Biocompatibility of a physiological pressure sensor

Biocompatibility of a physiological pressure sensor

Chao Yang a, Chunfeng Zhao a, Lester Wold b, Kenton R. Kaufman a,*a Orthopedic Biomechanics Laboratory, Mayo Clinic, 200 First Street, SW, Rochester, MN 55905, USA

b Laboratory Medicine and Pathology, Mayo Clinic, Rochester, MN 55905, USA

Received 18 July 2002; received in revised form 26 February 2003; accepted 11 March 2003

Biosensors and Bioelectronics 19 (2003) 51�/58

www.elsevier.com/locate/bios

Abstract

A newly developed fiber optic micropressure sensor was evaluated for biocompatibility using the International Organization for

Standardization (ISO) test standard 10993-6. The test material and an inert control (fused silica glass) were tested in New Zealand

white rabbits. Four test specimens were implanted in the paravertebral muscles on one side of the spine about 2�/5 cm from the mid-

line and parallel to the spinal column. Similarly, four control specimens were implanted on the opposite side. The implantation

periods were 1, 4, and 12 weeks to ensure a steady state biological tissue response. Four animals were tested at each time period.

Macroscopic and microscopic observations were performed to compare the biological reactions between the test and control

materials. There was an inflammatory reaction at 1 week which subsided at 4 weeks. There was fibrous tissue growth near the

implant that also decreased over time. Most importantly, there was no significant difference in the biological response between the

test and control materials. Therefore, we conclude that the pressure microsensor is biocompatible.

# 2003 Published by Elsevier Science B.V.

Keywords: Biocompatibility; Silica glass; Fiber optic pressure sensor; Intramuscular pressure; Tissue response

1. Introduction

Measurements of biological fluid pressure are of

interest to clinicians in cardiology, pulmonology, gas-

troenterology, urology, neurology and rehabilitation.

Commercially available pressure transducers are too

large for optimum comfort (Crenshaw et al., 1992).

Microsensor technology is now available to construct

transducers that are much smaller. A fiber optic micro-

sensor has been developed for measurement of physio-

logical pressures (Kaufman et al., 2002). The

measurement method is based on an extrinsic Fabry�/

Perot interferometric (EFPI) technique (Murphy et al.,

1991). EFPI technology is a distance measurement

technique based on the formation of a low-finesse

Fabry�/Perot cavity between the polished end of a fiber

and a reflective surface, which is the undersurface of a

diaphragm (Fig. 1). Light is passed through the fiber,

where a portion of the light is reflected off the fiber/air

interface (R1). The remaining light propagates through

the air gap between the fiber and the reflective surface

and is reflected back into the fiber (R2). R1 is the

reference reflection while R2 is the sensing reflection.

Fluid pressure causes displacement of the diaphragm.

These two light waves interfere constructively or

destructively based on the path light difference traversed

by the sensing reflection relative to the reference

reflection, and travel back through the single mode

fiber to the demodulation unit. The return signal is

analyzed with a miniature spectrometer.

The sensor consists of an optical fiber and a dia-

phragm. The optical fiber is made with inert silica glass.

The diaphragm is made from polyimide. These materials

have been carefully selected with biocompatibility issues

in mind. However, biocompatibility testing is a neces-

sary step since this device will be used for measurement

of in-vivo pressures (Black, 1999; ISO 1992�/1994).

Although the main part of this sensor is made of inert

silica glass, the construction of this sensor requires a

special adhesive to attach the diaphragm and optical

fiber together. Therefore, the objective of this study was

to test the in-vivo host response for this sensor.

* Corresponding author. Tel.: �/1-507-284-2262; fax: �/1-507-266-

2227.

E-mail address: [email protected] (K.R. Kaufman).

0956-5663/03/$ - see front matter # 2003 Published by Elsevier Science B.V.

doi:10.1016/S0956-5663(03)00131-3

mailto:[email protected]

2. Materials and methods

The International Organization for Standardization

(ISO) test 10993-6, ‘Biological Evaluation of Medical

Devices-Part 6: Tests for local effects after implanta-

tion’, was followed in this test protocol (ISO 1992�/

1994). The test sensor was composed of fused silica

glass, polyimide, and an aliphatic amine-cured mineral

filled epoxy adhesive. The control material was totally

composed of inert silica glass (both of these materials

were provided by Luna Innovations, Blacksburg, VA).

