gold standard is exemplified by autologous bone, the limited amount of bone available, the morbidity of the donor site, and the high rate of resorption affect its use. Therefore, alternative materials for autologous bone grafts are required.

In recent years, it has been proposed that DDM be used as a potential bone graft. In Japan and Korea, this type of graft is widely used, as demonstrated by the number of

scientific works published in the literature by scientific authors [23]. It has been shown that the preparation process, the size, and shape of the dentin particles seem to influence their osteoinductive and osteoconductive properties [17].

In summary, the concentration of the phosphate group (inorganic component of dentin, corresponding to the spectral peak at 1021 cm−1) and the carbonate group (inorganic component of dentin, corresponding to the spectral peak at 872 cm−1) decreased with increasing time of demineralization (T0 > T2 > T5 > T10 > T60). In addition, with increasing demineralization time, the number of exposed dentinal tubules in the same surface area increased, and the particles become more homogeneous and smoother.

In the present study, the enamel and the cementum were removed from the extracted teeth. In fact, the hydroxyapatite in the enamel is structured as highly crystalline calcium phosphate, while the dentin contains hydroxyapatite in low crystalline calcium phosphate form. In the first case, the high crystalline content is not easy to decompose by osteoclasts. Consequently, the resorption rate is slow, and the material's osteoconductivity is reduced. On the contrary, hydroxyapatite in dentin has a low crystalline structure, and this makes its resorption easier [11]. Bone tissues also contain low crystalline apatite. Recently, Elfana et al. [24] performed a randomized clinical trial comparing autologous whole tooth grafts and the autologous demineralized dentin grafts. The histological results showed a higher amount of newly formed bone and a smaller number of remnant grafts in the autologous demineralized dentin grafts group. The authors hypothesized that in the autologous demineralized dentin grafts group, the lower mineral content made particle degradation faster than that in the autologous whole tooth grafts group, and this allowed the release of growth factors earlier.

Regarding the size of the particles, a clear consensus has not yet been reached on which precise size is most suitable for bone grafts. Shapoff et al. [25] stated that the particle size of bone grafts should be between 100 μm and 300 μm. Nam et al. [26] conducted an in vivo study testing the new bone formation capabilities of DDM grafts with different densities and particle sizes. The histomorphometric analysis demonstrated the superiority of the specimens with grafting particles sized between 250 μm and 1000 μm and spaces of 200 μm between the particles compared to the results obtained with grafts with larger particles (1000–2000 μm). Koga et al. [12] reported better results in terms of new bone formation with particle sizes between 1200 μm and 800 μm compared to those obtained with smaller particle sizes. The authors also observed that the smaller particles underwent faster resorption than the larger ones. Therefore, it was suggested that the larger-sized particles (1200–800 μm) offer a greater surface area than those of smaller sizes (180–212 μm and 425–600 μm) for the adhesion of osteoprogenitor cells and osteoblasts. In addition, the adhesion of these cells could prevent the absorption of DDM particles and start the formation of new bone. For these reasons, in our study, demineralized dentin particles ranging in size from 300 μm to 1200 μm were used, excluding the smaller-sized particles (<300 μm).

Many studies have reported different tooth processing methods. Generally, there are four main categories [17]: (i) extraction of non-collagenic proteins from dentin; (ii) demineralization; (iii) elimination of the organic matrix (denaturation); (iv) use of tooth particles without modification. Denaturation is a little-used method because it eliminates the proteins in the matrix, including the growth factors responsible for the osteoinductive capacity of the dentin itself. Currently, the most used protocol is demineralization, as demonstrated by a large number of studies in the literature. The most used demineralizing agents are: (i) EDTA [17,24,27]; (ii) HNO3 [12,27]; and (iii) HCl [28–30].

Demineralization does not affect the organic component of the dentin or damage the growth factors contained therein. This process increases the osteoinductivity of the dentin particles, since it promotes the release of growth factors [11], favors the adhesion of osteoblasts through the exposure of collagen fibers [12], and reduces dentin's antigenicity. Demineralization is necessary because crystalline hydroxyapatite inhibits the release of growth factors, such as BMPs [11]. It has been observed that the amount of time during

which a demineralizing agent acts influences the characteristics of dentin. An excessive demineralization can damage the dentin structure and adversely affect the composition and function of odontogenic factors. On the other hand, a mild demineralization produces a scaffold with poor osteoinductive capabilities [18].

