**1. Introduction**

The current manuscript represents an overview of literature reflecting the concern of gastroenterology physicians regarding the usefulness of fecal microbiota transplant as an appropriate and successful therapy in difficult to treat IBD patients. Inflammatory bowel disease (IBD) is a relapsing, remitting, and chronic disease that causes significant morbidity. The etiology of IBD is still unclear. The phenotype, the progression, and their development are multifactorial with environment and genetics. Nowadays, studies are confirming that the microbial influence in the

pathogenesis of IBD is increasing; this fact results from an inappropriate immune response towards components of the commensal microbiota. In IBD, the diversity of luminal microbiota is reduced. *Firmicutes* (bifidobacteria, lactobacillus, and *Faecalibacterium prausnitzii*) are especially decreased. On the opposite side, the mucosal adherent bacteria are increased [1, 2].

Both ulcerative colitis and Crohn's disease share many common features like bloody stools, diarrhea, fever, and abdominal pain, but each of them also has unique features. There are many differences between the two entities, the most important being the depth of involvement in the bowel wall and their location. Crohn's disease results in transmural ulceration of any portion of the gastrointestinal tract, but it affects most often the colon and terminal ileum; at the opposite side, ulcerative colitis affects the rectum, but it may extend beyond the sigmoid and into the sigmoid or include the entire colon into the cecum [1–3].


*Article* **High Accuracy and Safety of Intraoperative CT-Guided Navigation for Transpedicular Screw Placement in Revision Spinal Surgery**

**Yen-Yao Li 1,2,\*, Shih-Hao Chen 3, Kuo-Chin Huang 1,2, Chien-Yin Lee 1,4, Chin-Chang Cheng 5, Ching-Yu Lee 6,7, Meng-Huang Wu 6,7 and Tsung-Jen Huang 6,7**

	- Taoyuan City 333324, Taiwan

**Abstract:** Background: Intraoperative CT-guided navigation (iCT-navigation) has been reported to improve the accuracy and safety of transpedicular screw placement in primary spinal surgery. However, due to a disrupted bony anatomy and scarring tissue, revision spinal surgery can be challenging. The purpose of this study was to evaluate the accuracy and safety of iCT-navigation for screw placement at the virgin site versus the revision site in revision thoracolumbar spinal surgery. Method: In total, 254 screws were inserted in 27 revision surgeries, in which 114 (44.9%) screws were inserted at the site with previous laminectomy or posterolateral fusion (the revision site), 64 (25.2%) were inserted at the virgin site, and 76 (29.9%) were inserted to replace the pre-existing screws. CT scans were conducted for each patient after all screws were inserted to intraoperatively confirm the screw accuracy. Results: In total, 248 (97.6%) screws were considered accepted. The rate of accepted screws at the virgin site was 98.4% (63/64) versus 95.6% (109/114) at the revision site (*p*: 0.422). There were six (2.4%) unaccepted screws, which were immediately revised during the same operation. There was no neurological injury noted in our patients. Conclusion: With the use of iCT-navigation, the rate of accepted screws at the revision site was found to be comparable to that at the virgin site. We concluded that iCT-navigation could achieve high accuracy and safety for transpedicular screw placement in revision spinal surgery and allow for the immediate revision of unaccepted screws.

**Keywords:** thoracolumbar spine; revision spinal surgery; intraoperative computed tomography; navigation; transpedicular screw

#### **1. Introduction**

The transpedicular screw (TPS) system is one of the most widely used fixation devices in spinal surgery. It provides a three-column fixation [1] of the spine. It is essential to ensure the optimal trajectory of TPS insertion to offer the strongest bony purchase and prevent nerve tissue injury. TPS insertion can be challenging in revision spinal surgery [2], because the anatomical features used to identify the entry of a TPS may have been destroyed or obscured by the fusion mass or abundant scar tissue, causing a standard fluoroscopy to be insufficient to identify the proper entry of a TPS due to blurry images for altered or fused anatomy [2–4].

**Citation:** Li, Y.-Y.; Chen, S.-H.; Huang, K.-C.; Lee, C.-Y.; Cheng, C.-C.; Lee, C.-Y.; Wu, M.-H.; Huang, T.-J. High Accuracy and Safety of Intraoperative CT-Guided Navigation for Transpedicular Screw Placement in Revision Spinal Surgery. *J. Clin. Med.* **2022**, *11*, 5853. https:// doi.org/10.3390/jcm11195853

Academic Editors: Michał Strzelecki, Rafał Obuchowicz, Andrzej Urbanik and Adam Piorkowski

Received: 1 August 2022 Accepted: 30 September 2022 Published: 2 October 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/).

Several advanced image-guided techniques have emerged in the past two decades, including three-dimensional (3D) fluoroscopy [5,6], O-arm [7,8] and intraoperative computed tomography (iCT) [9,10]. One study [11] reported that the accuracy of the TPS was comparable for both primary and revision spine surgery using O-arm-guided navigation. However, some of the TPS placements in revision spinal surgery were performed at the virgin site for the extension of posterior instrumentation, instead of right over the revision site. Thus, the accuracy of TPS placement at the revision site could be different from that of the whole revision spine surgery. In the current study, we aimed to evaluate the accuracy and safety of iCT-navigation for TPS placement right over the revision site versus the virgin site in complex revision thoracolumbar spinal surgery.

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

We enrolled 27 consecutive patients who underwent TPS placement assisted with intraoperative CT (Siemens, Munich, Germany) integrated with navigation (Brainlab, Munich, Germany) for revision thoracolumbar spinal surgeries. The study was approved by the Institutional Review Board of the Chang Gung Memorial Hospital.

There were 27 patients (13 women and 14 men) in the present study, with an average age of 67 years (range 47–92 years). A total of 254 TPSs was inserted during the revision spinal surgeries, which were categorized into 3 groups: "revision site" screws (no. = 114) inserted over a previous laminectomy or posterolateral fusion field; "virgin site" screws (64) inserted at a site away from previous laminectomy or posterolateral fusion for the extension of posterior instrumentation; and "exchanged" screws (76), inserted to replace pre-existing screws with screws of a different style or brand through the previous trajectories.

#### *2.1. Surgical Technique*

After general anesthesia, the patient was positioned prone on a Wilson radiolucent frame (Mizuho OSI, California, CA, USA) on a Jackson table. A midline incision with paraspinal muscle stripping was performed to expose the previous operation site and extended to the virgin level if necessary. The pre-existing screws were exchanged for screws of a different brand, if necessary, which were inserted manually through previous trajectories without using iCT-navigation. Thereafter, the superior articular processes of the facets through which we planned to insert new TPSs were exposed. The reference array was then securely fixed at a vertebra distal to the expected lowest instrumented vertebra (usually L5 or S1) by tightly clamping its spinal process.

The operative field was scanned with the intraoperative CT, and the image data were transferred to the navigation system. A straight drilling guide was then registered to guide a 2.7 mm in diameter drill bit to create the optimal track for the TPS under iCT-navigation by checking the axial, sagittal and coronal views (Figure 1). A TPS with an optimal length and diameter was chosen and inserted through the pedicle trajectory after being adequately tapped. After completing all TPS placements, CT scans were conducted to confirm whether the TPSs were in the optimal position, which we called confirmatory CT scans (Figure 2).

#### *2.2. Assessment of TPS Placement*

Intraoperatively, with the aid of the confirmatory CT scan, we evaluated the TPS accuracy and divided the results into 2 conditions: "accepted" and "unaccepted". An accepted screw was defined as having a medial pedicle breach of ≤3 mm or a lateral pedicle or body wall breach of ≤3 mm. The position of the TPS was considered biomechanically optimal and safe from neurologic injury. An unaccepted screw was defined as having a medial pedicle breach of >3 mm or a lateral pedicle or body wall breach of >3 mm (Figure 3); the TPS could be biomechanically suboptimal and create a risk of nerve tissue injury, and, therefore, was immediately revised under iCT-navigation. After replacing the unaccepted screw, the second confirmatory CT scan was performed to ensure that the screw was in an accepted position.

**Figure 1.** The straight drilling guide (red arrow) was used to guide a 2.7 mm in diameter drill bit to create the optimal TPS track under iCT-navigation by checking the axial, sagittal and coronal views.

**Figure 2.** A 55-year-old male patient with post-laminectomy instability and spinal stenosis; X-ray (**a**) showed previous laminectomy of L4, L5 and partial L3. This patient underwent revision spinal surgery with the use of iCT-navigation; X-ray (**b**) showed the reference array fixed at the spinal process of S1 for iCT-navigation. Intraoperatively, after all screws were inserted, the confirmatory CT scan demonstrated sagittal and axial views of screws (**c**,**e**) at the virgin site and (**d**,**f**) at the revision site.

#### *2.3. Radiographic and Clinical Analyses*

The accuracy of the TPS placement was assessed with the first confirmatory CT. The rates of accepted screws and unaccepted screws were calculated for comparison between the virgin site and the revision site. Perioperative medical records were collected retrospectively, and postoperative neurological complications were evaluated.

**Figure 3.** Unaccepted screw (red star) was defined as (**a**) a medial pedicle breach of >3 mm or (**b**) a lateral pedicle or body wall breach of >3 mm.

#### *2.4. Statistical Analysis*

Descriptive statistics were reported as percentages to describe the rates of accepted screws and unaccepted screws. Fisher's exact test was used to compare the rates of accepted and unaccepted screws between the virgin site and the revision site. The test was considered significant if *p* < 0.05.

#### **3. Results**

In total, 27 patients were categorized using the etiology for revision spine surgery: 12 patients (44.4%) were diagnosed with postlaminectomy instability with or without posterolateral fusion; 9 (33.3%) patients had adjacent segmental disease with spinal stenosis or instability; 4 (14.8%) patients had malpositions of a previous TPS with clinical neurologic symptoms; and the remaining 2 (7.4%) patients had a refractory postoperative infection (Supplementary Table S1).

Of the total 254 TPSs, 248 screw placements were considered accepted (97.6%) by the first confirmatory CT, including all of the "exchanged screws" (N = 76). The rates of accepted screws were comparable both at the virgin site (63, 98.4%) and the revision site (109, 95.6%). There was no significant difference between both groups (*p* = 0.422).