The test material and control were a flat-shaped

cylinder, 10 mm in length and 360 mm in diameter

(Fig. 2).

Twelve healthy adult New Zealand white rabbits were

tested. Their body weight ranged from 3.0�/4.0 kg

(average 3.2 kg). The animals were divided into three

implantation periods (1, 4, and 12 weeks), with four

animals tested at each time period. The Mayo Institu-

tional Animal Use and Care Committee (IACUC)

approved all animal procedures. NIH guidelines for

the care and use of laboratory animals (NIH Publication

#85-23 Rev. 1985) have been observed. The persons

who performed the surgical procedure were trained to

obey animal welfare and handling protocols. The rabbits

were anesthetized using Ketamine (30�/50 mg/kg) and

then placed in sternal recumbency. The hair over the

dorsum, from mid-thorax to caudal lumbar region was

shaved. The shaved region then received a surgical

preparation with Betadine scrub and then Betadine

solution. The tip of the spinous process was palpated

to identify the lumbar spine (the space between the

lumbar spinous process is larger than that between the

thoracic spinous process). The sterile test sensors and

controls were percutaneously inserted at eight sites

along the lumbar paraspinal muscles (four on each

side). The implant was injected through an 18 gauge, 51

Fig. 1. Block diagram of the low-frequency, high resolution, fiber optic pressure measurement system. Light is conveyed via a fiber optic cable to the

sensor. Fluid pressure causes a diaphragm to deflect. The light is reflected back at the fiber optic interface and the sensor diaphragm. The returned

light is analyzed with a miniature spectrometer to determine the relative displacement of the diaphragm. The graphs represent the light spectrum that

is returned from the sensor before it is processed to determine the gap measurement.

Fig. 2. Configuration of test and control samples used in biocompat-

ibility tests.

C. Yang et al. / Biosensors and Bioelectronics 19 (2003) 51�/5852

mm I.V. needle (Terumo Corporation Tokyo, Japan) at

a 458 into the muscle, with the needle tip about 1.5 cm

beneath the skin. The implant was advanced using the

needle until the implant exited the tip of the I.V. needle.The implantation sites were recorded in reference to the

nearest vertebra. The injection sites were then covered

with sterile tissue. After the implantations were com-

plete, the animal was given a dose of Buprenorphine

(0.02�/0.05 mg/kg by I.V. means) for pain relief and

taken off the inhalation anesthesia. The animals recov-

ered from surgery on a warming pad and were placed in

their cages when fully recovered from the anesthesia.The rabbits were observed daily during the first post-

operative week and weekly thereafter. The observation

included body temperature, posture, activity, appetite,

and sensitivity to touch. If needed, analgesics were

administrated.

At 1, 4, and 12 weeks after implantation the animals

was anesthetized with Ketamine or Xylazine and

euthanized with Pentobarbital (�/100 mg/kg by I.V.).The rabbits were placed in a prone position. The hairs

over the dorsum, from mid-thorax to caudal lumbar

region, were shaved again. Any skin changes near the

implantation site were recorded. A vertical incision was

made along the dorsum midline. After exposure of the

paraspinal muscle, the nature and extent of any tissue

reaction was observed with the help of a low magnifica-

tion lens. The whole longissimus lumborum was dis-sected and wrapped with physiological saline-soaked

gauze, and fixed with 10% formalin solution for at least

48 h. The implants were found by sectioning the tissue

until the implant was identified. Approximately 1 cm3 of

paraspinal muscle at each implantation site was then

isolated and sent for histological preparation.

All specimens were formalin fixed and processed

routinely resulting in the tissue being embedded inparaffin. The slides were cut from the paraffin blocks

perpendicular to the long axis of the implant. The

section orientation in relation to the implant dimensions

and implant orientation was recorded.

The host response to the implant was determined by

measurement of the distance from the implant/tissue

interface to unaffected areas, i.e. normal tissue and

normal vascularity. The biological response parametersassessed and recorded included the: (a) extent of

fibrosis/fibrous capsule and inflammation; (b) degenera-

tion determined by changes in tissue morphology; (c)

number and distribution of inflammatory cells types,

namely macrophages, polymorphonuclear leukocytes,

lymphocytes, plasma cells, giant cells, as a function of

distance from the material/tissue interface; (d) presence

of necrosis determined by nuclear debris and/or capil-lary wall breakdown; (e) and other parameters such as

material debris, fatty infiltration and presence of

granulomata. All analyses were done with a digital

image analysis system (Microscope: Axioplan2; camera:

Axiocam; software: KS400. Carlzeiss Inc. Oberkochen

Germany).