Tanoue et al. [27] performed an FIB/SEM analysis of the demineralized (HNO3 2%) dentin matrix grafted in a rat calvaria bone defect model. This method allowed the 3D reconstruction of the interface between the implanted dentin particles and the surrounding bone. The diameters of the exposed dental tubules averaged 3 μm. Mesenchymal cells, such as osteoblasts and bone tissue cells, are between 10–20 μm in size, while osteoclasts are 20–100 μm. The results of this study showed that osteocytes surrounded the grafted dentin particles, forming a network on their surfaces. In addition, cytoplasmatic extensions of osteocytes were observed in the dentinal tubules contained within the dentin particles. The authors also hypothesized a possible biological sequence of events that occurs when demineralized dentin particles are grafted into the recipient site. The release of BMPs induces mesenchymal cells to differentiate into osteoblasts, as also shown by the immunohistochemical analysis of de Oliviera et al. [31]. These cells produce a matrix that undergoes mineralization and forms new bone. At this point, the osteoblasts differentiate into osteocytes, which adhere to the surface of the demineralized dentin particles, forming a network on their surface. Cytoplasmatic extensions from the network spread into the dentinal tubules contained within the dentin particles. Afterward, the dentin particles will be reabsorbed and replaced by new bone, as confirmed by Kim et al. [13].

BMPs belong to the large family of TGF-β. Urist [32] was the first to describe their biological activity. In the following years, considerable efforts were made to isolate these growth factors and study both their in vitro and in vivo features. The BMPs promote the differentiation of mesenchymal cells into osteoblasts and chondroblasts [33], participating in the development of bone and cartilage [34], rather than the formation of dental hard tissues [35]. Bessho et al. [36,37] showed that BMPs obtained from demineralized dentin have osteoinductive properties similar to those derived from bone tissue. The BMPs induced new bone formation through endochondral (indirect) and intramembranous (direct) ossification, as shown by the histological findings of Murata et al. [30].

In our study, the dentin particles underwent a demineralization process with 12% EDTA. This agent is widely used in endodontics as a chelating agent for the enlargement of the canals and the removal of the smear layer. The longer the EDTA works on dentin, the more evident its effects are on dentin, as demonstrated by the release of phosphorus [38]. The specimens in our study were divided based on the exposure time of EDTA (0 min; 2 min; 5 min; 10 min; 60 min). The effect of the demineralizing agent was demonstrated by the analysis of the weight of the specimens, the FT-MIR analysis, and also SEM.

The weight of each specimen was lower after the demineralization process. One of the most used methods to confirm demineralization is SEM [18]. The results of our SEM analysis showed that the longer the EDTA was allowed to work, the less rough, or smoother, the surfaces of the particles became. Furthermore, it was observed that the longer the duration of demineralization, the greater the number of dentinal tubules exposed on the same surface area of the particles.

Koga et al. [12] conducted an in vivo study comparing non-demineralized, partially demineralized, and completely demineralized dentin specimens. Similarly to our study, the SEM analysis showed that the surfaces of the demineralized dentin specimens were smoother while those of the non-demineralized dentin were rougher. Interestingly, the osteoblasts adhered only to the surfaces of demineralized dentin but not to those of nondemineralized dentin. The authors hypothesized that the exposure of collagen fibers following demineralization could promote the adhesion of osteoblasts. The specimens of partially and completely demineralized dentin matrix, above all, showed greater osteogenic power than the non-demineralized ones.

Similar SEM results were achieved by Tabatabei et al. [18] performed a SEM analysis of dentin particles, reaching results similar to those of our study. The specimens of dem-

ineralized dentin showed greater exposure of the dentinal tubules and smoother surfaces compared to the non-demineralized specimens. Furthermore, the surfaces of the demineralized dentin particles demonstrated better suitability for cell proliferation (human dental pulp stem cells) than the surfaces of the non-demineralized particles. The authors also reported that the surfaces of the demineralized dentin particles were less biocompatible than those of the deproteinized dentin particles. It was suggested that these results were due to the demineralization process, during which a certain amount of the proteins may have been denatured. The authors also assumed that if the release of proteins following the demineralization process exceeds a certain threshold, it could be unfavorable, resulting in lethality for the cells.