There were six screws (2.4%) considered unaccepted: one (1.6%) at the virgin site with a medial pedicle breach; and five (4.4%) at the revision site, including two with a medial pedicle breach and three with a lateral pedicle or body wall breach. The rates of unaccepted screws trended higher at the revision site than at the virgin site, but were not significantly different. All the unaccepted screws were revised immediately with the assistance of iCT-navigation, and, then, each proved to be accepted with a secondary confirmatory CT. No postoperative neurologic deterioration was observed related to the screw placement.

#### **4. Discussion**

With the use of iCT-navigation for revision spinal surgery, the rate of accepted screws was 97.6%, which was comparable to other studies [2,11] for revision surgery. Hsieh et al. [11] adopted an O-arm with navigation to guide TPS placement. They reported "good or fair" rates of screw placement (pedicle breach <3 mm) as 98.7% in the primary surgery group and 98.6% in the revision surgery group, which were not significantly different. Even compared with our previous studies for primary surgery [10,12], which reported the accuracy of TPS placement was 96–98%, the accuracy in the current study for revision

surgery was comparable. However, we divided our series of TPS placements into two groups for comparison: the revision site and the virgin site. The rate of accepted screws was 98.4% (63/64) at the virgin site and 95.6% (109/114) at the revision site (*p* = 0.422). There was one (1.6%) unaccepted screw noted at the virgin site, and five (4.4%) at the revision site. The rates of unaccepted screws trended higher at the revision site than at the virgin site, but were not significantly different. We assumed that some of the TPSs in revision spinal surgery could be inserted at the virgin site for the extension of posterior instrumentation instead of right over the revision site. Thus, the accuracy of TPS placement at the revision site could be different from or inferior to that at the virgin site across the whole revision spinal surgery.

In the current study, most of the unaccepted screws were noted at the revision site. We considered dense postoperative fibrotic tissue or scarring to be the principal causes resulting in screw malpositioning at the revision site. Once the navigation device bent around the dense scar during the process of TPS placement, the navigation had the potential to fail. The scar tissue caused excessive pressure on the instrument and usually resulted in the medial deviation of the screw trajectory. Furthermore, if the surgeon exerted excessive bending force on the instrument (the drill or awl) to apply soft tissue pressure, the instrument could skive off the pedicle and result in the lateral skidding of the trajectory. The skiving effect might explain why there were more unaccepted screws presenting with a lateral wall breach (3, 60%) at the revision site. We suggest that the skiving effect could be diminished with the adequate release of soft tissue tension, and by using a high-speed drill with a shard bit for drilling [13].

There have been few studies focused on the feasibility or accuracy of TPS placement in revision spinal surgery. Kim et al. [2] reported that the accuracy of TPS placement in revision surgery was excellent with the free-hand technique. The revision rates of misplaced screws were 1.1% in the virgin site screw group and 2.9% in the revision site group, which were not significantly different. However, the TPSs were evaluated with plain radiography, not CT, which might have underestimated the number of misplaced screws. Free-hand TPS insertions in revision spinal surgery are technically demanding, and it is laborious to identify the entry point of the TPS when anatomic features are distorted. With the use of iCT-navigation in the current study, surgeons were able to confidently insert the TPSs during revision spinal surgery. Hsieh et al. [11] adopted an O-arm (Medtronic, Inc., Minneapolis, MN, USA) with navigation to guide TPS placement. They collected postoperative CT scans to evaluate screw accuracy in revision spinal surgery. Interestingly, for the purpose of decreasing radiation, confirmatory intraoperative CT scans were not routinely performed in their study, unless there was concern for screw malpositioning. There were six poor screws (1.4%, pedicle breach of > 3 mm) found postoperatively in their study; fortunately, none caused neurological deficits and, thus, were not revised. In our study, unaccepted screws were identified intraoperatively with the first confirmatory CT. After the secondary confirmatory CT, all the revised screws were accepted.

Since unaccepted screws may not always result in neurologic deficits, it might not be necessary to revise all of them intraoperatively. Laine et al. [14] reported that no screw caused neurological problems if the pedicle breach was less than 4.0 mm on the postoperative CT scan. Castro et al. [15] found that a TPS with medial pedicle breach of more than 6 mm created a high risk of nerve root damage. In the current study, the criteria to revise a screw was a medial pedicle breach of > 3mm or a lateral pedicle or body wall of >3 mm, which might be appropriate. However, we should keep in mind that even a medial 2 mm breach can be symptomatic, whereas a lateral breach of approximately 5 mm often remains asymptomatic in some extreme case. Although all unaccepted screws were revised intraoperatively and successfully, the use of a secondary confirmatory CT increased the patients' exposure to radiation. Intraoperative neuromonitoring was available in the current study, by which the authors could have avoided revising some of the unaccepted screws and rescued a TPS with a pedicle breach of <3 mm and abnormal neuromonitoring alerts.

In spinal surgery with posterior instrumentation, the free-hand technique combined with intraoperative fluoroscopy is commonly used to confirm the entry point and trajectory of the TPS. The reported incidence of misplaced screws ranged from 10.5% to 21% [14,16–18]. Two-dimensional (2D) fluoroscopy may not provide sufficient image information to safely insert TPSs in spinal deformities [12,17,18] due to a remodeled bony structure or spinal fractures [10,16] because of hypermobile spinal segments and disrupted anatomy, such as facets or pedicle fractures. Furthermore, it would be technically challenging to use fluoroscopy for TPS placement in revision spine surgery due to the altered anatomy, indistinct bony landmarks and excessive scar tissue [2]. CT scans can provide adequate image information to evaluate TPS position, including the axial, coronal and sagittal planes [14], and is, therefore, the better choice for guiding TPS placement in cases of compromised spine anatomy, such as deformity, trauma, and revision spine surgery.

In the last two decades, several advanced imaging techniques have emerged that are able to provide 3D images. Usually, 3D fluoroscopy is limited by a small "field-of-view" scan area and an inferior image quality, with a screw accuracy inferior to iCT [19]. The O-arm is a cone-beam computed tomography, which has higher rates for screws revised intraoperatively [20], but has a similar accuracy for postoperative screw placement as compared with iCT. Therefore, we believe that iCT is the imaging technology of choice to provide the optimal image resolution for TPS placement. However, it cannot provide real-time imaging to immediately determine the screw trajectory and may further expose the patients to a greater radiation dosage.

Certain navigation techniques enable the diminution of CT scans for guiding TPS placement, while providing virtual but real-time images to guide TPS placement. The iCTnavigation technique is an effective method for TPS placement in complex spinal surgery for patients with spinal deformities [12] and unstable thoracolumbar spine fractures, with a high TPS accuracy of up to 98% [9,10]. It also allows for the immediate revision of unaccepted TPSs during surgery, followed by a secondary confirmatory CT scan. This option is highly practical in the case of immediate intraoperative revision, because returning to another surgery for screw revision would increase the risk for the patient. Although the confirmatory CT scan exposes patients to radiation, the medical staff are protected from radiation in a lead-shielded room.

There were several limitations in our study. It might be relatively expensive to set up the iCT-navigation system; however, the installation charge should be considered against the cost of reoperation for patients with misplaced screws. We should have measured the dose of radiation to which patients were exposed, which might have been relatively high, since the entire workflow of iCT-navigation surgery contained at least two CT scans, including the registration scan and confirmatory scan. According to our previous studies [10] for primary surgery, the mean dose of patient radiation exposure was 15.8 mSv, and the mean dose per single level was 2.7 mSv. Fortunately, according to the recommendation by the International Commission on Radiological Protection, the annual maximum permissible dose is 20 mSv per year with no single year exceeding 50 mSv. Hence, the radiation dose to the patient in our study might have been within the limit. Furthermore, although this study was purely an imaging-based investigation and did not focus on functional outcomes, it was found that there were no immediate postoperative neurologic deficits.

#### **5. Conclusions**

The use of iCT-navigation for TPS placement in revision thoracolumbar spinal surgery was a useful technique for improving the accuracy and safety of TPS placement and allowed for the immediate intraoperative revision of unaccepted screws. The rate of accepted screws at the revision site was comparable to that at the virgin site in revision spinal surgery, although the rate of unaccepted screws at the revision site trended higher than at the virgin site. No postoperative neurological injury was noted in our patients.

**Supplementary Materials:** The following supporting information can be downloaded at: https: //www.mdpi.com/article/10.3390/jcm11195853/s1, Table S1: Demographic and Perioperative Data.

**Author Contributions:** Conceptualization, Y.-Y.L.; methodology, Y.-Y.L. and S.-H.C., software, C.-C.C. and C.-Y.L. (Chien-Yin Lee); validation, K.-C.H.; formal analysis, C.-Y.L. (Ching-Yu Lee) and M.-H.W.; supervision, T.-J.H. 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 approved by the ethics committee of the Chang-Gung Memorial Hospital, Taiwan (approval no. 202100339B0).

**Informed Consent Statement:** Informed consent from individual patients was waived because of the retrospective study with anonymized data.

**Data Availability Statement:** The original data generated during the study are included in the article. Further inquiries can be directed to the corresponding author.

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

#### **References**

**Tomasz Wach 1,\*, Małgorzata Skorupska <sup>1</sup> and Grzegorz Trybek <sup>2</sup>**

**Abstract:** The reason why marginal bone loss (MBL) occurs after dental implant insertion without loading has not yet been clearly investigated. There are publications that confirm or reject the notion that there are factors that induce marginal bone loss, but no research investigates what exactly occurs in the bone surrounding the implant neck. In this study, 2196 samples of dental implant neck bone radiographs were analyzed. The follow-up period was 3 months without functional loading of the implant. Marginal bone loss was evaluated in relation to the torque used during the final phase of implant insertion. Radiographic texture features were also analyzed and evaluated. The analyses were performed individually for the anterior and posterior part of the alveolar crest in both the mandible and maxilla. After 3 months, an MBL relation with higher torque (higher than 40 Ncm; *p* < 0.05) was observed, but only in the lower jaw. The texture features Sum Average (SumAverg), Entropy, Difference Entropy (DifEntr), Long-Run Emphasis (LngREmph), Short-Run Emphasis (ShrtREmph), and discrete wavelet decomposition transform features were changed over time. This study presents that MBL is related to the torque value during dental implant insertion and the location of the procedure. The increasing values of SumAverg and LngREmph correlated with MBL, which were 64.21 to 64.35 and 1.71 to 2.01, respectively. The decreasing values of Entr, DifEntr, and ShrtREmph also correlated with MBL, which were 2.58 to 2.47, 1.11 to 1.01, and 0.88 to 0.84, respectively. The analyzed texture features may become good indicators of MBL in digital dental surgery.