A two-sample t-test was used to test for significant

differences between the two groups at each time period.The dependent variables were thickness of reaction

capsule and density of inflammatory cells. Statistical

significance was set at P�/0.05.

3. Result

3.1. General and macroscopic observations

All 12 rabbits recovered well after the surgery. All

grossly showed swelling around the insertion area on the

1st post-operative day. The swelling gradually subsided

by the 2nd day post-operatively. From the 2nd post-

operative day to the day of sacrificing, no abnormal

conditions, i.e. temperature, eating, and activity, were

observed. During post-mortem dissection, no gross

abnormalities were identified in the skin or the hostmuscle. No local swelling or fluid accumulation was

found around the implant sites and no suppurative

changes were identified. No infection occurred. For all

rabbits, the insertion hole could still be found on the

muscle fascia for all implants after 1, 4 and 12 weeks

(Fig. 3).

3.2. Cell responses and histology

Although there were no significant regimes of necro-

sis, granuloma formation, or soft tissue calcification

around any of the implants, apoptosis could occasion-

ally be seen in stromal cells in the 1 week test group.

However, the number of such observations was small

and did not consistently occur. For the 1 week group,

the inflammatory reaction was mild (Fig. 4). Monocytes,macrophages and eosinophils were the main inflamma-

tory cell types present for both test and control groups.

The capsule wall was composed of three layers. The

inner layer was mostly composed of macrophages and

monocytes. Eosinophils were located in the middle

layer. The outer layer was mainly composed of fibro-

blasts. There was no significant difference in inflamma-

tory cell composition or capsule thickness between thetest and control groups (Figs. 5�/7). At 1 week the mean

capsule thickness of test and control groups were 90.0

and 94.7 mm, respectively (t�/0.755, P �/0.05). The

mean density of macrophages and monocytes were 5479

and 5610/mm2, for the test and control groups, respec-

tively (t�/0.218, P �/0.05). The mean density of eosi-

nophils was 1655 and 1619/mm2, respectively (t�/0.201,

P �/0.05). The density of fibroblasts was 1621 and 1938/mm2, respectively (t�/1.409, P �/0.05).

For the week-4 group, the capsule thickness decreased

sharply. Unlike the week-1 sample, the capsule consisted

C. Yang et al. / Biosensors and Bioelectronics 19 (2003) 51�/58 53

Fig. 3. Comparison of insertion sites for test and control materials at 1, 4 and 12 weeks after implantation. The small box in each figure indicates the

region shown in Fig. 4.

Fig. 4. Close-up view of test and control sites at 1, 4 and 12 weeks after implantation. There was no significant difference in thickness of the fibrous

capsule or type and density of cellular response.


of two clearly distinctive layers (Fig. 4). The inner layer

was an inflammatory cell layer. The cell composition

was mainly macrophages and round cells. Lymphocytes

could be occasionally noticed at this time. Eosinophils

decreased. The outer layer was a thin fibrous layer

formed by fibroblasts and collagen fibers. The mean

capsule thickness for the test group (47.2 mm) and

control group (47.3 mm) was not significantly different

(t�/0.51, P �/0.05) (Fig. 5). The macrophages and otherinflammatory cell decreased at 4 weeks, so we counted

all the inflammatory cells together. The mean density of

inflammatory cells were 3243 and 3330/mm2, for the test

and control groups, respectively (t�/0.089, P �/0.05)

(Fig. 6). The mean density of fibroblasts was 4211 and

4655/mm2, respectively (t�/0.161, P �/0.05) (Fig. 7).

From 4 to 12 week, the thickness of the capsule

decreased slightly. The mean capsule thickness for thetest group was 30.0 mm, and the mean thickness of the

control group was 29.0 mm (Fig. 5). There was no

significantly difference between the groups (t�/0.249,

P �/0.05). The capsule was composed of 2�/3 layers of

cells, mainly fibroblasts. Macrophages were present but

decreased. The mean density of inflammatory cells was

2625 and 2500/mm2, for the test and control groups,

respectively (t�/0.238, P �/0.05) (Fig. 6). The meandensity of fibroblasts was 3970 and 3879/mm2, respec-

tively (t�/0.078, P �/0.05) (Fig. 7).