The FT-MIR analysis was performed to investigate the functional groups on the surfaces of dentin particles [18]. This analysis offered information regarding the degree of demineralization for each specimen. The results of our study showed that the inorganic mineral component (CO3 2− and PO4 3−) content decreased with increasing duration of exposure to 12% EDTA. In contrast, the protein component remained unchanged. The reference spectra values were considered on the basis of findings from previous investigations [19–21].

Clearly, demineralization is effective but within a certain range. If the dentin is poorly demineralized, it results in a poorly osteoinductive scaffold. Nevertheless, if the demineralization is excessive, the substrate becomes ineffective. Therefore, an appropriate balance must be achieved. The demineralized dentine particles have a porous structure due to the presence of the dentinal tubules. This means that the dentin particles may act as an osteoconductive scaffold, which allows cells proliferation. Literature on the effects of different demineralization agent exposure times on the biological properties of dentin particles is limited. In fact, most of the studies compare specimens of demineralized dentin, non-demineralized dentin, and deproteinized dentin. To the best of our knowledge, there are no in vitro or in vivo studies regarding the effect of different durations of demineralization on dentine particles. Although the present study has some limitations, as the limited demineralized substances used and the possibility of inaccurate removal of enamel and cementum from the specimens, it can be considered a pilot study for future in vivo studies to find the best degree of demineralization that can be used as a graft material for bone regeneration.

It seems that the use of DDM offers several advantages. Both dentin and alveolar bone tissue derive from the neural crest and also share similarities in composition. Dentin consists of (i) 70% inorganic components (hydroxyapatite, tricalcium phosphate, octacalcium phosphate, and amorphous calcium phosphate); (ii) 20% organic components (mainly collagen I; collagen III and V in small quantities; non-collagenic proteins); and (iii) 10% water [10]. Non-collagenous proteins of dentin are known to be involved in bone calcification [39]. Dentin has shown not only osteoconductive but also osteoinductive properties [11,13]. Clearly, DDM can be considered an autologous graft since it is obtained from the extracted tooth of the same patient. Therefore, there is no risk of a cross-infection or rejection reaction, as confirmed by retrospective clinical studies of Lee et al. [40] and Kim et al. [41]. Although the results of in vitro and in vivo studies suggested that DDM can be considered as an alternative to autologous bone grafts [42] and xenografts [15,43], the number of clinical studies is still limited. Hence, further studies are still needed to validate the performance of DDM for clinical purposes.

#### **5. Conclusions**

Demineralized dentin matrix could be considered as a suitable alternative to autologous bone grafts and xenografts. According to previous studies, an adequate balance should be achieved in the demineralization process of dentin particles. However, due to the lack of scientific data, this study described the chemical and surface characterization of DDM after different demineralization processes. The following can be concluded:

**Author Contributions:** Conceptualization, F.B; methodology, L.M., E.M.S., F.B. and S.M.; software, S.M.; validation, L.M. and F.B.; formal analysis, E.M.S. and R.M.; investigation, E.M.S. and S.M.; resources, L.M. and F.B.; data curation, E.M.S. and F.B.; writing—original draft preparation, E.M.S. and F.B.; writing—review and editing, E.M.S. and R.M.; visualization, F.B.; supervision, L.M. and F.B.; project administration, L.M. and F.B.; funding acquisition, L.M. and F.B. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** The data presented in this study are available on request from the corresponding author.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

## *Article* **Prospective Pilot Study of Immediately Provisionalized Restorations of Trabecular Metal-Enhanced Titanium Dental Implants: A 5-Year Follow-Up Report**

**Peter van der Schoor 1,†, Markus Schlee 2,3 and Hai-Bo Wen 4,\***

**Abstract:** Porous tantalum trabecular metal biomaterial has a similar structure to trabecular bone, and was recently added to titanium dental implants as a surface enhancement. The purpose of this prospective pilot study was to describe 5-year survival results and crestal bone level changes around immediately-provisionalized Trabecular Metal Dental Implants. Eligible patients were adults in need of ≥1 implants in the posterior jaw. A non-occluding single acrylic provisional crown was in place for up to 14 days before final restoration. Clinical evaluations with radiographs were conducted at each follow-up visit (1 month, 3 months, 6 months, and 1 to 5 years). The primary endpoint was implant survival, characterized using the Kaplan-Meier method. The secondary endpoint was changes in crestal bone level, evaluated using a paired *t*-test to compare mean crestal bone levels between the baseline, 6-month, and annual follow-up values. In total, 30 patients (37 implants) were treated. Mean patient age was 45.5 years, and 63% were female. There was one implant failure; cumulative survival at 5 years was 97.2%. After the initial bone loss of 0.40 mm in the first 6 months, there were no statistically significant changes in crestal bone level over time up to 5 years of follow-up.