**Keywords:** dental implant; torque; marginal bone loss; intraoral radiographs; radiomics; texture analysis; bone remodeling

#### **1. Introduction**

Presently, the most common procedure in oral surgery after wisdom tooth extraction are dental implants [1]. With the increasing number of dental implant placements, more and more post-operative complications also occur. One of the major complications is marginal bone loss (MBL) next to the dental implant neck [2]. Marginal bone loss is a condition where the bone surrounding the implant neck atrophies. It is affected by different factors, e.g., smoking, diabetes mellitus, vitamin D and 25-hydroxycholecaliferol level, implant placement technique, region of the jaw, and also torque during the surgical procedure [3–6]. MBL may occur after few years, but also after the first 3 months of healing. The osteotomy techniques used for implant placement may also differ and may affect MBL [7]. MBL that occurs after a few years may be a condition related to more factors, e.g., prosthetic restoration and gingival vertical and horizontal width. MBL may be associated with a high torque value [8,9] if the correct procedure steps have not been followed and the appropriate primary stability of the implant has occurred [10].

The visual assessment of the surrounding implant neck bone on radiographs may not be sufficient or reliable. Another way to evaluate radiographic images is to check how the radiographic textures change over time—the shade level of the pixels can be analyzed.

**Citation:** Wach, T.; Skorupska, M.; Trybek, G. Are Torque-Induced Bone Texture Alterations Related to Early Marginal Jawbone Loss? *J. Clin. Med.* **2022**, *11*, 6158. https://doi.org/ 10.3390/jcm11206158

Academic Editors: James Kit-Hon Tsoi, Michał Strzelecki, Rafał Obuchowicz, Andrzej Urbanik and Adam Piorkowski

Received: 7 August 2022 Accepted: 17 October 2022 Published: 19 October 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/).

This analysescan be used, for example, to check how the texture of the bone substitute materials have changed after several months, translating this change into healing process progress [11–14].

The first aim of this study is to check whether torque and the location of the implant have an influence on MBL and to determine whether there is a change of radiographic texture in the bone surrounding the implant neck. The second aim is to check how the radiographic texture near the implant neck changes over time and to determine the prognostic factors of MBL in image texture.

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

In this study, 2196 samples of neck area implants were included and analyzed. A total of 504 males and 496 females aged between 15 and 86 years old were included in the study. All patients had undergone the same surgical procedure, namely, bone-level dental implant placement under local anesthesia (4% articaine with 1:100,000 adrenaline, 3 M ESPE AG, Seefeld, Germany). The patients were divided into two groups depending on whether MBL occurred after 3 months or not, and also into a mandible and maxilla group where the anterior and posterior parts were distinguished:

#### MBL appearance (YES) if MLB is >0 MBL appearance (NO) if MBL is =0

The inclusion criteria were two-dimensional radiographs taken immediately after surgery and 3 months later, the measurement of torque value immediately after dental implant placement, and laboratory tests to check patients' vitamin levels, ions, and hormones: parathormone (PTH), where the norm is 10 to 60 pg/mL; thyrotropin (TSH), where the norm is 0.23–4.0 μU/mL; calcium in serum (Ca2+), where the norm is 9–11 mg/dL; glycated hemoglobin (HbA1c), where the norm is <5%; and vitamin 25(OH)D3 (D3), where the norm is 31–50 ng/mL. Spine densitometry, where the T-score can be examined, was also considered. The T-score shows the ratio between the bone mineral density (BMD) of the examined patient and the average BMD for young patients. A normal value for normal bone is >−1.0, osteopenia is indicated by values between −1.0 and −2.5, and scores < −2.5 indicate osteoporosis. The exclusion criteria included a lack of X-rays, defective X-ray images in the visual assessment, lack of a clear torque value, and lack of laboratory tests. In this study, only patients with proper values from the laboratory tests were included.

Surgery was done under local anesthesia, Septanest + A 1:100.00, by one surgeon according to the recommended protocols. The healing process was carried out under a closed mucoperiosteal flap, unloaded in two-stage implants. The thickness of the soft tissue did not affect the healing process or MBL in the first stage of healing. Table 1 presents the implants used in this study and their technical features. The data were confirmed at www.spotimplant.com/en/dental-implant-identification, accessed on 5 March 2022.

Two-dimensional X-ray images were taken immediately after surgery (00M) and 3 months later (03M). Radiographs were taken using the DIGORA OPTIME radiography system (TYPE DXR-50, SOREDEX, Helsinki, Finland). The radiographs were taken in the standardized way [15] with the following parameters: 7 mA, 70 mV, and 0.1 s (the focus apparatus was from Instrumentarium Dental, Tuusula, Finland). Positioners were used to take images repeatedly, with the X-ray beam at a 90◦ angle to the surface of the phosphor plate. The texture of the X-ray images was analyzed in the MaZda 4.6 software, developed by the University of Technology in Łód´z, Poland [16], to check how the features changed over the 3 months of observation. A limitation of the study is that the laboratory tests were not checked after 3 months.

The analyses were performed in a few steps: first, all of the X-rays were edited (leveled) (Figure 1); next, the MBL near the implant neck area was measured (Figure 2), and then the X-ray image was loaded into MaZda in a bitmap file format. Next, the region of interest (ROI) was marked near the neck on the mesial and/or distal side of the implant (5–6 mm height) (Figure 3a). The ROI was marked on the RTG image immediately after inserting the dental implant and after 3 months of healing (Figure 3b). Any bone loss after 3 months was evaluated through radiographic analysis, as the vertical differences between the implant platforms and the first bone contact with the implant surface. The ROIs were normalized (μ ± 3σ) to share the same average (μ) and standard deviation (σ) of optical density within the ROI. The selected image texture features—sum of squares (SumOfSqrs), sum of average (SumAverg), entropy, different entropy (DifEntr), long-run emphasis moment (LngREmph), and short-run emphasis moment (ShrtREmph)—in the ROIs were calculated for the reference bone and for the bone near the implant neck. The Haar wavelet decomposition (LH, HL, LL, HH) was also performed and statistically analyzed after 3 months of observation. All features were gathered from four angles: 0◦, 45◦, 90◦, and 135◦ from done pixel and the average value was later calculated.

$$SumAverg = \sum\_{i=1}^{2N\_\\$} ip\_{x+y}(i)$$

$$SumOfsqrs = \sum\_{N\_\S}^{N\_\S} \cdot \sum\_{j=1}^{N\_\S} (i - \mu\_x)^2 \ p(i, j)$$

$$Entropy = -\sum\_{i=1}^{N\_\S} \sum\_{j=1}^{N\_\S} p(i, j) \log(p(i, j))$$

$$DiffEntr = -\sum\_{i=1}^{N\_\S} p\_{x-y}(i) \log(p\_{x-y}(i))$$

where Σ is sum, *N* is the number of levels of optical density in the radiograph, *i* and *j* are the optical density of pixels five-image-point distant one from another, *p* is probability, and log is logarithm [11].

$$LngREmph = \left(\sum\_{i=1}^{N\_\mathcal{S}} \sum\_{j=1}^{N\_r} j^2 p(i, j)\right) / \mathcal{C}$$

$$ShrtRemph = \left(\sum\_{i=1}^{N\_\mathcal{S}} \sum\_{j=1}^{N\_r} j^{-2} p(i, j)\right) / \mathcal{C}$$

where Σ is sum, *N* is the number of series of pixels with density level *i* and length *j*. *Ng* is the number of levels for image density (8 bits, i.e., 256 gray levels), *Nr* is the number of pixels in series, *p* is the probability, and *C* is the coefficient, as below:

$$\mathcal{C} = \sum\_{j=1}^{N\_r} \sum\_{i=1}^{N\_g} p(i, j)$$

#### *Statistical Analysis*

The Kruskal–Wallis test (to compare time-dependent alternations in medians) was applied for statistical analysis. Next, a multiple comparison procedure was used to determine which means were significantly different from the others. The method discriminates among the variables Fisher's least significant difference (LSD) procedure. The difference was considered significant if *p* < 0.05. Stargraphics Centurion XVI (Statgraphics-StatPoint Technologies, Inc., The Plains, VA, USA) was used for the statistical analyses.

**Figure 1.** Geometrical alignment of radiograph image. The green line marked on the implants indicates the long axis of the implant.

**Figure 2.** Measuring of marginal bone loss on the radiographic images. Red lines indicate the implant platform to the bottom of the bone loss cavity.

(**a**)

**Figure 3.** (**a**) Marking a ROI. ROIs were marked near the implant neck area. Green area—mesial implant neck area; red area—distal implant neck area; blue area—reference bone. Abbreviations: ROI—region of interest. (**b**) Marking a region of interest on RTG image immediately after inserting the implant and 3 months after the first stage of the healing process. Green area—mesial implant neck area; red area—distal implant neck area. At the bottom of the marked area on the right (03M), it can be noticed that MBL occurred and is analyzed. Abbreviations: MBL—marginal bone loss; 00M—0 months of observation; 03M—3 months of observation.

#### **3. Results**

#### *3.1. Marginal Bone Loss Statistical Evaluation*

The statistical evaluation revealed that the amount of marginal bone loss was between 0 and 8.05 mm after 3 months of observation; the average MBL for the mandible was 0.29 ± 0.98 mm and for the maxilla was 0.23 ± 0.91 mm, which were both statistically significant at *p* < 0.01. The MBL means for the anterior and posterior parts of the jaw were 0.32 ± 1 mm and 0.22 ± 0.91 mm, respectively, where *p* was lower than 0.01, which means that that was statistically significant.

#### *3.2. Torque Statistical Evaluation*

It was also noticed that when MBL occurred, higher torque was observed in the mandible group (mean 46.77 Ncm ± 14) than in the maxilla group (mean 40.5 Ncm ± 11.9), with statistical significance (*p* < 0.05); the torque value during the implantation procedure was between 5 Ncm and 90 Ncm (Table 2).

**Table 2.** Average values for marginal bone loss and torque. Values are presented for mandible, maxilla, and for anterior and posterior areas of the jaw. Since the *p*-value is greater than or equal to 0.05, there is no statistically significant relation.

Abbreviations: *p*—the probability of obtaining test results at least as extreme as the results actually observed, under the assumption that the null hypothesis is correct.