4. Discussion

This biocompatibility test focused on the direct

interactions between the substance of the implant andthe chemical and biological aspects of the host environ-

ment. Histologic and morphologic in-vivo examinations

are well-accepted methods to assess biocompatibility

giving reliable unequivocal results. Standards that

govern these types of tests have been detailed by various

federal and international agencies (ASTM, 1997; ISO

1992�/1994). The accepted international standard for

biocompatibility testing is ISO 10993, which is thestandard used in this study (ISO 1992�/1994).

This study followed the ISO 10993-6 standard with

respect to the type of animal used, study period, implant

site and surgical procedure (ISO 1992�/1994). However,

the dimension of the implant was 10 mm by 360 mm,

which was slightly thinner than the ISO 10993 recom-

mendation. We chose this sized implant as it reflects the

actual full-sized implant that would be used in futureclinic practice. In addition, the implant size allowed us

to insert it through a needle probably resulting in a

decrease in the acute inflammatory response as com-

pared to an open surgery procedure. The sensor will

probably always be inserted with a needle in clinical

practice. We chose silica glass as the negative control

since this material is similar to our test material. In

addition, silica is a well-accepted inert material used as acontrol in many previous studies (Hakkinen et al., 1988;

Ito et al., 1987b; Kobayashi et al., 1997a, 1999, 1997b;

Matsuda et al., 1987; Seitz et al., 1982).

Fig. 5. Capsule thickness surrounding test and control implants at 1, 4

and 12 weeks after implantation. There was no significant difference in

capsule thickness between test and control at each time point studied.

Fig. 6. Cellular density of inflammatory cells in the reaction zone.

There was no significant difference in cellular response surrounding the

test and control implants at each time period.

Fig. 7. Cellular density of fibroblasts in the reaction zone throughout

the test period. There was no significant difference in cellular density

between the test and control implants at each time period.


Biocompatibility testing can be done using sensitive

in-vitro cell culture techniques. These tests can serve as

precursors for more involved, more costly, and more

time-consuming animal biocompatibility tests. We chosean in-vivo animal test because the morphology and

function of cells in-vivo depend to a degree on the

chemistry of the pericellular environment and the nature

of the substrate supporting and surrounding the cells.

Lack of these environmental factors may cause cells to

become inactive or to revert, at least functionally (Black,

1999). Tissue culture experiments are considered by

many to be too sensitive if used alone since failure inthese tests would eliminate many useful materials that

are presently in use. Effects on division, growth, and

function of cells in culture can be easily and quantita-

tively estimated. But such tests should not be considered

‘fail safe’. A material that displays some toxic signs in

culture may yet prove acceptable because of the capacity

of the whole animal to buffer local effects. In the present

study, an in-vivo test was adopted to circumvent someof these problems (Wilson et al., 1981).

The in-vivo biocompatibility of a material can be

evaluated by analyzing the cell population present,

measuring the mediator and metabolite cells excreted,

and analyzing the morphologic characteristics of the

tissue and the capsule thickness around the implant.

Different methods or scoring systems have been used to

evaluate the host response (Wood et al., 1970; Boswaldet al., 1999; Ito et al., 1987a; Wilson et al., 1981;

Schmalz and Schmalz, 1981; Haggerty and Lusted,

1989; Mathur et al., 1997; Allen et al., 2001; Woodward

and Salthouse, 1986). No one system seems to be well

accepted. In our study, the average capsule thickness

decreased sharply from the 1st week to the 4th week for

both the sensor and control, indicating that the tissue

went from an acute, insertion procedure related inflam-matory phase to a relatively stable phase. At week 1,

macrophages and a small amount of monocytes were at

the inner layer. Their number was much higher at 1

week compared to 4 and 12 weeks. Woodward and

Salthouse observed similar changes (Woodward and

Salthouse, 1986). They pointed out that activated

macrophages will be observed at the surface of all

implant biomaterials, but their numbers will be propor-tional to the toxicity of the implant material. With

smooth-walled, nontoxic materials, the macrophage

population will be largely replaced by fibroblasts and

later by a fibrous capsule (Woodward and Salthouse,

1986). This is the type of cellular cascade we observed. It

was very important to cut the samples perpendicular to

the long axis of the implant. This way we could measure

the real thickness of the capsule. Yet, some samples werenot sectioned ideally. In these cases, we measured the

thinnest part of the capsule, which represented the real

thickness of the capsule. The area chosen for cell

counting was also important. We chose a rectangular

area with the long axis perpendicular to the capsule.