**Keywords:** dental implant; osseoincorporation; osseointegration; trabecular metal

#### **1. Introduction**

Porous tantalum has been used in the orthopedic industry since the 1980s for joint arthroplasty, bone augmentation, and as an integral part of orthopedic implants to facilitate bone ingrowth [1]. Although tantalum is highly biocompatible and resistant to corrosion, its use in orthopedics was initially limited due to cost and difficulties in manipulating solid tantalum [2]. However, in the early 1990s porous tantalum trabecular metal (PTTM) was introduced [2]. Porous tantalum trabecular metal, known commercially as Trabecular Metal (TM) Material (Zimmer Trabecular Metal Technology, Inc., Parsippany, NJ, USA) is a highly porous biomaterial made from elemental tantalum, with a similar structure to trabecular bone by having a repeating pattern of regular, interconnecting pores, which provide a high volume (70% to 80%) of porosity [1–3]. Using a proprietary chemical deposition process, elemental tantalum is deposited onto a substrate, creating a nanotextured surface topography to build the TM material [4].

Early nonclinical (canine and bovine) research showed that the multi-dimensional enhancement of implant surfaces with PTTM allowed for excellent bone ingrowth within porous tantalum structures [1,3,5,6] and that the material was chemically stable and biocompatible [7,8]. Porous tantalum trabecular metal has excellent biocompatibility and high frictional characteristics, which make it conducive to biologic fixation [1,2], and it has been used in clinical applications for orthopedics for several decades. PTTM has been recently added to titanium dental implants as an enhancement to titanium implant surfaces [2].

**Citation:** van der Schoor, P.; Schlee, M.; Wen, H.-B. Prospective Pilot Study of Immediately Provisionalized Restorations of Trabecular Metal-Enhanced Titanium Dental Implants: A 5-Year Follow-Up Report. *Appl. Sci.* **2022**, *12*, 942. https:// doi.org/10.3390/app12030942

Academic Editor: Vittorio Checchi

Received: 21 December 2021 Accepted: 15 January 2022 Published: 18 January 2022

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

As originally described by Bencharit et al., using PTTM as an enhancement for dental implants can potentially provide multiple advantages, including rapid endothelial budding, as well as ingrowth and endothelial neovascularization, both of which are critical for promoting new osseous tissue formation [2]. Based on the extensive prior clinical use of PTTM in orthopedics, a titanium alloy dental implant with a PTTM midsection was developed (Trabecular Metal Dental Implant, Zimmer Dental Inc., Carlsbad, CA, USA). The TM dental implant is manufactured from a combination of titanium alloy and trabecular metal and designed for integration into hard tissue through a process of bone apposition to the threaded implant surface and by bone ingrowth into the porous surface. These dental implants are indicated for delayed and immediate restoration.

There has been limited clinical research on TM dental implants, with a maximum of 1-year follow-up results reported. In 2013, the authors of the current report published the 1st year results of a 5-year descriptive, pilot study to evaluate TM dental implants which were immediately restored with a provisional restoration and fully loaded within two weeks. Their evaluation showed that during the first year of the study, PTTM-enhanced titanium dental implants had a 97.3% cumulative implant survival rate and appeared to be able to withstand the clinical demands of immediate loading [4]. The authors noted that longer-term follow-up was needed to better evaluate the performance and clinical characteristics of the TM dental implant. The present report describes the 5-year survival results and crestal bone loss associated with immediately-provisionalized and fully loaded (within two weeks) TM dental implants in a descriptive 5-year pilot study.

#### **2. Materials and Methods**

#### *2.1. Study Design and Eligibility Criteria*

This was a prospective, two-center, 5-year, nonrandomized, descriptive pilot study. Enrollment was open to all qualifying male or female patients of at least 18 years of age with adequate bone volume to support an implant in the posterior jaw. Other inclusion criteria included having residual facial and palatal/lingual plates at least 1.5 mm thick after osteotomy preparation, having a healed extraction site greater than six months postextraction (with or without grafting), being able to provide informed consent, and ability to attend evaluation visits.