#### *3.3. Torque Value and Marginal Bone Loss as a Dependency*

Marginal bone loss is related to the torque value during the implant placement (CC = 0.06, R2 = 0.3%, *p* < 0.05). If the torque increases, then MBL occurs more often (Figure 4). The study also showed that the average torque in the group with detected bone loss was higher than in the group without bone loss (42.01 and 40.04, respectively, where *p* was lower than 0.05, which means that was statistically significant) (Figures 5 and 6).

**Figure 4.** Dependence of marginal bone loss appearance from torque value after the dental implant insertion after 3 months of observation. The greatest number of marginal bone loss samples occurred in implants with a torque value equal to 45 Ncm. The higher the torque, the higher the MBL occurrence (with statistical significance). Abbreviations: MBL—marginal bone loss.

**Figure 5.** Dependence of average torque to marginal bone loss appearance after 3 months of observation. The mean where marginal bone loss was detected was higher than 42 Ncm. When the torque did not have an influence on MBL, the average torque was 40 Ncm or lower. Abbreviations: MBL—marginal bone loss.

**Figure 6.** Figure presenting two radiographs. On the left is the radiograph image taken immediately after dental implant placement at 00M. The level of the bone at the beginning of the observation is marked by a green line. On the right is the RTG at 3 months of observation after implant placement. The red lines mark the marginal bone loss in relation to the previous bone level. Abbreviations: RTG—radiograph image; 00M—0 months of observation.

#### *3.4. Implant Placement Region in Jaw*

After splitting the implant samples into two groups, the mandible and the maxilla group, the statistical evaluation revealed that marginal bone loss only correlated with torque in the mandible group (*p* < 0.01), even though most samples with MBL in the maxilla group occurred with a torque higher than 40 Ncm (Figure 7). There was no correlation between the groups where augmentation was performed before implant placement (*p* > 0.05).

**Figure 7.** The graph on the left shows the dependence of marginal bone loss from dental implant insertion placement torque in the upper jaw (maxilla). There was no statistically significant difference (*p* > 0.05). The graph on the right presents the dependence of marginal bone loss on dental implant insertion placement torque in the lower jaw (mandible). There was statistical significance (*p* < 0.05).

Not all of the analyzed texture features were statistically significantly related to the marginal bone loss (*p* < 0.05). A *p*-value lower than 0.05 was observed for the following features after 3 months of observation:

Texture features where the *p*-value was higher than 0.05, which means that this is not statistically significant, were:

As the reference, textural features for the trabecular bone were analyzed. The basal values for SumAverg, Entropy, DifEntr, LngREmph, and ShrtREmph are presented in Table 3.

**Table 3.** Reference texture feature values. Since the *p*-value is greater than or equal to 0.05, there is no statistically significant relation.

Abbreviations: *p*—the probability of obtaining test results at least as extreme as the results actually observed, under the assumption that the null hypothesis is correct.

#### *3.5. SumAverg Changes*

The SumAverg for the implant neck area was 64.21 ± 2.9 at 00M. After 3 months, the SumAverg for the implant neck area with MBL was higher than after implantation, presenting 64.35 ± 3.54 and 64.16 ± 3.85, where MBL was not detected, and this was

statistically significant (*p* < 0.05). There was no statistical difference between the maxilla and mandible groups (*p* > 0.05).

#### *3.6. Entropy Changes*

The Entropy for the implant neck area was 2.58 ± 0.19, and after implantation changed to 2.47 ± 0.21 where MBL was detected and 2.52 ± 0.2 where MBL was not detected, which was statistically significant. Statistically significant (*p* < 0.01) changes were also noticed in the maxilla and mandible groups with MBL (2.58 ± 0.12 changed to 2.52 ± 0.16, *p* > 0.05; and 2.58 ± 0.14 changed to 2.42 ± 0.16, *p* < 0.01, respectively).

#### *3.7. DifEntr Changes*

The DifEntr after implantation at 00M near the implant neck area was 1.11 ± 0.16 and changed to 1.01 ± 0.15 with MBL, and to 1.04 ± 0.16 where MBL was not present (*p* < 0.05). The DifEntr in the mandible group with MBL was 1.07 ± 0.15 and changed to 0.95 ± 0.14, which was statistically significant (*p* < 0.05).

#### *3.8. LngREmph Changes*

The LngREmph value at 00M was 1.71 ± 0.57 and changed to 2.01 ± 0.55 in the area with MBL, and 1.97 ± 0.75 where MBL did not appear (*p* < 0.05). The LngREmph value in the mandible group changed from 1.74 ± 0.64 to 2.11 ± 0.75 and there was no statistically significant difference (*p* > 0.05).

#### *3.9. ShrtREmph Changes*

The ShrtREmph also changed significantly: at 00M it was 0.88 ± 0.05, and after 3 months changed to 0.84 ± 0.05 for the area where MBL appeared and 0.85 ± 0.06 for the implant neck area without MBL. The ShrtREmph in the mandible group where MBL was correlated with MBL changed from 0.88 ± 0.6 to 0.84 ± 0.06, which was not statistically significant (*p* > 0.05).

#### *3.10. Wavelet Decomposition Changes*

The value for WavEnLH\_s-4 after implantation for the maxilla group was 142.92 ± 111.54 and changed to 118.04 ± 87.24, which was statistically significant (*p* < 0.01). In the mandible group with MBL, the WavEnLH\_s-4 changed from 134.35 ± 86.67 to 112.18 ± 116.14, which was statistically significant (*p* < 0.05).

The WavEnLH\_s-5 in the mandible group changed from 345.23 ± 203.47 to 212.15 ± 185.22, which was statistically significant (*p* < 0.05). The WavEnHH\_s-5 in the mandible group statistically significantly changed from 61.28 ± 65.77 to 35.75 ± 38.44, where the *p*-value was lower than 0.05 (*p* < 0.05).

Interestingly, the Wavelets 6 disappeared after 3 months of observation in the whole group with marginal bone loss (in the mandible and maxilla groups). The Wavelets 6 were also not detected in the reference trabecular bone. This texture index may be an indicator for cortical bone or changes in structure during the observation (Tables 4–6).

**Table 4.** Texture feature values at 00M, 03M, and for reference trabecular bone. Since the *p*-value is greater than or equal to 0.05, there is no statistically significant relation.

Abbreviations: 00M—0 months of observation; 03M—3 months of observation, MBL—marginal bone loss; *p*—the probability of obtaining test results at least as extreme as the results actually observed, under the assumption that the null hypothesis is correct.

**Table 5.** Texture feature values for mandible and maxilla groups after 3 months.

Abbreviations: 03M—3 months of observation; *p*—the probability of obtaining test results at least as extreme as the results actually observed, under the assumption that the null hypothesis is correct; MBL—marginal bone loss.

**Table 6.** Texture feature values for mandible and maxilla groups on the day of surgery.

Abbreviations: 00M—the day of surgery; *p*—the probability of obtaining test results at least as extreme as the results actually observed, under the assumption that the null hypothesis is correct. MBL—marginal bone loss.

The research also showed that implant design also has an impact on marginal bone loss near the implant neck. This study compared several design properties: insertion implant level, neck microthreads, body shape, and body threads (thread shape). Apex shape, apex hole, apex groove, and connection type were not taken into account, as in the authors' opinion, these implant features do not have an impact on marginal bone loss in the early period of healing without exposing the implant for oral cavity conditions.

The statistical evaluation showed that MBL in the case of bone-level implants was 0.26 ± 0.97 mm, and for subcrestal implants was 0.09 ± 0.51 mm, which was statistically significant (*p* < 0.05).

Greater MBL occurred near the neck of implants without microthreads (0.31 ± 0.92 mm) than near the implants where microthreads were present (0.25 ± 0.94 mm), and this was statistically significant at *p* < 0.05.

The marginal bone loss for implants without body threads was higher (0.99 ± 0.77 mm) than for the implants where the body threads were V-shaped (0.15 ± 0.64 mm). MBL was also shown in the case of square-, buttress-, and reverse buttress-shaped body threads at 0.28 ± 0.93 mm, 0.67 ± 1.75, and 0.25 ± 0.97 mm, respectively. Statistical significance at *p* < 0.05 was noticed between the reverse buttress and no threads; between square threads and no threads; and between V-shaped threads and no threads.

The research also showed that the MBL in the case of straight-body implants was 0.20 ± 0.49 mm and in the case of tapered implants was 0.25 ± 0.95 mm. There was no statistical significance, as *p* > 0.05 (Table 7).

Taking into account torque as a factor that can lead to MBL, some implant design features were examined depending on the insertion region and the torque used (higher or lower than 45 Ncm). Titanium alloy, the level of implant placement, the presence of microthreads on the implant neck, implant body shape, and also the design of the threads on the implant body were checked.

**Table 7.** Comparison of marginal bone loss depending on selected implant design features.

Abbreviations: MBL—marginal bone loss; *p*—the probability of obtaining test results at least as extreme as the results actually observed, under the assumption that the null hypothesis is correct.

It was noticed that significantly higher MBL occurred in the anterior part of the mandible when the torque was lower than 45 Ncm, where dental implants made of titanium alloy Grade 4 (mean 1.06 ± 0.94 mm) were used. There was no correlation between MBL and titanium alloy in the maxilla samples. In the rest of the samples, titanium alloy was not correlated with MBL due to the localization of the implant or the insertion torque.

The research shows that the level of implant placement has an impact in the case of maxilla and mandible implants. In the anterior of the maxilla, greater MBL was noticed according to the tissue implant level (0.97 ± 0.88 mm) and the smallest MBL was found where subcrestal implants were used (mean 0.24 ± 1.14 mm), but only in cases when the torque was higher than 45 Ncm. In mandible-inserted implants, a subcrestal location had an impact on MBL in cases where the torque was lower than 45 Ncm (mean 1.06 ± 2.14 mm). The level of implant placement had no influence in the maxilla, either for the anterior part with a torque lower than 45 Ncm or the posterior part regardless of the torque used; and for the mandible, the anterior part with a torque higher than 45 Ncm and the posterior part regardless of the torque used.

Higher MBL was noticed in the anterior part of the maxilla when torque higher than 45 Ncm was used for implants without microthreads (mean 0.82 ± 1.76 mm). The presence or absence of neck microthreads in the case of the mandible, the posterior part of the maxilla, and the anterior part of the maxilla with torque lower than 45 Ncm was not correlated with MBL.

The shape of the implant body was not correlated with MBL in the case of maxilla and mandible dental implants regardless of the torque used during the insertions.