This area included all the layers of the reaction zones

from the inner part to outer part. This avoided

inappropriate sampling of a particular area and omis-sion of another area within the reaction zone. This

method ensured that every cell type was counted in the

evaluation.

Measuring the capsule membrane thickness around

the implant is a basic and important tool for estimating

biocompatibility. Investigators have used different mea-

suring techniques (Ellies et al., 1988; Christel and

Meunier, 1989; Benghuzzi, 1996; Ryhanen et al.,1998), but the results show similar outcomes. In our

study, the capsule thickness dropped sharply from 1 to 4

weeks, and slowly to 12 weeks. Even at 12 weeks, the

capsule was still about 30 mm for the test group and 29

mm for control group. This supports the contention that

after 4 weeks, no new chemicals were released from the

implant, and the tissue underwent a stable inflammation

process. As Woodward pointed out, the interfacebetween the host and the foreign body is permanent,

and the cells and collagen participating in it turn over

extremely slowly after the interface response is estab-

lished (Woodward and Salthouse, 1986). This was

confirmed by Ryhanen et al. (1998), where after 26

weeks there was still a thin layer of fibrous capsule

around the implant.

In this study, although the difference in number ofinflammatory cells between the two groups was not

significant, the cell numbers of the control group

appeared to be higher than the test group at 1 and 4

weeks. For the 1 week group, some apoptosis was

noticed in the test group, which is probably due to

chemical release. Since the amount of apoptosis was

very small, it could not be used for statistical analysis.

More importantly, there was no significant difference inthe inflammatory process of the two groups at either 4

or 12 weeks.

The cells at the interface were mainly macrophages,

eosinophils, and fibroblasts. Other monocytes, giant

cells and lymphophils were only occasionally present.

This cellular distribution has been reported previously

(Wood et al., 1970; Wilson et al., 1981; Christel and

Meunier, 1989; Mathur et al., 1997; Ryhanen et al.,1998). Polymorphonuclear granulocytes migrate into the

damaged tissues, where their phagocytic activity re-

moves tissue debris, deleterious foreign material, and

pathogens. Like some other studies, most of the poly-

morphonuclear cells present in our case were eosinophils

but not neutrophils. This indicates that no infection was

present. Although eosinophils were present at the

implant-tissue interface early, the eosinophils tended todisappear by the 12 weeks observations.

Monocytes migrated to the interface where they

assumed the role of macrophages. Macrophages play a

very important role in acute inflammation and probably


in final biocompatibility. They can release mediators,

which can in turn activate other cells. Macrophages are

found to affect fibroblast and lymphocyte activity. The

amounts of macrophages present at the different timeperiods did not differ between the test and control

material. Very few giant cells were identified, suggesting

that the amount of ‘foreign body debris’ was small.

Fibroblasts migrated to the implant-tissue interface in

the early phase of healing and gathered around the

implants. Other granulation tissue was slowly replaced

by fibroblast proliferation and collagen deposition. In

optimal situations, an inert biomaterial forms a thin,relatively avascular and acellular fibrous scar capsule at

the tissue�/plant interface. This kind of capsule was

similar to what we saw in our 12 weeks group.

The basic biomaterial�/cell interaction is still poorly

understood. If a material is clearly toxic, it induces a

strong local response. On the other hand, if there are

toxic elements in the material which dissolve slowly,

they might give symptoms many years after the implantis inserted. Thus, the results of a single study are only

suggestive in nature. However, our test material was

predominantly composed of materials that are known to

be inert. Therefore, we do not anticipate that a longer

duration study will change the cell response documented

in this study.

5. Summary

Through our study, we conclude that this microsensoris biocompatible. The microsensor has a bio-reaction

which was histopathologically similar to the inert silica

glass in term of inflammatory migration and fibro-

capsule formation. Based on these results, this device

can be used for in-vivo human applications.

Acknowledgements

This study was supported by NIH grant R01H03476.We thank James E. Tarara for his help with the semi-

quantitative image analysis. Appreciation is also ex-

pressed to Tom Wavering and Luna Innovations for

their collaboration on the design, development and

manufacturing of this pressure microsensor.

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Biocompatibility of a physiological pressure sensor