Exclusion criteria for enrollment into the study included bruxism, fresh extraction sites, smokers, uncontrolled systemic disease, bleeding disorders, use of biphosponates, and pregnancy. At the time of surgery, an additional exclusion criterion was employed; patients with type 4 bone or implants with <35 Ncm of insertion torque were also excluded. After a patient was deemed medically, dentally, psychologically, functionally, and anatomically to be a good candidate for dental implant therapy, informed consent was obtained, and the patient was enrolled. Up to 40 patients from two sites (one in Germany and one in the Netherlands) were planned for enrollment.

#### *2.2. Data Collection*

Before implant surgery, patient demographics, and medical and dental conditions were recorded. During the implant surgery, insertion torque values, implant size, and bone density were recorded. Radiographs were taken to evaluate mesial and distal bone levels. All data requested for this study were collected using electronic case report forms (eCRFs). Data collection and management were performed by the Sponsor.

#### *2.3. Implant Procedure*

Pre-operative photographs and radiographs were taken with an individualized imaging holder (Figure 1). All patients received pre-operative oral antibiotic therapy with either clindamycin or amoxicillin [9]. At the time of the surgical implant placement, the subject was administered anesthesia. As described in an earlier publication, implants of either 4.7 mm or 6.0 mm diameters were placed in either premolar or molar sites, in either jaw [4]. Implants were TM dental implants with a 0.5 mm machined collar, microtextured (MTX)

surface, and microgrooves, and were placed according to the surgical manual using a one-stage (nonsubmerged) technique. Implant insertion torque and resonance-frequency analysis values were recorded at the time of placement. As noted earlier, if it was discovered that a patient had type 4 bone at the implant site, they were excluded from the study. Appropriate analgesics and postoperative physical therapy were prescribed as needed. Post-implantation antibiotic therapy was not prescribed unless indicated for infection. After the implant placement, a final impression was taken and sent to the dental lab for fabrication of the final restoration.

**Figure 1.** An individualized imaging holder was made and stored for each patient during the 5-year study period.

#### *2.4. Immediate Provisional and Final Restoration*

A standard contoured abutment was placed as the final abutment with 30 Ncm preload, as recommended. This was the abutment for both the provisional and final crown. A nonoccluding single acrylic crown provisional prosthesis was immediately delivered and the soft tissues were sutured around the provisional restoration. All immediate provisional restorations were placed within 48 h of surgery. The provisional prosthesis was in place for 7 to 14 days to allow adequate time for soft tissue healing. Periapical radiographs were taken. Seven to 14 days after the implant procedure, the provisional prosthesis and sutures were removed. If the implant appeared stable, the final restoration was delivered (Figure 2). The final restorations were fully occluding, single-tooth ceramometal restorations. Any final occlusal adjustments were made at this visit.

**Figure 2.** Shows the intraoral photos and periapical radiographs taken pre-operatively, at final restoration, and at 5-years follow-up. This female patient with missing posterior mandibular first and second molars received two TM implants (diameter 6.0 mm and length 10 mm for both). Final restoration was completed 13 days after implant placement consisting of two cement-retained crowns. The 5-year follow-up showed good soft tissue health and bone maintenance.

#### *2.5. Follow-Up Visits*

Subjects were scheduled to be re-evaluated at 1 month, 3 months, 6 months, and for annual visits for 5 years after delivery of the provisional restoration. During each follow-up visit, the study investigator took photographs in the facial, occlusal, and lingual views, assessed for any clinical complications, and took standardized periapical radiographs (Rinn, Dentsply, York, PA, USA) (Figure 1). An individualized imaging holder was made and stored for each patient during the 5-year study period. When mobility was suspected, opposing force from two hand instruments was applied to the post and any overt indication of mobility was recorded as a failure.

#### *2.6. Study Endpoints*

The primary endpoint was implant survival after 5 years of occlusal loading. Implant failure was defined as the presence of persistent pain, mobility, or loss or removal of the implant.

The secondary study endpoint was change in crestal bone level; mesial and distal levels were determined radiographically and a mean (per implant site) was calculated. Crestal bone loss was summarized as the mean of radiographic mesial and distal changes in bone height from baseline (implant surgery) values. Mean crestal bone loss served as the secondary endpoint, and was computed by averaging the mean bone loss per implant.