The research also showed that in maxilla implants with buttress threads inserted in the anterior part with a torque higher than 45 Ncm, and in the anterior part of the mandible where implants with V-shaped threads on the implant body were inserted with a torque lower than 45 Ncm, the MBL was greater and was statistically significant (mean 2.19 ± 3.34 mm and 1.06 ± 2.24 mm, respectively).

#### **4. Discussion**

The question is: Does the torque value of dental implants affect early marginal bone loss after 3 months of healing? There are publications about crestal bone stability claiming that there is no relation between implant torque insertion value and marginal bone loss [17]. In this study, 2196 samples of implant neck areas were analyzed and proved that there is a statistical relationship between a torque higher than 40 Ncm and marginal bone loss after 3 months of healing. Due to the lack of prosthetic loading, it can be declared that high torque during implant insertion is the main surgeon-related factor of MBL near the implant neck after 3 months of healing. Additionally, it was presented that dental implants inserted with a torque higher than 40 Ncm in the lower jaw were more susceptible to marginal bone loss, even though there are studies indicating that there is no correlation between marginal bone loss in the maxilla vs. the mandible [18,19].

The radiological or clinical evaluation of marginal bone loss is not always possible using only visual assessment. Sometimes, changes in the morphological part of the bone or bone substitute materials are not visible. Tomasz Wach et al. and Kozakiewicz et al. show that changes overcome during the healing process can be detected using texture features, already on the level of pixels [12,14,20]. Some of the texture features (inter-pixel relation in the optical density environment) can also be a good indicator of bone structure changes near the implant neck after 3 months of healing.

SumAverg and Entropy change over 3 months. It can be noticed that the decreasing value of SumAverg in cases where MBL was not detected can be a sign that the tissue around the dental implant neck is similar to intact trabecular bone tissue. The increasing value of Entropy also shows that in the group without MBL, the tissue around the dental implant neck after 3 months is more similar to trabecular bone. Referring to the other publications, it can also mean that the texture feature values of tissue near the implant neck after 3 months of healing in the group where MBL was detected approach cortical bone texture feature values [13,14].

The decreasing values of DifEntr and increasing values of LngREmph texture features are further proof that the tissue around the implant neck in the group with marginal bone loss is not similar to the trabecular one. For confirmation, Kozakiewicz et al. also proved that increasing LngREmph texture feature values and decreasing values of DifEntr are one of the signs of corticalization, which is correlated with bone loss [21]. In cases where MBL was detected and was correlated with corticalization, more longitudinal objects were observed (increasing LngREmph) as well as a decrease in chaotic patterns (decreased DifEntr).

Texture feature values that become similar to the reference of cortical bone can be an indicator of MBL. The process called corticalization near the implant neck surface is related to MBL [21]. On the other hand, an increasing value of the Entropy feature can be one of the signs that chaotic patterns have increased. Greater entropy means greater chaos, which can be equal to the decrease in bone structure.

It can be noticed that the values of the texture features in the group where marginal bone loss appeared were similar to the texture features of the reference scale for cortical bone. It is also related to higher torque when placing an implant. This means that higher torque during implant placement may lead to the densification of tissue around the implant neck. Condensed bone can be the first step to cortical bone formation, which is correlated with MBL.

Wavelet decomposition may be a healing process indicator. It can show that the trabecular part of the bone becomes part of the structure surrounding the dental implants. It also indicates the changes of long and small objects: longitudinal or circular shades [14,22]. It was noticed in Wach and Kozakiewicz's study [14] that the higher the scale of wavelet decomposition, the larger the object appears in the texture. The subbands HH and LL indicate circular shades, while LH and HL indicate longitudinal ones. The decrease in wavelets in scale 4 and 5 in subband LH in the mandible group where MBL was detected may be related to the disappearance of bone structure. The disappearance of the decomposition of wavelets in scale 6 in our research was interesting. Our research can lead to the conclusion that the wavelets in scale 6 are an indicator of the healing process that occur later than 3 months post-operation. Wavelets in scale 6 should be observed in the next observation period.

The subcrestal implant level placement looks promising. This study shows that the lowest MBL appeared near this kind of implant. It is likely that the level of implant placement is not the only factor to have an impact on MBL. There are also studies that show that implant level placement does not have an influence on MBL near the implants [6]. MBL before the loading and exposure of the implant may be a result of different stresses of the implant neck.

Implant design features and structural features may also have an impact on marginal bone loss regardless of the torque used or the region of insertion [23,24], or may have preservation influence on the bone around the implants [25]. Taking into consideration that torque was not correlated with marginal bone loss in the case of implants in the maxilla, the research shows a correlation between MBL and implant design in the anterior part of the maxilla where the torque was higher than 45 Ncm. This proves that marginal bone loss in the maxilla is not correlated with a high insertion torque, or high torque along with some other implant feature can result in greater MBL. In the implant samples analyzed in the lower jaw, dental implant design had an insignificant impact on MBL in cases where the torque was lower than 45 Ncm. This is further proof that insertion torque (higher than 45 Ncm) has an impact on MBL near the implant neck after the first 3 months of healing.

One limitation of the study is that the laboratory tests were not checked after 3 months. Another limitation is that the radiograph texture analyses were not compared to the histopathological examination of bone near the implant neck. Clinical marginal bone loss in mm was not carried out because of trauma—the authors checked the marginal bone loss before the second stage of treatment—before exposing the implant. There was also a limited number of samples, as not all patients came back after 3 months of healing and not all of the pictures taken were qualified after visual assessment for analyses. The BMI of the patients was also not taken into account. The research needs further evaluation of other local factors that could have an impact on marginal bone resorption.

#### **5. Conclusions**

Marginal bone loss is related to higher torque value during the implant placement procedure (higher than 40 Ncm torque was closely associated with MBL), but only in implants located in the lower jaw. The texture feature values that change over the healing process are closely related to the occurrence of MBL. It can be concluded that there are some texture features that can be used as indicators of problems near the inserted implants. The texture feature values (SumAverg, Entropy, DifEntr, and LngREmph) can indicate the likelihood of MBL occurrence. Haar wavelet decompositions should be observed in the next period of observation. Selected implant design features near the implant neck may have a positive impact on marginal bone, lower early bone loss associated with the neck microthreads and body threads. The level of implant placement also has an effect on early MBL. Additionally, the statistics verified the correlation between the design of the implant, insertion region, and torque, showing that cases where the implants were inserted with higher torque were more vulnerable to MBL near the implant neck in the early period of healing. This study is also the beginning of research into the correlation between texture features near the implant neck on the day of surgery and future bone loss.

**Author Contributions:** Conceptualization, T.W.; data curation, M.S. and T.W.; formal analysis T.W.; funding acquisition, T.W.; investigation, M.S. and T.W.; methodology, T.W.; resources, T.W.; software T.W.; supervision, G.T.; validation, T.W.; visualization, G.T. and T.W.; writing—original draft, T.W.; writing—review and editing, T.W. and G.T. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research was funded by the Medical University of Lodz (grant numbers 503/5-061- 02/503-51-001-18, 503/5-061-02/503-51-001-17, and 503/5-061-02/503-51-002-18).

**Institutional Review Board Statement:** The study was conducted according to the guidelines of the Declaration of Helsinki and approved by the Institutional Ethics Committee of the Medical University of Lodz, PL (protocol no. RNN 485/11/KB and date of approval: 14 June 2011).

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

**Data Availability Statement:** The data on which this study is based will be made available upon request at https://www.researchgate.net/profile/Tomasz-Wach, access date 16 October 2022.

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

#### **References**


*Review* **Modification of Cellulose Micro- and Nanomaterials to Improve Properties of Aliphatic Polyesters/Cellulose Composites: A Review**

**Mariia Stepanova and Evgenia Korzhikova-Vlakh \***

Institute of Macromolecular Compounds, Russian Academy of Sciences, Bolshoy pr. 31, 199004 St. Petersburg, Russia<EMAIL_ADDRESS>
**\*** Correspondence<EMAIL_ADDRESS>
**Abstract:** Aliphatic polyesters/cellulose composites have attracted a lot attention due to the perspectives of their application in biomedicine and the production of disposable materials, food packaging, etc. Both aliphatic polyesters and cellulose are biocompatible and biodegradable polymers, which makes them highly promising for the production of "green" composite materials. However, the main challenge in obtaining composites with favorable properties is the poor compatibility of these polymers. Unlike cellulose, which is very hydrophilic, aliphatic polyesters exhibit strong hydrophobic properties. In recent times, the modification of cellulose micro- and nanomaterials is widely considered as a tool to enhance interfacial biocompatibility with aliphatic polyesters and, consequently, improve the properties of composites. This review summarizes the main types and properties of cellulose micro- and nanomaterials as well as aliphatic polyesters used to produce composites with cellulose. In addition, the methods for noncovalent and covalent modification of cellulose materials with small molecules, polymers and nanoparticles have been comprehensively overviewed and discussed. Composite fabrication techniques, as well as the effect of cellulose modification on the mechanical and thermal properties, rate of degradation, and biological compatibility have been also analyzed.

**Keywords:** microcrystalline cellulose; nanocrystalline cellulose; cellulose fibers; cellulose modification; aliphatic polyesters; polyhydroxyalkanoates; poly(lactic acid); poly(ε-caprolactone); poly(glycolic acid); poly(lactic acid-co-glycolic acid); poly(hydroxybutyrate); poly(butylene succinate); (bio)composites; "green" materials; mechanical properties; thermal properties; degradation; biocompatibility

### **1. Introduction**

In recent decades, aliphatic polyesters have attracted enormous interest as an alternative to plastics derived from petroleum [1]. Aliphatic polyesters are biocompatible, biodegradable, and have an excellent ability to a number of processing techniques allowing the production of electrospun nanofibers, films, filaments, nonwoven materials, 3D-printed materials of different shapes, molded and pressed materials, nanocomposite bulk materials, etc. [2–4]. Degradation to nontoxic products, the possibility of recycling, thermoplasticity, nontoxicity, comparability of some parameters with poly(ethylene terephthalate) (PET) and polypropylene (PP) [5–7], low flammability, smoke and refractive index, and dyeability [8] are among other positive features of aliphatic polyesters. In sum, these advantages make aliphatic polyesters very attractive polymers for obtaining biomedical (drug-delivery systems, suture threads, scaffolds for tissue engineering, etc.) [5,9–11] and environmentally friendly materials (packaging and disposable items such as clothing, tableware, etc.) [8,12,13]. However, their high hydrophobicity, insufficient thermal stability, and mechanical and barrier properties limit their wide application for technical purposes. The most powerful way to modify the properties of aliphatic polyesters is to obtain various

**Citation:** Stepanova, M.; Korzhikova-Vlakh, E. Modification of

Cellulose Micro- and Nanomaterials to Improve Properties of Aliphatic Polyesters/Cellulose Composites: A Review. *Polymers* **2022**, *14*, 1477. https://doi.org/10.3390/ polym14071477

Academic Editor: Debora Puglia

Received: 28 February 2022 Accepted: 31 March 2022 Published: 5 April 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/).

composites [14]. In this case, the properties of the matrix polymer can be adjusted by the selection of a certain filler. For example, metals [15], carbon nanotubes [16], graphene [17] and its derivatives [18], ceramics [19], and different organic nanoparticles [20–22] are considered to improve the properties of interest.

Despite the variety of potential fillers, the most attention is paid to micro- and nanomaterials that are nontoxic and inexpensive, which makes it possible to produce "green" biocomposites on an industrial scale. Cellulose micro- and nanomaterials are among the most potential fillers for producing such environmentally friendly and biocompatible composites [23–26]. Excellent mechanical properties, a large specific surface area of cellulosic materials, the possibility to obtain them from the wastes of various industries, as well as biodegradability and biocompatibility make their application as reinforcing materials for a variety of areas, including biomedicine and obtaining "green" materials, relevant. However, the hydrophilicity of cellulose impairs significantly its dispersion in hydrophobic polyesters, which leads to cellulose aggregation, poor dispersion in the matrix polyester, and as a consequence, unsatisfactory material properties [27,28]. This obstacle can be overcome by modifying cellulose materials to improve their compatibility with hydrophobic polymer matrices, and as a result, to provide a more homogeneous dispersion of the filler.

In recent years, much attention has been paid to improving the compatibility of cellulose with aliphatic polyesters by covalent and noncovalent modification [29–38]. Modification of the cellulose surface, in turn, affects the properties of the cellulose filler and allows the properties of aliphatic polyester/cellulose composites to be adjusted in a wide range. Recently, several review articles devoted to the composites based on poly(lactic acid) and cellulose [39,40] with special focusing on the processing techniques [41,42] and biofiber's properties [43] have been published. Some reviews have partially discussed cellulose modification [39,44]; however, progress in this area has not been extensively overviewed.

In this comprehensive review, we have summarized the progress on the various approaches reported for the modifications of cellulose micro- and nanomaterials and the further preparation of composites with aliphatic polyesters. The different techniques such as adsorption, covalent modification with small molecules, grafting with polymers, and modification with inorganic and organic nanoparticles have been discussed. Unlike most reviews that consider only poly(lactic acid) (PLA), we have also included other aliphatic polyesters used to produce composites with modified cellulose, e.g., poly(glycolic acid), poly(ε-caprolactone), poly(hydroxybutyrate), poly(butylene succinate) and their copolymers. Furthermore, the effect of modification on various properties of composites, such as mechanical, thermal, degradation and biological ones, have been analyzed.

#### **2. Cellulose Micro- and Nanomaterials**

It is known that cellulose is the most abundant polysaccharide on our planet. Its main sources are primarily plants, including wood, as well as algae, tunicate, and metabolic products of some bacteria [45,46]. The highest cellulose content (more than 90%) is characteristic of "bacterial cellulose" (BC), while for other sources this value does not exceed 80% (plant—30–80%, tunicate—about 60%, algae—8–47%) [46–49]. The exception is mature cotton fibers, which consist almost entirely of cellulose (88.0–96.5%) [50]. Accordingly, BC and mature cotton are characterized by fewer impurities, such as lignin and hemicellulose, which are present in large amounts in plant and algae cellulose [46,50]. Another feature of BC is the presence of a finer mesh structure [46]. Furthermore, the degree of crystallinity for cellulose from different sources also varies quite a lot. Regardless of the source, cellulose is a linear homopolysaccharide and consists of β-1,4-glycosidic bonded anhydro-D-glucose units [49,51]. A large number of Van der Waals and hydrogen interactions are formed between and within the polymeric cellulose chains, which lead to the formation of three-dimensional hierarchical structures, the structural unit of which is an elementary fibril [52]. Elementary fibrils, also called elementary nanofibrils, are threadlike bundles of cellulose molecules consisting of alternating crystalline and a number of amorphous domains providing fiber flexibility [26]. Elementary fibrils due to aggregation are packed

into microfibrils, which, in turn also aggregate, and this leads to the formation of cellulose fiber [53]. The widths of elementary fibrils and microfibrils range from 1.5 to 5 nm [54,55] and from 10 to 30 nm [49,54], respectively, and the width and length of microfibril aggregates can reach the order of 100 nm and tens of micrometers [49,51,53,55], respectively. The large number of hydroxyl groups (three reactive groups in each monomeric unit) and the supramolecular structure of cellulose determine its physical and chemical properties (insolubility in water and basic solvents, semicrystallinity, good mechanical properties, relative reactivity) [26,55–57].

Depending on the origin and the method of isolation, the degree of polymerization (DP) of cellulose and the molecular orientation of its chains can be different. For native cellulose, the most common crystalline structure is cellulose I, which under the influence of sodiumhydroxide solution or recrystallization changes to the most stable crystalline state, cellulose II. More details about the different forms of cellulose can be found elsewhere [49,51,55,58,59]. The degree of polymerization of cellulose ranges from a few hundred to several tens of thousands [46,49,58]. Given the structure of cellulose, cellulose objects can be produced as fibers, micro/nanofibrils, and micro/nanocrystals [26,44], which vary in degree of polymerization, crystallinity, and shape [60]. Figure 1 schematically demonstrates the general hierarchical structure and structure of a single polymer chain of cellulose with a list of the main cellulose micro- and nanomaterials obtained.

**Figure 1.** Cellulose from source to molecule and micro- and nanomaterials.

The nomenclature used to designate the various types of micro- and nanocellulose materials is currently ambiguous. Thus, cellulose nanocrystals (CNC) are called nanocrystalline cellulose (NCC), cellulose nanowhiskers (CNW), cellulose whiskers, nanocrystals, nanofibers, nanoparticles, nanorods, rod-like cellulose crystals, cellulose microcrystals (CMC), cellulose microcrystallites, cellulose microfibrils (CMF) [26,39,51,55]; cellulose microfibrils (CMF) are called microfibrillated cellulose (MFC), microfibrillar cellulose, nanofibrillated cellulose (NFC), cellulose nanofibrils (CNF) [26,55]; a synonym of microcrystalline cellulose (MCC) is whiskers [53]. Some time ago, the Technical Association of the Pulp and Paper Industry proposed to standardize the terminology used (nomenclature and abbreviation). According to the recommendations (WI 3021), depending on the dimensions (width (w) and length/width ratio (L/w)), cellulose materials are divided into: cellulose nanocrystals (CNC, w = 3–10 nm, L/w > 5), cellulose nanofibrils (CNF, w = 5–30 nm, L/w > 50), cellulose microcrystals (CMC, w = 10–15 µm, L/w < 2), cellulose microfibrils (CMF, w = 10–100 nm, L/w > 50) [61]. The main cellulose types found in the literature and used in the production of composite materials are shown in Figure 2.

**Figure 2.** Overview of cellulose micro- and nanomaterials commonly used as fillers to prepare composite materials. Electron micrographs of (**a**) sisal fiber (scanning electron microscopy (SEM), reproduced from [62] under the terms of the Creative Commons CC BY license), (**b**) tunicate whiskers (transmission electron microscopy (TEM), reproduced from [63] with permission of American Chemical Society), (**c**) sugar beet CMF (TEM, reproduced from [64] with permission of Elsevier), (**d**) CMC, commercial (SEM, reproduced from [65] with permission of John Wiley & Sons, Inc), (**e**) wood CNF (TEM, reproduced from [66] with permission of American Chemical Society), (**f**) CNC from ramie fibers (TEM, reproduced from [67] under the terms of the Creative Commons CC BY license).

The size, type, and consequently the physical and chemical properties of the resulting cellulose materials are influenced by the source of origin, processing, and extraction method [44]. For instance, the use of mechanical action alone or its combination with chemical treatment of previously purified cellulose pulp/fibers (e.g., carboxymethylation or TEMPO-mediated oxidation) and/or enzymatic hydrolysis results in thin long flexible micro- (CMF) or nanofibrillar (CNF) structures with alternating crystalline and noncrystalline domains. In turn, acidic hydrolysis produces stiffer particles with a high degree of crystallinity (CMC and CNC), which are the result of the action of acid on both amorphous and crystalline domains. Thus, in the first case, the obtained cellulose micro- and nanofibrils retain the inherent semicrystallinity and high aspect ratio (L/w, over 50 for CMF and CNF) [39,46,49,55], while acid exposure reduces the number of defects in the structure and results in more highly crystalline materials with much lower L/w values (8 to 67 for CMC and CNC) [26,53,68].

Despite the existing terminology recommendations for cellulosic micro- and nanomaterials (see above), the use of terminology in the practice of current publications varies. Nevertheless, we have attempted to generalize the available data on the size of the various

cellulose-based materials used. The preparation methods and summarized descriptions and characteristics of the obtained micro- and nanocellulose materials found in literature are presented in Table 1 [26–28,39,44,46,69,70].

**Table 1.** Cellulose micro- and nanomaterials.

Abbreviations: CNC: cellulose nanocrystals (nanocrystalline cellulose); CMC: cellulose microcrystals (microcrystalline cellulose); CNF: cellulose nanofibers; CMF: cellulose microfibers.

The degree of crystallinity and DP for all obtained materials depends largely on the source of the cellulose as well as the processing technique. The found values are very scattered. For example, DPs for micro- and nanoobjects in the literature range from 100 to 15,000 [46,70], and the degrees of crystallinity vary from a few dozen to more than 90% [26,27,55,70]. For instance, the degree of crystallinity for BC and tunicin (cellulose from tunicate) is 80–100%, for cellulose from algae it is more than 70%, and cellulose from plants it is 40–60% [39,46].