All periapical radiographs were evaluated by an independent radiologist using highresolution uncompressed image files. As described in an earlier paper, crestal bone levels were measured by calculating the distance from the implant shoulder to the first bone-toimplant contact [4]. Both mesial and distal measurements were made on each periapical radiograph, and the mean value was used in the analyses. The known height of the implant's tantalum section (4.8 mm) was used as the standardized dimension for calibration. The height of the tantalum section was measured on the image in pixels, and the ratio between the length in pixels and tantalum height of 4.8 mm was calculated. Because the two study sites used different radiographic image sensors, each site was calibrated differently: 0.0234 mm/pixel (4.8 mm/205.5 px = 0.0234 mm) for the first site (Germany) and 0.0349 mm/pixel (4.8 mm/137.5 px = 0.0349 mm) for the second site (the Netherlands). Bone height values measured in pixels were then multiplied by the calculated calibration factors to arrive at the final data values in millimeters. Each image was opened using US FDA-cleared image analysis software (OsiriX MD, Pixmeo SARL, Bernex, Switzerland) on a personal computer [4]. Measurements were recorded in a spreadsheet (Excel, Microsoft Corp., Redmond, WA, USA).

#### *2.7. Statistical Analyses*

Descriptive statistics (N, %, mean ± SD, min, max, median) were used to summarize the data. The implant was the unit of analysis, except for patient demographics. Implant survival was summarized through the characterization of failure over time, using the Kaplan–Meier method. Cumulative survival of the implants was estimated at each time of assessment, with corresponding 95% confidence intervals. A paired *t*-test was performed to analyze the initial bone loss from implant placement/immediate provisional restoration to 6 months of follow-up, and the additional bone loss was analyzed by comparing crestal bone loss between annual follow-up values and the 6-month value. Statistical significance was declared if the 2-sided *p*-value was <0.05.

Because this was a descriptive pilot study with no control group and no prior hypothesis, no formal sample size calculations were conducted. This pilot study was conducted to gather initial, exploratory data; the sample size was not selected according to any statistical power calculation. Instead, the sample size was determined according to feasibility and was thought to provide sufficient exploratory results to inform the development of future confirmatory studies. Analyses were performed using SAS 9.4 TS (SAS Institute Inc., Cary, NC, USA) or Minitab 18 (Minitab, LLC, State College, PA, USA).

#### *2.8. Human Subjects Protections*

The research was conducted in compliance with applicable regulations, and according to the principles of the Declaration of Helsinki. The pilot study was conducted in accordance with the respective government regulatory authorities and the local regional Institutional Review Boards for two study sites in Germany and the Netherlands (Protocol CSU2010- 07D Freiburger Ethik Kommission-an independent Ethics Committee-28 June 2010). All patient data were fully anonymized to safeguard patient confidentiality. All materials and procedures complied with local and international health and safety standards and good clinical practices. This report is structured in alignment with the Consolidated Standards of Reporting Trials (CONSORT) [10].

#### **3. Results**

After obtaining informed consent and verifying that potential subjects met the inclusion/exclusion criteria, 30 patients (37 implants) were treated per protocol. Each patient received one or two TM implants in premolar or molar sites (Figure 2). The mean patient age was 45.5 years (range 19 to 73 years), and the majority of patients were female (63.3%). Twenty-nine of the 30 patients were White, one patient was Asian. The most common comorbidity was hypersensitivity reactions/allergies (40%), followed by hypertension/hypotension (30%). Most of the implants (70.3%) were placed in the mandible. Regarding bone density, most implants (62.2%) were placed in type 2 bone, with the remaining placed in type 3 bone. The most common (and smallest) implant diameter that was used was 4.7 mm (64.9%), and the most common length was 10 mm (45.9%). The majority of implants (91.9%) received a provisional restoration on the final abutment at the time of surgery (Table 1). There were no surgical complications across the study cohort.

**Table 1.** Patient Demographics and Surgical Information.

SD = standard deviation. \* Subjectively assessed by the clinician based on radiographic evaluations and tactile sensations during implant placement.