The source of cellulose also has a significant impact on its mechanical properties. Young's tensile modulus for cellulose fibers varies from 5 to 200 GPa [39,46,71]. The highest values from this range are typical for tunicin fibers (from 110 GPa), while for fibers from other origins the elastic modulus does not exceed 115–130 GPa [39,46]. Elongation at break and tensile strength of cellulose fibers are in the ranges of 1–30% and 0.2–1.2 GPa, respectively [26,71]. The application of cellulose-fiber treatments that help to reduce the amorphous components in the chain packing thereby leads to a decrease in DP, an increase in crystallinity and, as a result, an increase in the mechanical properties of the resulting cellulose material compared to the original fibers [39,55]. The theoretically calculated Young's modulus (E) of an ideal cellulose crystal (along the axis of the cellulose chain) is 167.5 GPa [72]. According to the published data, the practically identified Young's modulus values for micro/nanocrystals range from 60 to 220 GPa [26,39,44,46,49,73]; for micro/nanofibers from 14 to 84 GPa [39,46,73]; and for a single tunicin microfibril, a value of about 150 GPa has been found [74]. The established tensile-strength data are in the range of 1–10 GPa for CMC and CNC [39,44,46,58]; about 2–6 GPa for cellulose nanofibers [66]; and about 4–8% elongation at break for wood-cellulose CNF has been reported [46]. The above data indicate the excellent mechanical properties of these micro- and nanoscale cellulose materials. The characteristics of crystalline cellulose (density 1.5–1.6 g/cm3 ) are close to—and in some cases significantly higher than—those of glass fibers used for composites (E about 70 GPa, density 2.6 g/cm3 ), Kevlar (E 60–125 GPa, density 1.45 g/cm3 ) and steel (E 200–220 GPa, density about 8 g/cm3 ) [61].

In addition to the above subgroups, cellulose materials such as amorphous nanocellulose (ANC) and cellulose nanoyarn (CNY) are also mentioned in the literature. ANC are obtained from regenerated cellulose by acid hydrolysis and ultrasonic treatment and are generally elliptical particles 50–200 nm wide with DP of 60–70 with an amorphous structure

and hence poor mechanical properties. CNYs are electrospun nanofibers; their width and DP range from 500–800 nm and 300–600 nm, respectively [26]. CNFs can be made into cellulose filaments by flow-focusing, wet-extrusion, or spinning processes, but their mechanical properties are also inferior to the more highly crystalline forms of micro- and nanocellulose materials [46,75]. Figure 3 shows these three forms of cellulosic materials.

**Figure 3.** SEM images of (**a**) ANC from CMC (reproduced from [76] under the terms of the Creative Commons CC BY license), (**b**) CNY from carboxymethyl cellulose sodium salt with polyethylene oxide at ratio 1:1 (reproduced from [77] with permission of John Wiley & Sons), (**c**) cellulose filament from carboxymethylated CNF (reproduced from [78] under the terms of the Creative Commons CC BY license).

#### **3. Aliphatic Polyesters**

Aliphatic polyesters (polyhydroxyalkanoates) are biocompatible and biodegradable materials. Currently, they are of interest for the fabrication of biomedical and environmentally friendly materials to replace petroleum-based plastics. Polylactide, polyglycolide, polyhydroxybutyrate, polycaprolactone, poly(butylene succinate) and some copolymers based on them are among the widely considered aliphatic polyesters [79–83]. The chemical structures of key aliphatic polyesters are illustrated in Figure 4, while their key characteristics are summarized in Table 2.

**Figure 4.** Structures of PLA, PGA, PLGA, P3HB, PHBV, PCL, and PBS.

#### *3.1. Poly(lactic acid)*

Poly(lactic acid) (PLA) or polylactide is a synthetic, thermoplastic, biocompatible, and biodegradable aliphatic polyester derived from renewable lactic acid [7,11,84]. PLA can be produced by lactic-acid condensation or by the ring-opening polymerization (ROP) of lactide (lactic-acid cyclic dimer) [5,7,11,14,44]. Polycondensation of lactic acid allows the synthesis of only a low-molecular-weight polymer due to the side reaction of hydrolysis preventing the production of high-molecular-weight polymer chains. In contrast to polycondensation, ROP provides PLA with high molecular weight but requires the use of catalysts [7]. A combination of these methods is commonly used to produce PLA on an industrial scale [3]. In this case, a low-molecular-weight polymer is synthesized from lactic acid (2-hydroxypropionic acid) by polycondensation, then the formed polymer is depolymerized to form lactide, which is further used to produce PLA of high molecular weights by ROP [3,7].

The presence of optical isomers of lactic acid and lactide (L-lactic acid, D-lactic acid, L,L-lactide or L-lactide, D,D-lactide or D-lactide, D,L-lactide or meso-lactide) leads to obtaining PLA of four types: isotactic and optically active poly-L-lactide (PLLA) and poly-Dlactide (PDLA), syndiotactic and atactic optically inactive poly-D,L-lactide (PDLLA) [44,84,85]. Equimolar racemic mixture of L- and D-enantiomers of lactide (rac-lactide) is also designated as D,L-lactide [7,85]. Chemical structures of stereoisomers of monomers and PLA can be found elsewhere [44].

The molecular structure (enantiopure) of PLA and heat treatment affect its crystallinity [3,84]. Optically inactive PDLLA is amorphous, whereas stereoregular PLLA and PDLA are capable of homocrystallization [84], forming α- (highest thermodynamic stability), β- or γ-crystalline forms depending on composition [3,6]. Blending PLLA and PDLA leads to their cocrystallization and the formation of a stereocomplex with a different crystal structure, characterized by an increase in melting temperature (*Tm*) by about 50 ◦C relative to the homopolymer PLLA or PDLA [6,101]. The parameters of the crystalline forms of PLA can be found in detail elsewhere [6]. Stereochemistry has a tremendous influence on the supramolecular structure of PLA. The presence of more than 10 mol% of the links in the polymer chain different from the basic optical form leads to a significant decrease in crystallinity [84,102]. Branching of the polymer chain also leads to a rather significant decrease in the crystallinity and glass-transition temperature (*Tg*) of PLA [102]. The crystallinity of PLA is also affected by the molecular weight of the polymer, various treatments, the introduction of nucleation agents, plasticizers into the matrix, and for final products, PLA crystallization can be initiated by temperature annealing [3,6,103].

Crystallinity, as well as the parameters listed above, determine the physical (thermal, rheological, barrier, etc.) and mechanical properties, as well as the degradability of PLA [3, 7,44,104,105]. The high degree of crystallinity of polylactide leads to excellent thermal and mechanical properties [44]. However, a high degree of crystallinity is not always necessary and is determined by the application of the final polymer product. For example, the rapid crystallization can complicate stretching the product by blow molding, can diminish the optical transparency of the product, such as a bottle, and can increase the degradation time of the polymer, which may limit its use for some biomedical applications. At the same time, the presence of thermal stability due to high crystallinity is very important for products formed by injection molding [44]. It has been reported that the increase in molecular weight and crystallinity is accompanied by an increase in viscosity and softening point, so that the behavior of PLA in the melt becomes similar to polystyrene [85]. The thermal stability of PLA is similar to poly(vinyl chloride) (PVC), but significantly inferior to PP, polyethylene (PE), polystyrene (PS), and PET [85].

The glass-transition temperature and melting point of PLA are important parameters for predicting the material properties [84]. Both are influenced by the molecular weight of the polymer. *Tg* and *Tm* increase sharply when the number-average molecular weight (*Mn*) increases to 80,000 and 100,000, respectively, and then remain almost unchanged [102]. As optical purity decreases with a constant molecular weight, a decrease in glass-transition

temperature is observed. Moreover, PDLA is characterized by lower *Tg* values than PLLA with the same molecular weight [3,102]. In turn, *Tm* is more influenced by the amorphous state of PLA than *Tg*, due to the lack or complete absence of a crystalline phase [102].

The commercially available type of PLA produced on a large scale is mainly PLLA [6,84], since about 90% of all lactic acid is produced from renewable sources by microorganisms as L-isomer [84]. Thus, a commercial polylactide is a semicrystalline polymer with a *Tg* of 55 to 65 ◦C and a *Tm* of 140 to 180 ◦C depending on the amount of the D-enantiomer impurity [94]. In comparison with petrochemical-based plastics, PLA has a slow crystallization rate, low impact strength, low thermal resistance, and low glass-transition temperature and fragility [11,94]. For instance, to consider the substitution of PET by PLA for packaging fabrication, the barrier properties of PLA need to be improved. Typically, aliphatic polyesters with molecular weights greater than 60,000 are used for packaging, agricultural, and biomedical applications [11].

#### *3.2. Poly(glycolic acid)*

Poly(glycolic acid) or polyglycolide (PGA) and its copolymer with PLA (PLGA) is among the most widely studied and used polymers [84,86]. PGA is a semicrystalline, biodegradable, biocompatible aliphatic polyester that differs from PLA in the absence of a methyl group in the monomer unit (glycolic acid residue) [86,93]. PGA can be obtained by polycondensation of glycolic acid (difficult to obtain high molecular weights), ROP of glycolide (more economical, but pure monomer is required), and solid-phase polycondensation of halogen acetates (low degree of polymerization) [86]. The synthesis conditions determine PGA molecular weight, crystallinity, *Tm* and *Tg*, and terminal groups. The growth of the molecular weight of PGA contributes to an increase in crystallinity, mechanical properties and a decrease in the biodegradation rate. Acceptable mechanical properties of PGA are achieved at molecular weights greater than 30,000 [86]. Due to the high degradation rate of PGA, its synthesis is more difficult and expensive than for PLA.

Mechanical, thermal, degradation properties and density of PGA are determined by molecular weights, dispersity, and crystallinity. PGA is characterized by high crystallinity. The most common crystallization degree is 45–55%, but 77% has also been reported. Due to the stabilized crystal cage, PGA has a high melting point (220–230 ◦C) and poor solubility (soluble only in highly fluorinated solvents such as hexafluoroisopropanol) [86]. The glasstransition temperature (*Tg*) of PGA is higher than the ambient temperature, but close to human body temperature (35–40 ◦C), which makes the material elastic when introduced into the human body (e.g., implantation) [86,93]. PGA is characterized by poor thermal stability because *Tm* is very high and close to the degradation temperature [90,92]. The lack of solubility in conventional organic solvents and the narrow processing window of PGA melt create a problem in obtaining products based on it [93]. At the same time, the supramolecular structure of PGA provides excellent mechanical properties [86,93]. For example, the elastic modulus (*E*) of PGA is higher than that of other synthetic biodegradable polymers (PLLA, PDLLA, poly-ε-caprolactone) and is 6–7 GPa [86,93]. The high density of PGA (1.50–1.71 g/cm3 ), due to the molecular-packing features, provides high gas-tight properties of the polymer, exceeding this parameter of polyethylene terephthalate (PET) by 100 times [86,92].