Radiographic evaluations of mean crestal bone levels over time (from baseline time of implant placement/provisional restoration to the 5-year follow-up evaluation) showed an initial bone loss of 0.40 mm in the first 6 months, after which there were no statistically significant changes in crestal bone level over time (Table 2). The mean crestal bone value levels changed minimally (≤0.05 mm) from 6 months to any of the annual follow-up visits. Repeated Measures Analysis of Variance with Tukey Pairwise Comparison confirmed there is no statistical difference among the crestal bone level values of the 6-month and annual follow-up visits (details not shown). There was also no statistically significant difference in crestal bone level changes between those who had received 4.7 diameter versus 6 mm diameter implants (data not shown). Additionally, no implants were observed to have pain, mobility, or peri-implant radiolucency.

**Table 2.** Changes in Crestal Bone Levels Over the 5-Year Follow-Up Period.

1 Analyses were conducted using a paired *t*-test. The study population for the bone loss value dataset had several missing values from follow-up visits, which is why there is a slight inconsistency in sample size between the crestal bone level dataset and the survival analysis dataset.

At the end of the 5-year follow-up period, 25 patients remained for evaluation (Table 3). The Kaplan-Meier cumulative survival rate of all implants in the study was 97.2% at the 3-month, 6-month, 1-year, 2-year, 3-year, 4-year, and 5-year evaluation visits; one implant failed to osseointegrate. The patient with a failed implant was a 46-year-old White male who had received a 6 × 10 implant in a mandibular molar site. The implant was placed into type 2 bone without complications. The implant failed to integrate and was removed 30 days after the implant surgery. The patient took no concomitant medications (outside of post-surgical analgesics) and had no recorded comorbidities.

1 The Kaplan-Meier method was used to calculate cumulative implant survival. One implant failed within 2 months after loading.

There were no other failures. All 25 implants that completed 5-years of follow-up were surviving and clinically stable, with no indicators of pain, mobility, radiolucency, or clinically meaningful crestal bone loss.

#### **4. Discussion**

The results from this descriptive pilot study of 37 implants showed that immediate provisional restoration (out of occlusion, with final restoration and loading within two weeks) of titanium dental implants with surfaces enhanced with PTTM were clinically effective, with a 5-year survival rate of 97.2%. As a comparison to non-TM systems, a recent systematic review and meta-analysis of 18 dental (non-TM) implant studies evaluated the 10-year survival of contemporary, two-piece titanium systems; the summary estimate for 10-year survival at the implant level was 96.4% (95% CI 95.2–97.5%) [11].

Professional consensus in dental implantology affirms that the success of implants is dependent on the presence and preservation of surrounding bone, particularly in the crestal area; however, a primary challenge of dental implant treatment is the bone resorption, particularly saucerization, that occurs after insertion [12]. Results from the present study showed an initial mean bone loss of 0.40 mm (SD = 0.48) after 6 months and minimal bone level changes over time up to 5 years in function. While the small sample size of this uncontrolled study limits direct comparisons to findings in other studies, these findings suggest that crestal bone loss around TM implants is lower than traditional titanium implant systems [13,14], and that the unique properties of PTTM may foster enhanced vital bone and surrounding tissue ingrowth and potentially reduce the amount of bone remodeling which occurs during the first year following implant surgery [15].

A recent 1-year retrospective case-control study by Edelmann et al. reported a 0.28-mm mean bone gain in a TM implant cohort, and multivariate logistic regression analysis demonstrated that the odds of having bone loss were 64% less in a TM group compared to a non-TM implant control group [16]. Another recent study by Bencharit et al. compared osteogenesis gene expression in a small group of patients who had received both TM and non-TM implants; the authors found that, compared to traditional titanium alloy, trephine samples from the TM implant group displayed higher expression of genes specific to neovascularization, growth factors, and osteogenesis [17].

To further contextualize these crestal bone level changes when compared to non-TM implants, contemporary literature was surveilled for studies with similar sample sizes and follow-up periods. An analysis by Payer et al. found that a cohort of 24 patients with 40 screw-type (non-TM) implants that were immediately provisionalized in molar and premolar sites experienced an overall survival rate of 95% at 5 years, with significant marginal bone loss (1.06 mm) in the first year [18].

To compare the crestal bone findings in this report with historical standards, longitudinal studies on the original Brånemark implant showed that one year after abutment connection, crestal bone loss frequently extended 1.5 mm to 2 mm below the implantabutment junction [19]. This became the unofficial industry standard for acceptable crestal bone loss because most major implant systems experienced a similar bone loss phenomenon during the early years of their products. However, more contemporary implant surfaces typically include a rougher surface to foster better osseointegration, as improving the roughness of an implant surface has been correlated with less peri-implant bone loss [2]. Additionally, because the abutment was immediately placed with no intention of removal during the follow-up period, essentially the "one abutment/one-time" method was used in this study. The abutment was secured not only by the fixation screw, but also with a friction-fit connection. This type of "one-piece" implant has been shown to have less initial bone loss [20].

The TM implants used in this study have a PTTM shell that begins about 4.5 mm apical to the implant platform. A recent report by Fraser et al. compared osteogenic activity between titanium and TM-coated implants in a nonclinical study of rabbits. The authors reported higher vertical bone growth around TM implants compared to titanium implants,

and increased activity in upregulation of key osteogenic genes. Furthermore, TM implants had greater bone-implant contact at 4, 8, and 12 weeks and significantly greater removal torque at 8 and 12 weeks [21]. It is not yet fully known how the TM shell impacts periimplant bone healing and remodeling at the cervical aspect, though the current study shows promising results. The current clinical pilot study showed no statistically significant bone loss after the initial healing, which may in part be due to the high porosity and modulus of elasticity of the TM surface, which has similar properties to human cancellous bone. Prior studies have shown that these surface properties facilitate bone ingrowth and increased regions of bone-to-implant contact, potentially reducing stress on the coronal aspect of the implant due to rapid ingrowth of bone in the TM shell, though further confirmatory research with larger sample sizes is needed [1,3,5]. Taken together, these promising findings of increased osteogenesis indicators and crestal bone stability around TM implants warrant further study with larger study populations.

In this study, TM dental implants were clinically effective when fully loaded within two weeks in a small population with a variety of health risk factors that may be typically encountered in routine clinical practice.

#### *Limitations*

Patients were followed for 5 years post-restoration; during that interval, the number of evaluable implants went from 37 to 25, an attrition rate of 32%. This is an unfortunate challenge when obtaining long-term survival data, and may lead to a bias in the results [10]. Although the study investigators were aware of several patients moving away from the study site area, the reasons for other patients being lost-to-follow-up was unknown, as were their outcomes, which could potentially bias the study results and conclusions.

Additionally, because implants in this study were only placed in type 2 and 3 bone, the clinical response to TM implants in other bone types is unknown, possibly reducing the generalizability of the results.

#### **5. Conclusions**

Immediate provisional restoration (with final restoration within two weeks) of titanium dental implants with surfaces enhanced with PTTM were clinically effective in a small study population, with a 5-year survival rate of 97.2% and minimal crestal bone loss. These findings of potentially increased osteogenesis indicators and crestal bone stability around TM implants warrant further study with larger study populations.

**Author Contributions:** Conceptualization, P.v.d.S., M.S. and H.-B.W.; methodology, P.v.d.S., M.S. and H.-B.W.; investigation, P.v.d.S. and M.S.; data curation, P.v.d.S. and M.S.; writing—original draft preparation, P.v.d.S., M.S. and H.-B.W.; writing—review and editing, P.v.d.S., M.S. and H.-B.W.; project administration, P.v.d.S. and M.S. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research was funded by Zimmer Biomet Dental, Palm Beach Gardens, FL, USA.

**Institutional Review Board Statement:** The study was conducted in accordance with the Declaration of Helsinki and approved by an independent Ethics Committee (Freiburger Ethik Kommission) for two study sites in Germany and the Netherlands (Protocol CSU2010-07D, 28 June 2010).

**Informed Consent Statement:** Informed consent was obtained from all subjects involved in the study.

**Data Availability Statement:** The data presented in this study are not publicly available but can be requested from the corresponding author.

**Acknowledgments:** The authors are greatly indebted to Alexandra R M van der Schoor for her invaluable support in patient follow-up and clinical data collection and would also like to thank Na Ren and Cristina Matthews at Zimmer Biomet for their help in statistical analysis.

**Conflicts of Interest:** P.v.d.S. and M.S. declare no conflict of interest. H.-B.W. is an employee of Zimmer Biomet Dental.

#### **References**

## *Review* **Finding the Perfect Membrane: Current Knowledge on Barrier Membranes in Regenerative Procedures: A Descriptive Review**

**Sorina-Mihaela Solomon 1,†, Irina-Georgeta Sufaru 1,\*,†, Silvia Teslaru 1,†, Cristina Mihaela Ghiciuc 2,† and Celina Silvia Stafie 3,\*,†**