Currently, PGA is of great interest for renewable industry and biomedical applications due to its thermal properties, biocompatibility, biodegradation, nontoxicity, excellent mechanical characteristics, and low gas permeability. The obstacles of PGA application are overcome by making PGA-based copolymers and composites [86,93,106]. For example, by copolymerizing glycolide and various enantiomers of lactide and varying their ratios, the properties of the resulting PLGA (stiffness, crystallinity, melting point, and biodegradation rate) can be set. For example, PLGA demonstrates mechanical properties similar to those of human calcareous bone. In addition, PLGA is widely used as implants, micro- and nanoparticles for drug delivery [107,108].

### *3.3. Poly(hydroxybutyrate)*

Poly(3-hydroxybutyrate) (P3HB, PHB) is a thermoplastic, biodegradable, semicrystalline, linear microbial aliphatic polyester [5,8,79,87]. Biosynthesis in cells of natural/transgenic plants and bacterial fermentation, including the use of genetically modified microorganisms, are the main ways for the production of P3HB [8,79,87]. In particular, P3HB is produced by prokaryotic microorganisms from sugar-based media (agricultural industrial wastes, hydrolysates of some wood) and other carbon sources in the form of inclusion bodies, which serve as intracellular bacterial depots storing carbon and energy [5,79,87]. Microorganisms may accumulate up to 40–50% of P3HB from the dry-cell mass, and in the case of *Alcaligenes eutrophus* the accumulation may reach up to 96% of the dry cell mass [87]. Depending on the conditions and isolation forms, the resulting P3HB can have different characteristics (molecular weight, crystallinity, mechanical properties and ability to biodegrade) [8,79].

P3HB exhibits optical activity due to a chiral central carbon, and the main natural configuration is poly(D-3-hydroxybutyrate) [87]. The stereostructure and tacticity of P3HB can be specified by chemical synthesis, obtaining isotactic, syndiotactic, or atactic PHB [8,88,109]. The number of monomeric units in P3HB can vary in different range: (1) over than 10,000 for P3HB produced as cytosolic inclusions of bacteria, (2) 100–300 for P3HB from cell membranes and (3) up to 30 monomeric units for P3HB for other natural sources, including human tissues [8]. P3HB with a number of monomeric units greater than 1000 (molecular weight greater than 100,000) can be obtained chemically from β-butyrolactone [8,109].

The linear structure of the P3HB chain ensures its high crystallinity with the presence of an amorphous phase in addition to the crystalline phase [79]. The crystallinity of P3HB can vary in a wide range from 50 to 80%, and as with PLA, has a significant effect on the mechanical properties [79,87]. P3HB is generally characterized as a strong and stiff polymer with low thermal stability and low crystallization rate. Secondary crystallization of P3HB occurs at room temperature with the formation of amorphous lamellae, leading to polymer brittleness [87,110]. P3HB has piezoelectric properties and is also characterized by good resistance to acids, bases, and ultraviolet radiation [87]. In addition, P3HB has better barrier properties than PP, PE, PVC, and PET, and is characterized by some other properties similar to or superior to those of PP and PE [79]. In addition to these advantages, P3HB is a biocompatible and nontoxic polyester, which makes it a promising environmentally friendly alternative to petrochemical polymers, and also demonstrates suitability for various tissueengineering and other biomedical applications (scaffolds, surgical threads, drug-delivery systems, surgical mesh, etc.).

The low thermal stability is due to the close values of the melting and degradation temperatures of P3HB, which leads to a narrow heat-treatment window. Thus, thermal degradation of the polymer melt can occur during processing [110,111]. This problem can be partially solved by using a random copolymer of 3-hydroxybutyric and 3-hydroxyvaleric acids (poly(3-hydroxybutyrate-co-3-hydroxyvalerate) or poly(3-hydroxybutyric acid-co-3 hydroxyvaleric acid) abbreviated as PHBV or PHBHV. PHBV as well as P3HB homopolymer is characterized by high crystallinity, biocompatibility, biodegradability in different environments, good barrier properties, nontoxicity, UV stability, similarity to P3HB in solubility and chemical stability, hydrophobicity, and low impact resistance and fragility. Unlike P3HB, PHBV has a lower *Tm*, higher surface tension and flexibility [111]. Thus, PHBV appears to be technologically more attractive and of interest as materials for biomedical applications, agriculture, and packaging materials, and has been developed on an industrial scale [88]. PHBV copolymer can be produced by various microorganisms [111], including recombinant strains, in amounts up to 80% of dry-cell weight, and PHBV composition can vary over a wide range depending on substrate composition [111].

In addition to PHBV, lower melting-point values are observed for another copolymer, namely poly(3-hydroxybutyrate-co-4-hydroxybutyrate) (P(3HB-co-4HB)). Thus, an increase in the number of 4-hydroxybutyrate units (from 0 to 38 mol%) provides a significant decrease in *Tm* (from 176 to 54 ◦C), and with a further increase in the proportion of 4HB the melting temperature of copolymers practically does not change. P(3HB-co-4HB) is a

thermoplastic biodegradable aliphatic polyester produced by bacterial fermentation. The ratio of monomeric units is largely determined by the substrate used. Moreover, during biosynthesis the production of a mixture of copolymer compositions with a wide range of monomer compositions is observed, including the presence of P4HB, which significantly affects the characteristics of isolated polymers [112]. Regarding P3HB, P4HB is a relatively new material, which can also be obtained by polycondensation of 4-hydroxybutyric acid or ROP γ-butyrolactone. P4HB is also nontoxic, biocompatible, biodegradable, UV-resistant, demonstrates relatively good barrier properties, and exhibits optical activity, and therefore can be used for biomedical and packaging applications [79].


*3.4. Model Outputs*

Of the seven activities that were agreed upon by the steering committee, five were implemented by the design team by the end of the study period; two were not completed because they needed other resources to implement (i.e., waiting on information technology department to complete tasks). Surveys at follow up from 105 store workers (67% response rate) showed the activities noticed most often by workers were ones that were implemented in the breakroom: the new employee refrigerator and discounted bottles of water. Activities that were delivered in other areas of the store were implemented intermittently, and not noticed by many workers. Only six workers said they did not notice any of the implemented program activities. Similarly, utilization of the activities was higher for those implemented in the breakroom and for the activities that did not require much extra effort by the workers. During store worker interviews, workers either were excited about the new health activities and wanted to see more implemented or they had not heard of them. Those that had not heard of the activities indicated that direct communication from store management about new health opportunities may be more useful than printed materials placed throughout the store. Only one of

the five workers reported health behavior changes based on an implemented program activity; other workers said that they appreciated the effort, but that none of the implemented activities impacted their personal behaviors.


**6. Summary**

482 Solar Radiation

When sizing the battery bank, the ampere-hours efficiency (columbic efficiency) of new battery is considered to be 100%. During the charge/discharge cycle of the battery, the battery is charged by receiving an input voltage from the thermocouple system and the same number of voltage is delivered at a lower output voltage. The battery efficiency (battery's voltaic efficiency) is expressed as the ratio of average voltage output to the

The daily load requirements determine the necessary battery bank capacity while the system voltage determines the battery bank voltage and the number of cells to be connected in series. The product of Daily Load (DL) requirement and the number of no-sun days (N)

Total Useable Capacity = DL X N (Ah) The ampere-hours efficiency (columbic efficiency) of new battery is considered to be 100%. The daily load requirement determines the necessary battery bank capacity. The total

An assumption of a lighting programme in poultry house where power is needed for four out of every seven days in a week was made, and the inverse relationship between voltage and amperage was used to determine the average daily current requirement by multiplying current by a factor of 4/7 to yield a net value in Ah, the average daily current required to satisfy the load demand of 5.14Ah as calculated from given relations. An average of 4½ hrs of full sunshine hours per day round the years is taken for a non-critical system. The thermocouple array load capable of generating the required load demand is obtained by

Thermocouple load = average daily current requirement /4.5 (A). When a peak sunshine hour of 4.5 hours/day is required, the thermocouple array designed is capable of generating a measured 1.14A, capable of providing a glow continuously to satisfy the load demand of 5.14Ah. At increased number of sunshine hours above 4 ½ hours,

The conversion efficiency is defined as the ratio of electrical power output and the heat flux

Δ*T* corresponds to the temperature difference between the hot and the cold side of the TEG, *A* is the TEG area and *h* is the overall heat transfer coefficient given by (the ratio of total thermal conductivity (5.4 x 10 -4 W/m2/°C) of the materials of the thermoelectric generators and the thickness (0.015m) of the TEG. The electrical power output (P=174.06 W h). The

������ (38)

more current generation is possible whereby the battery could be recharged.

� = � � <sup>=</sup> �

**5.1.2 Battery size requirement and efficiency**

gives the total useable capacity (TUC) of the battery i.e.

useable capacity (TUC) of the battery is 22.84Ah

dividing the average daily current requirement by 4.5.

**5.1.4 The system conversion efficiency**

through the entire TEG surface.

**5.1.3 Daily load requirement**

average input voltage.

The conversion efficiency of the cell is 15%. This is comparable to other solar TEG system efficiency. The research work indicates the possibility of the utilization of asphalt bonded thermocouples to generate enough current for lighting programme in small scale agricultural undertaking such as poultry house illumination. The output voltage across the thermocouple generator can be increased to higher value enough to provide energy for other low thermal processes. Asphalt heat absorber will be a promising solar harvest cell when the surface is polished and made more sensitive to wider photon energy range (1.3- 1.5eV) for increased efficiency.

From the economical point of view, there exists a huge discrepancy between the costs of commercial thermoelectric generators compare with asphalts embedded TEG. The commercial TEG is by a factor of 10 more expensive than the asphalt TEG. The properties of asphalt TEGs are comparable with that of commercial TEG, even though the asphalt TEG used in this study has a smaller area than most commercial TEGs, therefore more asphalt TEGs per unit area can be mounted for increased overall performance at a cheaper price. Further research on antireflection coatings and stacking of different cells with band gaps covering the incident energy of the photons would be a good attempt at achieving higher efficiency.